Abstract:
An electrical method and apparatus for stimulating cardiac cells causing contraction to force hemodynamic output during fibrillation, hemodynamically compromising tachycardia, or asystole. Forcing fields are applied to the heart to give cardiac output on an emergency basis until the arrhythmia ceases or other intervention takes place. The device is used as a stand alone external or internal device, or as a backup to an ICD, atrial defibrillator, or an anti-tachycardia pacemaker. The method and apparatus maintain some cardiac output and not necessarily defibrillation.

Description:
This is a Continuation-in-Part of application Ser. No. 08/754,712, filed Dec. 6, 1996, U.S. Pat. No. 5,978,708, which in turn is a continuation of Ser. No. 08/543,001, filed Oct. 13, 1995, abandoned, which in turn is a continuation of Ser. No. 08/251,349, filed May 31, 1994, abandoned. 
    
    
     BACKGROUND OF THE INVENTION 
     1. Field of the Invention 
     The invention relates to the field of therapies for cardiac arrhythmias, and more particularly, to a method and an apparatus for forcing cardiac output by delivering a pulsatile electrical field to the heart during fibrillation of a hemodynamically compromising tachycardia. 
     2. Background Information 
     Approximately 400,000 Americans succumb to ventricular fibrillation each year. It is known that ventricular fibrillation, a usually fatal heart arrhythmia, can only be terminated by the application of an electrical shock delivered to the heart. This is through electrodes applied to the chest connected to an external defibrillator or electrodes implanted within the body connected to an implantable cardioverter defibrillator (ICD). Paramedics cannot usually respond rapidly enough with their external defibrillators to restore life. New methods of dealing with this problem include less expensive external defibrillators (and thus more readily available) and smaller implantable defibrillators. Since the first use on humans of a completely implantable cardiac defibrillator in 1980, research has focused on making them continually smaller and more efficient by reducing the defibrillation threshold energy level. The goal has been to reduce the size of the implantable device so that it could be implanted prophylactically, I.E., in high risk patients before an episode of ventricular fibrillation. 
     An ICD includes an electrical pulse generator and an arrhythmia detection circuit coupled to the heart by a series of two or more electrodes implanted in the body. A battery power supply, and one or more charge storage capacitors are used for delivering defibrillation shocks in the form of electrical current pulses to the heart. These devices try to restore normal rhythm from the fibrillation. While it works well at restoring normal function, the ICD is large in size and not practical for a truly prophylactic device. A small device capable of maintaining minimal cardiac output, in high risk patients, prior to admission into an emergency room is needed. 
     In addition, external defribillators are limited in their performance. The typical paramedic defibrillation may be delayed by 10 minutes. At this time defibrillation may be irrelevant since the rhythm is often advanced to asystole. In asystole, there is little or no electrical activity and certainly no cardiac pumping. 
     There is a need for a new method and apparatus for dealing with ventricular fibrillation. The defibrillation approach does not work satisfactorily. External devices are too slow in arrival and implantable defibrillators are excessively large (and expensive) for prophylactic use. 
     SUMMARY OF THE INVENTION 
     The invention provides an electrial method of stimulating cardiac cells causing contraction to force hemodynamic output during fibrillation, hemodynamically compromising tachycardia, or asystole. Forcing fields are applied to the heart to give cardiac output on an emergency basis until the arrhythmia ceases or other intervention takes place. The device is usable as a stand alone external or internal device or as a backup to an ICD, atrial defibrillator, or an anti-tachycardia pacemaker. 
     The goal of the invention is maintaining some cardiac output and not necessarily defibrillation. The method is referred to as Electrical Cardiac Output Forcing and the apparatus is the Electrical Cardiac Output Forcer (ECOF). 
     In the implantable embodiment, a forcing field is generated by applying approximately 50 volts to the heart at a rate of approximately 100-180 beats per minute. These fields are applied after detection of an arrhythmia and maintained for up to several hours. This will generate a cardiac output which is a fraction of the normal maximum capacity. The heart has a 4 or 5 times reserve capacity so a fraction of normal pumping activity will maintain life and consciousness. 
     The implantable embodiment is implanted in high risk patients who have never had fibrillation. If they do fibrillate, the ECOF device forces a cardiac output for a period of up to several hours, thus giving the patient enough time to get to a hospital. That patient would then be a candidate for an implantable cardioverter defibrillator (ICD). The ECOF differs from the ICD in that it is primarily intended for a single usage in forcing cardiac output over a period of hours, while the ICD is designed to furnish hundreds of defibrillation shocks over a period of years. 
     Insofar as is known, no prior attempts have been made at forcing pulses during any type of fibrillation. Some workers in the field have experimented for research purposes with local pacing during fibrillation. For example, Kirchhof did local pacing during atrial fibrillation in dog hearts (Circulation 1993; 88: 736-749). He used 0.5 mm diameter electrodes and pacing stimuli. As expected, small areas around the heart were captured but no pumping action was expected or detected. Similar results have been obtained in the ventricle by Ken Knight (Journal of the American College of Cardiology 1994; 283A). 
     Various researchers have tried multiple pulse defibrillation without success in reducing the energy thresholds, for example, Schuder (Cardiovascular Research; 1970, 4, 497-501), Kugelberg (Medical &amp; Biological Engineering; 1968, 6, 167-169 and Acta Chirurgica Scandinavia; 1967, 372), Resnekov (Cardiovascular Research; 1968, 2, 261-264), and Geddes (Journal of Applied Physiology; 1973, 34, 8-11). 
     More recently, Sweeney (U.S. Pat. No. 4,996,984) has experimented with multiple (primarily dual) shocks of timing calculated from the fibrillation rate. None of these approaches has been able to significantly reduce voltages from conventional defibrillation shocks. Importantly, none of these approaches anticipated the idea that the individual pulses might force cardiac output or could sustain life indefinitely. 
     Some have considered the use of smaller pulses, before the shock, to reduce the energy required for a defibrillation shock (Kroll, European Application No. 540266), but never anticipated eliminating the defibrillation shock itself or anticipated that the pulses themselves could maintain cardiac output. Some have suggested using higher voltage pulses to terminate ventricular tachycardias, but no suggestion was made of an application with fibrillation or of obtaining cardiac output (Kroll WO 93/19809) and Duffin (WO 93/06886). 
     The benefits of this invention will become clear from the following description by reference to the drawings. 
    
    
     BRIEF DESCRIPTION OF THE DRAWINGS 
     FIG. 1 is a block diagram illustrating a system constructed in acordance with the principles of the present invention. 
     FIG. 2 a  shows the connection of an implantable embodiment of the device to the heart in an epicardial patch configuration. 
     FIG. 2 b  shows the connection of an implantable embodiment of the device to the heart using an endocardial lead system and the device housing as an electrode. 
     FIG. 3 a  shows the connection of an external embodiment of the invention. 
     FIG. 3 b  shows a representative cardiac output detection configuration. 
     FIG. 4 is a diagram showing a representative pulsatile electrical signal. 
     FIG. 5 is a flowchart illustrating one embodiment of the method of the invention. 
     FIG. 6 is a diagram showing the expected effect of a 50 V pulse on the heart during diastole. 
     FIG. 7 is a diagram showing the expected effect of a 50 V pulse on the heart during systole. 
     FIG. 8 is a diagram showing the expected effect of a 50 V pulse on the heart during fibrillation. 
     FIGS. 9 a  and  9   b  show various waveforms useful for the electrical cardiac output forcing method and apparatus. 
     FIG. 10 shows the device used as a backup to an atrial defibrillator. 
    
    
     DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS 
     The present invention will now be described more fully hereinafter with reference to the accompanying drawings, in which preferred embodiments of the invention are shown. This invention may, however, be embodied in many different forms and should not be construed as limited to the embodiments set forth herein. Rather, applicants provide these embodiments so that this disclosure will be thorough and complete, and will convey the scope of the invention to those skilled in the art. 
     FIG. 1 is a block diagram illustrating a system  10  constructed in accordance with the principles of the present invention. The device circuitry is connected to the heart  40  via a series of leads; output lead  32 , presure sense lead  34 , and ECG sense lead  36 . The electronic circuit includes a conventional ECG amplifier  30  for amplifying cardiac signals. The amplified cardiac signals are analyzed by a conventional arrhythmia detector  20  which determines if an arrhythmia is present. The arrhythmia detector  20  may be one of several types well known to those skilled in the art and is preferably able to distinguish between different types of arrhythmias. For example; fibrillation, tachycardia or asystole. The circuit also contains an optional pressure sensing section  28  which amplifies and conditions a signal from an optional pressure sensor from within the heart or artery. The output of the pressure sense circuit  28  is fed to a cardiac output detection circuit  18  which analyzes the data and determines an estimate of the cardiac output. Data from the arrhythmia detector circuit  20  and the cardiac output detection circuit  18  is fed to the microprocessor  16 . The microprocessor  16  determines if Electrical Cardiac Output Forcing (ECOF) is appropriate. If forcing is indicated, the microprocessor  16  prompts the output controll  22  to charge a capacitor within the output circuit  26  via the capacitor charger  24 . The output control  22  directs the output circuitry  26  to deliver the pulses to the heart  40  via the output leads  32 . The microporcesor  16  may communicate with external sources via a telemetry circuit  14  within the device  10 . The power for the device  10  is supplied by an internal battery  12 . 
     FIG. 2 a  is a diagram showing the connection of an implantable embodiment of the device  130  to the heart  40  in an epicardial patch configuration. In this thoracotomy configuration, current passes through an output lead pair  32  to electrode patches  42  which direct the current through the heart  40 . There is an optional pressure sense lead  34  which passes the signal from an optional pressure transducer  46  which lies in the heart  40 . The ECG is monitored by sense electrodes  44  and passed to the device  130  by a lead  36 . The area of the electrodes  42  is at least 0.5 cm 2 . The size of the electrode is greater than that of a pacing lead and no more than that of a defibrillation electrode or between approximately 0.5 cm 2  and 20 cm 2  each. 
     FIG. 2 b  shows a non-thoracotomy system embodiment of the invention. In this system, the current passes from a coil electrode  52  in the heart  40  to the housing of the device  140 . An endocardial lead  50  combines the ECG sensing lead and the pulse output lead. The ECG is monitored by sense electrodes  44  in the heart  40  and passes through the endocardial lead  50 . There is an optional pressure transducer  46  in the heart  40  which passes a signal to the device  140  via optional lead  34 . 
     FIG. 3 a  shows an external embodiment of the invention. External patch electrodes  54  are placed on the chest to deliver current to the heart  40  through output lead  32 . The ECG is monitored by surface electrodes  56  and passed to the device  150  by a lead  36 . Alternately, the ECG could be monitored by the external patch electrodes  54 . An optional pressure sensor  46  passes a pressure signal via an optional pressure sense lead  34 . This embodiment could be used as a substitute (due to its small size) for an external defibrillator and keep a patient alive until arrival at a hospital. Also, the system could precede the external defibrillator by generating output in patients in asystole until blood flow and rhythm are restored. 
     FIG. 3 b  shows another means of detecting cardiac output which may be useful to the external embodiment, in particular, of this invention. FIG. 3 b  illustrates use of a relatively high frequency (such as about 10-50 kHz) impedance measurement across the chest area of the patient, with either 2 or more electrodes, for example patch electrodes  54   b.  Surface electrodes  56   b  sense and deliver the signal, which may be rectified to highlight the cardiac mechanical frequencies (1-10 Hz) at amplifier  57 ′. In one embodiment, this may be a 30 kHz selective amplifier. Processing circuitry may further include rectifier means  57 ″ and filter  57 ″′, which could include a 1-20 Hz filter or similar filter means. 
     A series of forcing pulses  60  are shown in FIG.  4 . The pulses are approximately 50 V in amplitude with a spacing of approximately 500 ms. The 50 V and the 500 ms pulse spacing are chosen as illustrative for an implantable embodiment. The forcing pulse interval is chosen to maximize cardiac output within the limits of device circuitry and the response of the heart muscle. An interval of 500 ms corresponds to a heart rate of 120 beats per minute. This will produce a greater output than a typical resting rate of 60 beats per minute. However, a rate of 240 beats per minute would produce a lower output due to mechanical limitations of the heart. Thus a practical range is 60 to 200 beats per minute is appropriate. The pulses could also be timed to coincide with the natural pumping of the atria, thus improving overall cardiac output. 
     The higher voltage, the higher the forcing fields, and therefore a greater number of heart cells contracting producing greater cardiac output. However, the higher voltage produces greater patient discomfort and extraneous muscle twitching. 
     Implantable batteries are also limited to a certain power output and energy storage. If an output pulse is 50 V and the electrode impedance is 50 Ω, the power during the pulse is P=V 2 /R=50V*50V/50Ω=50W. If the pulse has a duration of 2 ms then the energy per pulse is 0.1J. If two pulses are delivered every second, the charger must be capable of delivering 0.2J per second which is 200 mW. This is well within the limits of an implantable battery. An implantable battery can typically deliver 5 W of power. However, 200 V pulses at 3 per second would require 4.8 W which is near the limit of the battery and charging circuitry. A typical implantable battery energy capacity is 10,000 J. Delivering forcing pulses at a rate of 4.8 W would deplete the battery in only 35 minutes (10,000J/4.8W=2083 seconds). Thirty five minutes may not be enough time to transport the patient to a hospital. Therefore 200 V represents the highest practical voltage for continuous operation in an implantable embodiment, although voltages of up to 350 V could be used for short periods and adjusted down when hemodynamic output is verified. A practical lower limit is about 10 A. During normal sinus rhythm, 10 V delivered through the patches would pace. However, during fibrillation the 10 V could not pace and only cells very near the electrodes would be captured. This would be insufficient for forcing cardiac output. 
     These calculations also suggest other differences between an implantable ECOF and an ICD. With a battery storing 10,000 J and ECOF pulse having 0.1 J, this ECOF would be capable of delivering 100,000 pulses. An ICD can only deliver 200-400 shocks of about 30 J. The ECOF is also very different from an implantable pacemaker which typically delivers 150,000,000 pacing pulses (5 years at 60 BPM) each of about 0.00005 J. 
     For an external ECOF the calculations are similar, but scaled up. The typical ECOF pulse would have a voltage of 100 V with a range of 25-500 V. With electrode impedances of 50 Ω the power during the pulse is P=V 2 /R=100V*100V/50Ω=200 W with a range of 12.5-5,000 W. If the pulse has a duration of 2-5 ms, then the energy per pulse is 0.02-25 J. This is much less than the American Heart Association recommended output of 360 J for an external defibrillator. 
     This is also different from an external transthoracic pacemaker. These devices are rated by current and typically have an output range of 30-140 mA. Most patients are paced by pulses of 40-70 mA of current. An example of a modern external thoracic pacemaker is given by Freeman in application WO 93/01861. Assuming an electrical impedance of 50Ω and the ECOF voltage range of 25-500 V, then the ECOF current range would be 500 mA 59 10 A. Since electrode impedance increases with lower voltage, the 25 V ECOF pulse would probably see an impedance of 100Ω thereby giving a lower current of 250 mA. 
     However, it is now recognized that use of external defibrillation has a homogeneous current advantage over ICDs, due to the relatively poor electrical field coverage of the ICDs. Accordingly, it is believed that the external ECOF-type of pulses described herein have an added advantage over pulses delivered with implantable systems. A further advantage of external delivery of pulses exists with regard to the potential for skeletal muscle and diaphragm contraction to assist with cardiac output. This type of contraction augments the cardiac contraction normally attributed to both implanted and external pulse delivery. These advantages accumulate to present an external ECOF ratio which allows for use of voltage levels between a range of about 20-2000 volts in a combined external ECOF and defibrillation device. However, without the traditional higher voltage defibrillation requirement, it is likely that an external ECOF-type device may only require a delivery capability of between about 20-1000 volts, with possible initial capture pulses that may be higher or lower than that upper range. For example, if an external ECOF-type ratio is no longer considered to be between 5-10, and is only assigned a value of 2, then the 10 volt internal minimum becomes a 20 volt external minimum, and the internal typical delivery range of 20-200 volts becomes a typical external delivery range of 40-400 volts. In similar manner, a representative value for a maximum internal delivery of an ECOF-type pulse might be 350 V, with a comparable external value being only 700 V using this ratio. When considered in the context of being a non-invasive therapy, the external ECOF-type of application is quite advantageous and energy efficient. This is particularly so in view of the unexpected ratio described above which is improved over the previous known ratios of AED/ICD voltage ratio values of between about 4-10. 
     FIG. 5 is a flowchart illustrating the method of the invention, which is provided for purposes of illustration only. One skilled in the art will recognize from the discussion that alternative embodiments may be employed without departing from the principles of the invention. The flow diagram shown in FIG. 5 represents a method of automatically treating a heart which is in fibrillation, tachycardia, or asystole and thereby pumping inefficiently or not at all. Electrodes are attached  69  and diagnoses the presence of an arrhythmia  70 . A series of cardiac output forcing electric pulses  72  is automatically delivered. It should be understood that the therapy  72  may be delivered for any output within the ranges stated herein which is appropriate for the particular cardiac dysrhythmia. After delivery of forcing pulses (for example, such as either ten pulses at a rate of 60-200 BPM or a series of pulses for a certain time, e.g., 20 seconds) in the first block  72 , the status of the heart is determined  74 . If an anomaly such as arrhythmia or low cardiac output is still present and/or there exists low pressure within the heart, more forcing pulses are delivered  78 . Such delivery may include, for example, perhaps 100 pulses of 200 V for 20 seconds, although these values may be varied according to the recipient&#39;s condition. If the heart is pumping at a safe level, the therapy ceases and exits  76 . Note that this means that the ECOF successfully defibrillated the patient&#39;s heart even though this may not be the only goal of this therapy. This could be tested in patients who were scheduled to receive an ICD, in a hospital setting. Those patients who are defibrillated by ECOF pulse therapy could then receive the ECOF instead of the larger ICD. After the therapy  78  has been delivered, the pressure and ECG is again monitored  74 . If the therapy  78  is successful, it ceases and exits  76 . If the therapy  78  is unsuccessful in producing a safe level of pumping efficiency, the method proceeds to a third delivery mode  80  which may comprise a continuous cardiac assist mode, a defibrillate mode, or some combination of the above. Indeed, in one embodiment of the method depicted in FIG. 5, the first step of delivery of pulses is an ECOF-type of pulse sequence, the second step of delivery of pulses is a high voltage defibrillating shock type of pulse, and the third step of delivery of pulses is an assist-type of ECOF-like pulse sequence. In another embodiment of this method, it is possible to deliver a shock level pulse one or several times, then deliver a forcing or ECOF type of pulse sequence one or more times, and then again deliver a shock level pulse one or several times. The inclusion of the forcing pulses between shocks replenishes the blood in the vessels during the therapy. 
     The therapy may only be stopped by an external command, for example, a telemetry signal or a magnet which is applied to the chest activating a magnetic reed switch  82  which terminates the therapy and exits  76 , or some other appropriate termination means is activated. To minimize patient discomfort and maximize battery life, the forcing voltage could be adjusted down when sufficient pressure signals or adequate flow measured by other means were detected, for example, the pressure sense transducer could be replaced by an oxygen detector or a doppler flow measuring device. The pulse rate could also be adjusted to maximize output. 
     FIG. 6 is a diagram showing the effect of a 50 V forcing pulse on the heart  40  during electrical diastole (cells at rest). The current is passed through the heart  40  by the electrodes  42 . Approximately 60% of cardiac cells  90  would be captured by a 50 V pulse if the cells were in diastole. The captured cells  90  mostly lie in the direct path between the electrodes  42  and near the electrodes  42  where the field strengths are highest. Of course over a time period of about 100 ms these directly captured cells then propagate an activation wavefront to stimulate the rest of the heart. This so called far-field pacing is irrelevant here as the hearts, of interest, are in fibrillation and not in diastole. 
     FIG. 7 is a diagram showing the effect of a 50 V forcing pulse on the heart during electrical systole (cells already stimulated). The current is passed through the heart  40  by the electrodes  42 . Approximately 20% of cardiac cells  100  would be captured by a 50 V pulse if the cells were in systole. The captured cells  100  are nearest each electrode  42  where the field strengths are highest. Capture in systolic cells means that their activation potential is extended. This capture requires significantly higher fields (10 V/cm) than those required for diastolic cell capture (1 V/cm). 
     FIG. 8 is a diagram showing the effect of a 50 V forcing pulse on the heart during fibrillation. During fibrillation there are always cells in systole and diastole simultaneously. But, the vast majority are in systole. This diagram assumes 50% of the cells are in diastole which applies only after several capturing pulses. The current is passed through the heart  40  by the electrodes  42 . 100% of the cells  110  nearest the electrodes  42  would be captured due to the high field strength. As shown in FIG. 7, even systolic cells are captured by high field strengths. 50% of the cells  112  in the direct path betwen the electrodes  42  would be captured if it is assumed that 50% of all cells are in diastole. If roughly 60% of cardiac cells are captured by a 50 V pulse when the cells are in diastole, and 20% are captured when in systole, and if 50% are in systole and 50% in diastole, 40% would be captured during fibrillation. This calculation is shown in the following table. The last two columns give the mechanical action resulting and the contribution to forcing a cardiac output. 
     Considering the cardiac cells that are originally in diastole, (rows A &amp; B) in the table below, the A row represents the diastolic cells that are not captured by the forcing pulse. If 50% of the heart&#39;s cells are in diastole and 40% of those are not captured that is 20% of the total cells. These cells will, however, shortly contract on their own (from previous wavefronts or new ones) providing a positive gain in mechanical action and therefore cardiac output. The B row corresponds to the diastolic cells that are captured. If 60% of the diastolic cells (50% of total) contract due to the forcing field this is 30% of the total heart cells. These cells provide the biggest gain in mechanical action and cardiac output. Next considering the activity of the systolic cells (rows C &amp; D), if 50% of the heart&#39;s cells are in systole and 80% of those are not captured (row C), that is 40% of the heart&#39;s cells. These cells soon relax and negate a portion of the cardiac output. The systolic cells that are captured (row D) are 10% of the heart&#39;s cells (20% of 50%). These cells will hold their contraction and be neutral to cardiac output. The net result is a gain in contraction which forces cardiac output. 
     
       
         
               
               
               
               
               
               
               
             
           
               
                   
               
               
                   
                 Percentage 
                 Status of 
                 Percentage 
                   
                   
                 Forcing 
               
               
                 Original 
                 of the 
                 the 
                 of the 
                 Percentage 
                   
                 Cardiac 
               
               
                 status of 
                 cardiac 
                 cardiac 
                 original 
                 of the total 
                 Mechanical 
                 Output 
               
               
                 the cells 
                 cells 
                 cells 
                 status 
                 cells 
                 Action 
                 Effect 
               
               
                   
               
             
             
               
                 (A) 
                 50% 
                 Diastolic 
                 40% to 50% 
                 20% 
                 Will start 
                 Positive 
               
               
                 Diastolic 
                   
                 non- 
                   
                   
                 to contract 
                 (+) 
               
               
                   
                   
                 captured 
                   
                   
                 on own 
               
               
                 (B) 
                   
                 Diastolic 
                 60% of 50% 
                 30% 
                 contract 
                 positive 
               
               
                 Diastolic 
                   
                 captured 
                   
                   
                   
                 (++) 
               
               
                 (C) 
                 50% 
                 Systolic 
                 80% of 50% 
                 40% 
                 Will start 
                 Negative 
               
               
                 Systolic 
                   
                 non- 
                   
                   
                 to relax on 
                 (−) 
               
               
                   
                   
                 captured 
                   
                   
                 own 
               
               
                 (D) 
                   
                 Systolic 
                 20% of 50% 
                 10% 
                 hold 
                 neutral 
               
               
                 Systolic 
                   
                 captured 
                   
                   
                   
                 (0) 
               
               
                 Total 
                 100% 
                   
                 100% 
                 100% 
                 More 
                 Positive 
               
               
                   
                   
                   
                   
                   
                 contraction 
                 (++) 
               
               
                   
               
             
          
         
       
     
     The net result over a 200 ms mechanical response is given in the next table. The major contribution is in row (B) from the captured diastolic cells contracting. 
     
       
         
               
               
               
               
             
           
               
                   
               
               
                   
                 Status of the 
                   
                 Description of 
               
               
                 Row 
                 Cardiac Cells 
                 Change in Output 
                 Activity  
               
               
                   
               
             
             
               
                 A 
                 Diastolic non- 
                 +5% 
                 Positive. Some 
               
               
                   
                 captured 
                   
                 cells will begin to 
               
               
                   
                   
                   
                 contract on their 
               
               
                   
                   
                   
                 own. 
               
               
                 B 
                 Diastolic captured 
                 +30% 
                 Positive. Cells 
               
               
                   
                   
                   
                 contract due to 
               
               
                   
                   
                   
                 forcing field. 
               
               
                 C 
                 Systolic non- 
                 −5% 
                 Negative. Some 
               
               
                   
                 captured 
                   
                 cells will begin to 
               
               
                   
                   
                   
                 relax on their own. 
               
               
                 D 
                 Systolic captured 
                 0% 
                 Neutral. Cells 
               
               
                   
                   
                   
                 hold contraction 
               
               
                   
                   
                   
                 due to forcing 
               
               
                   
                   
                   
                 field. 
               
               
                 Net Gain 
                   
                 +30% 
                 A net gain in 
               
               
                   
                   
                   
                 cardiac output due 
               
               
                   
                   
                   
                 to forcing fields.  
               
               
                   
               
             
          
         
       
     
     The 30% net pumping action should be sufficient to maintain survival and consciousness, because the heart has a 4-5 times reserve capacity. 
     FIG. 9 depicts examples of waveforms designed to minimize the twitching of the chest muscles which can be very uncomfortable to the patient. In FIG. 9 a  is seen a low harmonic pulse waveform  120  which has a very gradual “foot”  122  and a gradual peak  124 . Such a pulse has less high frequency energy components and thus is less likely to stimulate the skeletal muscle. 
     FIG. 9 b  shows a technique of going to the opposite extreme. Here, each compount forcing pulse  126  is actually composed of 50 very short spikes  128  each of which is 20 μs in width with a 20 μs spacing. The heart will tend to average out these thin pulses and “see” a 2 ms wide forcing pulse. The skeletal muscle, however, is not efficiently stimulated by these extremely narrow pulses. The skeletal muscle will not average out this signal either. This approach could help minimize skeletal muscle twitching and discomfort. 
     An alternative system would be to charge the capacitor to 300 V for the first pulse to capture many cells therefore putting those cells into diastole after a delay of 100-200 ms. At this point the voltage could be lowered to 100 V and delivered every 100 ms. A 3 watt DC-DC converter with a 67% efficiency could provide 100 ms interval forcing pulses assuming a 50 Ω resistance and 1 ms pulse (0.2J). This rate is too fast for forcing cardiac output due to mechanical limitations, but is very effective for electrical capture. After sufficient capture, the rate of forcing pulses could be slowed down to 100-170 beats per minute for optimum cardiac output. 
     The Electrical Cardiac Output Forcing device (ECOF) could also be used to help patients with atrial fibrillation. As an alternative embodiment to the ventricular placement of FIG. 2 b,  the electrode coil  52  and sensing electrodes  44  could be placed in the atrium. The device could then function to force atrial output. Even though atrial fibrillation is not instantly fatal like ventricular fibrillation is, clots can build up in the atria which can eventually lead to strokes. Cardiac output forcing of the atria on a daily basis may limit this problem. It is also possible that after a number of forcing pulses the atria would return to a normal rhythm. There is however, no urgency as is the case with ventricular fibrillation. 
     A second use of this invention for atrial defibrillation is shown in FIG.  10 . As before in FIG. 2 b,  the ECOF  160  is shown connected to the heart  40  via endocardial lead  50 . Again forcing coil electrode  52  and sensing electrodes  44  are in the right ventricle. In addition a large atrial coil electrode  130  and atrial sensing electrodes  132  are in the right atrium. These would be used for conventional atrial defibrillation. One of the big concerns with atrial defibrillation is that in a few cases, an atrial defibrillation shock causes ventricular fibrillation. If this happens, the patient dies within minutes. With the ECOF approach, for the left ventricle, one could maintain output in the patient for several hours and thus have enough time for transport to a hospital or external defibrillation. Thus the ECOF approach in the ventricle could provide a safety backup to atrial defibrillation. 
     Many cardiac patients have no known risk of ventricular fibrillation, but suffer regularly from ventricular tachycardia. Accordingly, these people can be treated with anti-tachycardia pacing (ATP). Unfortunately, occasionally ATP will cause a ventricular fibrillation. Then a large defibrillation shock must be applied. Thus it is not considered safe to implant a pure ATP device and these patients instead receive a full size ICD. The ECOF approach also serves as a safety backup and thus allow the implantation of true ATP devices. The system is depicted in FIG. 2 b , although the pressure sensor  46  would typically not be needed. 
     Low energy cardioverters can also be used to treat ventricular tachycardias. These devices are also not considered safe as stand alone devices due to the fact that they may not terminate the rhythm or that they may cause fibrillation. The ECOF method also could be used as a safety backup thus allowing the implantation of cardioverters without defibrillation capabilities. Such a system is shown in FIG. 2 b.    
     It should be understood that various alternatives to the embodiments of the invention described herein may be employed in practicing the invention. For example, while most of the discussion is in the context of an implantable device, the concepts of the invention are also applicable to external delivery systems. The use of ECOF-type of low voltage forcing pulses allows the choice of optimum therapy for different patient needs while only delivering the minimum voltage necessary to a patient in order to achieve the desired outcome. It is intended that the following claims define the scope of the invention and that structures and methods within the scope of these claims and their equivalents be covered thereby.