CA1106921A - Cardioverting apparatus - Google Patents

Abstract

ABSTRACT OF THE DISCLOSURE

A two-stage apparatus for cardioverting a malfunctioning heart conserves power by a first-stage detector for discrimina-ting between a normal cardiac state and a state requiring cardioversion; a powering circuit for continually powering the first-stage detector; a seoond-stage detector for further discriminating between a normal cardiac state and a state requiring cardioversion, normally in a stand-by condition and going to an active full power sensing condition upon initia-tion by the first-stage detector circuitry for enabling the second-stage detector and for thereby bringing the second-stage detector from its stand-by condition to its full power active sensing condition; a control device for powering the second-stage detector only upon said first-stage detector sensing a cardiac state requiring cardioversion; a device for delivering cardioverting energy to the heart; and circuitry for actuating the delivery means upon the second-stage detector sensing a cardiac state requiring cardioversion.

Description

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This is a division of Patent Application 262,273 filed September 29, 1976.
Ventricular fibrillation (VF) is a lethal cardiac arrhythmia for which the only known efficacious treatment is electrical countershock. A victim of VF outside of the hospital setting has little chance of survival since treatment must take place within a few minutes after the onset of the episode.
Fortunately, new techniques and devices are being devised to help deal with this life threatening condition.
Among these are computer techniques which aid in the identifica-tion of high risk VF patients, anti-arrhythmic drugs which can be prophylactiaally administered to these patients, programs for widespread cardio-pulmonary resuscitation training and implantable devices which can automatically detect VF and deliver cardioverting countersocks.
Many of the known techniques, such as defibrillation in a hospital setting, or defibrillation by a paramedic as part of a resuscitation program, rely upon the human detection of VF. This detection has typicall~ been accomplished by a trained operator interpreting an ECG from an oscilloscope tracing. However, there are situations where such an approach to reversing VF is impossible or impractical. There is ac-cordingly a great need for an electronic device able to ac-curately detect VF or other life threatening arrhythmias from an input ECG where such a traditional approach is unfeasible.
For example, an external defibrillator could be built with an interlock to its discharge switch so that a shock can be delivered only after the presence of VF has been confirmed by a detector receiving an ECG signal from the paddles. Such 2~

a defibrillator could safely be used by even an untrained operator.
With regard to the automatic implantable defibril-lator, techniques have been developed which are generally acceptable for detecting VF and discriminating between life threatening arrhythmias and other cardiac malfunctions. ~et there is considerable room for improvement with regard to detecting and discriminating VF from other non-fatal arrhythmias.
Accordingly, another use for such a detector as noted above would be in the totally implantable automatic defibrillator.
Previous approaches to VF detection for implantable devices have had certain drawbacks. Fundamental questions, particularly important to an automatic implantable defibrillator, relate to potential failure modes, the risks to a patient should the device reach one of these failure modes, and speci-fically to whether failures should occur in a passive or an active manner. Obviously, failures must be minimized, but they still must be considered. In this regard, it is believed preferable that potential sensing failures lead to inherent passivity of a defibrillating device.
In many known VF detectors and automatic implantable defibrillators, the primary detection schemes would result in active mode failures unless other lock-out circuitry is provided. Examples are R-wave sensors, pressure sensors, and elastomeric contraction sensors.
There is accordingly a great need for a VF detector which is accurate in its detection of VF or other life threaten-ing arrhythmias, so that failure modes may be passive.
There is disclosed herein an accurate simple, VF
detector which at least partly mitigates at least some of Z~L

the drawbacks of known VF detectors.
The present detector may be embodied in a system for measuring the electrical activity of the heart which can reliably discriminate between hemodynamically efficient and inefficient arrhythmias, being particularly sensitive to ven-tricular fibrillation. Though presented as a part of an auto-matic implantable defibrillator, it should be appreciated that the present detector is not limited to this specific application. For example, certain other arrhythmias, or tachy-arrhythmias can easily be identified by utilizing the present teachings.
Customarily, the term electrocardiogram (ECG) implies the use of electrodes on the body surface to obtain electrical signal indicative of heart activity. The term electrogram, on the other hand, generally refers to measurements made at the surface of the heart. As used herein, "ECG" is defined broadly, and refers to any measurement of the electrical activity of the heart, notwithstanding the source or technique of the measurement.
With the present detector, VF may be detected with a degree of accuracy never before possible, and hence inherent passive failure modes can be afforded. The detector may operate independently of the concepts of QRS detection and heart rate calculations to maximize accuracy. As is known, these concepts are particularly difficult to define during ventricular fibril~
lation. Furthermore, high-amplitude P and T-waves can in-accurately be sensed as R-waves, leading to false VF diagnosis.
The VF detector described hereinafter has simple circuitry to minimize component count and therefore the possibility of electronic component failure. And, the circuitry of the z~

hereinafter described VF detector is easily adaptable to low power operation.
Reference is also made hereinafter to the principle of probability density function. Briefly, the probability density function defines the fraction of time on the average that a given signal spends between two amplitude limits. It has been noted that the probability densit~ function of an ECG changes markedly between ventricular fibrillation and normal cardiac rhythm. A probability density need not be represented by the entire function, but rather, can be sampled at discrete values of amplitude. As employed herein, the entire function and the sampled form of the function a_e used interchangeably. As will become apparent from the following description, VF may be detected by monitoring a sampled pro-bability density. The probability density function can be monitored at any number of levels, but in a simple arrangement monitoring is accomplished at one level, near zero, which can be defined as the ECG baseline. In this instance, the ECG is filtered, providing a first derivative of the ECG, and in this manner moving any secondary probability density function peaks toward the desired zero.
Howevex, a detector for accurately discriminating between a normal cardiac state and a cardiac state requiring cardioversion may require a relatively large input power to perform its function.
It is accordingly an object of the present invention to provide apparatus for cardioverting a malfunctioning heart which enables the heart to be constantly monitored for the occurance of a cardiac state requiring cardioversion but without the continuous consumption of a relatively large amount of power.

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According to the present invention, there is provided a two-stage apparatus for cardioverting a malfunctioning heart in such a manner that power is conserved, the apparatus com-prising first-stage detector means for discriminating between a normal cardiac state and a state requiring cardioversion;
power/ing means for continually powering the first-stage de-tector means; second-stage detector means for further dis-criminating between a normal cardiac state and a state re-quiring cardioversion, normally in a stand-by condition and going to an active full power sensing condition upon initiation by the first-stage detector means; means for enabling the second-stage detector means and for thereby bringing the second-stage detector means from its stand-by condition to its full power active sensing condition; control means for powering the second-stage detector only upon the first-stage detector means sensing a cardiac state requiring cardioversion; delivery means for delivering cardioverting energy to the heart; and means for actuating the delivery means upon the second-stage detector means sensing a cardiac state requiring cardioversion.
The second-stage detector may be in the form of an impedance sensor which measures impedance between cardiac electrodes. It has been found that the impedance due to cardiac contractions is related to stroke volume. The impedance sensor, however, requires a relatively large input power to perform its sensing function. By means of the present invention, the impedance sensor may remain idle for the greater majority of time.
Embodiments of the invention will now be more par-ticularly described with reference to the accompanying drawings, given by way of example, in which:

Figure la is a tracing of a s~uare wave given for exemplary purposes;
Figure lb is a plot of the probability density function of the wave illustrated in Figure la;
Figure 2a is a typical catheter sensed ECG trace;
Figure 2b is a plot of the pro~ability density function of the ECG trace illustrated in Figure 2a;
Figure 3a is an ECG trace representing ventricular fibrillation;
Figure 3b is a curve representing the probability density function of the ECG trace illustrated in Figure 3a;
Figure 4 is a block diagram of a probability density function detector;
Figure 5 is a detailed circuit schematic of the detector illustrated in Figure 4;
Figure 6a is a curve of an exemplary input ECG sig-nal to the detector circuit of Figures 4 and 5, showing both normal cardiac rhythm and fibrillation;
Figures 6b through 6e are curves representing sig-nals at select locations in the circuit illustrated in Figures 4 and 5 based upon the ECG input illustrated in Figure 6a;
Figure 7 is a block diagram of a circuit for develop-ing probability density function traces for the input of an oscilloscope;
Figures 8a through 8d are curves illustrating an ideal example of filtering an ECG trace to move the probability density function to zero;
Figure 9a is a curve similar to that illustrated in Figure 2a, but representing the ECG trace after filtering;
Figure 9b is a probability density function similar ~$36~32~

to that shown in Figure 2b, but illustrating the function of the filtered ECG of Figure 9a;
Figure 10 is a block diagram of a phase lock loop second-stage detector;
Figures lla through llf represent signals at select locations in the circuit illustrated in Figure 10;
Figure 12 is a block diagram of a second-stage im-pedance sensor for detecting ventricular fibrillation; and Figures 13a through 13f are traces for explaining the operation of the impedance detector illustrated in Figure 12.
The probability density function cardiac arrhythmia detector will first be described. However, before embarking upon a detailed explanation of the circuit, there follows a brief discussion of the theory of probability density.
The detector system described hereinafter, is based upon a series of measurements on the ECG. The measurements are known in the literature as the probability density function, denoted as KX(X). If X(t) is a function of time, then KX(X) can be interpreted as a function that defines the fraction of time on the average that X(t) spends between two limits.
For example, the area under KX(X) between X=Xl and X=X2 is the fraction of time that X(t) spends between the limits X
and X2. Looking at the simplified example illustrated in Figure la, it can be seen that X(t) is always either at the levels X=B or X=A, and that the waveform spends half of its time at each one of these limits. The probability density function for this example is illustrated in Figure lb, wherein the continuous function of time X(t) has been mapped into a function of the amplitude-time distribution of X(t).

The present inventors have recognized that the pro-bability density function of an ECG changed markedly between normal cardiac rhythm and ventricular fibrillation. In this regard, the attention of the reader is directed to Figure 2a which illustrates a typical ECG trace, Figure 2b which shows the probability density function of the ECG illustrated in Figure 2a, Figure 3a which illustrate~s an exemplary ECG
trace representing ventricular fibrillation, and Figure 3b which is the probability density function of the trace illus-trated in Figure 3a. It will be noted that when comparingnormal cardiac rhythm with ventricular fibrillation, the great-est changes occur in the respective ECG traces at X=0, or at the baseline of the ECG signal. This is markedly reflected in the probability density functions as can be seen when com-paring Figures 2b and 3b.
In a most simplified arrangement the probability density is sampled at one value of x, namely X=0 or at the baseline of a filtered ECG. As will be later explained when reference is made to Figures 8 and 9, the filter in its most basic form provides the derivative of the ECG. Physiologically then, sampling the probability density of the filtered ECG
at X=0 corresponds to detecting the presence of relative isoelectric segments in the ECG. These isoelectric segments disappear during severe tachyarrhythmias such as fibrillation.
It should be noted that many other sampling levels are available for X other than zero, as will be explained below, and hence the number or level of sampling points are not in any way intended to be limited.
The present probability density function detector shown in block form in Figure 4, and in detailed schematic 2~

form in Figure 5. A representative set of waveforms is illus-trated in Figure 6.
The detector circuit is shown generally at 10, having a first stage ECG section 12, followed by a window stage 14, integrator stage 16 and threshold detector stage 18. The input to the detector 10 is by way of terminal 20 which leads directly to an ECG preamplifier 22. The~ output from pre-amplifier 22 is fed to a gain control circuit 24, and then in parallel to a first filter 26 and combined peak-to-peak detector and second filter 28.
The ECG section 12 has a bandpass filter character-istic. Most important in this bandpass characteristic is the highpass section which is designed to reject low frequency ECG components such as ST segments and to provide an approxi-mation of the first derivative. The automatic gain control circuit 24 is provided to normalize the probability density function over a known and fixed range of amplitude. To facili-tate understanding of the simplified block diagram of Figure 4, the respective transfer characteristics for the four discrete sections are provided immediately beneath each section.
With particular reference now to Figure 5, it can be seen that amplifier 42 serves as the main gain block, with capacitors 44, 46 and 48, and resistors 50, 52 and 54 serving as the bandpass elements. Gain control is provided by N-type junction field effect transistor 56 which shunts part of the ECG signal to ground through capacitor 58. This partial shunting results in a voltage divider effect with resistor 50. A typical endocardial electrogram which w~uld appear at terminal 20 and the corresponding output of the ECG section 12 which would appear at terminal 30 are illustrated in Figures g _ z~

6a and 6b, respectively. It should be apparent that the filters in the ECG section 12 concentrate the cardiac signal to a significant degree along the time axis.
After initial amplification and filtering in ECG
section 12, the signal at terminal 30 passes through window section 14 which comprises a window comparator 32. Window comparator 32 is designed to provide a digital signal at its output terminal 34, the sense of which depends upon whether the input to the comparator 32 lies inside or outside a band centered about a given window level introduced at terminal 60. In the simplified embodiment of the present invention, the window level at terminal 60 is chosen at the ECG base-line. The band can be seen between "+a" and "-a" in Figure 6b, and the resultant digital signal developed by window com-parator 32 and appearing at terminal 34 can be seen in Figure 6c. It will be noted that the digital output of comparator 32 goes to a fixed level whenever the filtered signal leaves the designated band. The sizes of resistors 62 and 63 set the band width "a". It can also be seen in Figure 6c that upon the onset of fibrillation, very little time is spent by the filtered ECG signal inside the designated band, corres-ponding to the lower value of the probability density function at X=0 as shown in Figures lb and 3b.
The digital signal appearing at terminal 34 is then integrated by integrator 36 with respect to a bias level, and produces a signal at output terminal 38 such as that illus-trated in Figure 6d. As can be seen, this output signal takes the form of a ramp when fibrillation begins. This output signal at termianl 38 in turn becomes an input to the threshold detector, or comparator 40. Detector 40 then switches when ~-f~ Z~

the ramp signal at terminal 38 reaches a given threshold level.
Hysteresis is provided in the threshold detector stage 40 for a latching function so that the ramp must fall past level Vt to VL (shown on the transfer characteristic beneath de-tector 40) for fibrillation detection to cease. This is in-dicated in Figure 6 during the period of inactive fibriIla-tion shown in Figure 6a wherein the trace of Figure 6d falls beneath the upper switching threshold of detector 40. Still, the output of detector 40 is high, resulting from the noted hysteresis characteristic.
It should be noted that the above-described simpli-fied detector configuration provides inherent passive failure mode behaviour and remains inactive if no ECG is applied.
Also, the inventive detector is independent of heart rate definition and its inherent ambiguity during VF. Accordingly, the probability density function VF detector overcomes major disadvantages common in known VF detectors.
From the previous discussion, it should be apparent that the probability density function provides another tool for viewing the original time-amplitude function. All of the discrete characteristics of the original signal are re-tained, but are displayed in a different format. Thus informa-tion of general diagnostic significanceis inherent in the presentation and in some instances can be more readily seen or measured automatically. The attention of the reader is therefore directed to Figure 7, which illustrates a circuit in greatly amplified block form which can be used to provide complete displays of probability density functions. These traces of probabilitydensity provide a great deal of informa-tion in the detection and study of tachyarrythmias.

In Figure 7, an input signal is introduced at inputterminal 62, to be then passed through an automatic gain control circuit 64. In this manner, input signals of different ampli-tude can be handled by the overall circuit. On the probability density display, signal amplitude appears on the abscissa, and therefore, AGC will normalize the width of the display.
A digital storage element 66 follows AGC 64, and serves to provide a repetitive sou~ce of the input signal.
The storage element 66 stores approximately two-seconds of ECG data in a digital memory, and continually repeats this data. In this manner, the same data is repeatedly provided to a window comparator 68.
The window comparator 68 provides a logical "1"
whenever its input signal lies within a narrow band centered around a band center "X". By then passing the output of the window comparator 68 through a simple low-pass filter 70, a voltage is developed, proportional to the average time that the input stays within the designated band. This signal is fed to the "vertical" input of oscilloscope 72. This is pre-cisely analogous to the definition of the distribution asdefined above. Sweeping the band center "X" slowly through the range of the input signal by means of a wave generator 74, provides a continuous display on the oscilloscope screen.
The band center is coupled into the "horizontal" input of the oscilloscope 72.
The respective traces ofFigures 2, 3 and 9 were developed from the circuit illustrated in Figure 7. The trace of Figure 2a, as noted above, represents an electrogram recorded from an intracardiac catheter. The corresponding density function is shown in Figure 2b. Several regions have been ~p~z~

identified on the respective traces, representing the same cardiac events but on the two different display patterns.
For example, t~e region "B" of Eigure 2a is the most negative peak of the R-wave. The signal spends very little time at this value, thus the corresponding peak of the probability density curve of Figure 2b is small. The peak at "A" is repre-sentative of the ST segment, and as is readily seen, the ECG
signal spends more time on the average at this level than at region "B". Accordingly, the peak is higher at "A" in Figure 2b. The ECG dwells longest at the baseline identified at "C" in Figure 2a, and the zero peak is the largest in Figure 2b.
It should by now be evident that the absence of a peak at zero in the probability density function may be utilized as being characteristic of abnormal cardiac rhythym.
By taking the derivative of the original ECG input signal, the function of the filter in the arrangement described above, the zero peak is considerably emphasized. Following up the example of Figure 1, reference should be made to Figure 8.
In Figure 8a, there is illustrated a square wave alternating between "+A" and "-A". The probability density function of this square wave is given in Figure 8b and is similar to that shown in Eigure lb. Since the square wave spends no time at X=0, the probability density function has no peak at X=0.
Figure 8c represents an impulse train which is developed by taking the derivative of the square wave illustrated in Figure 8a. The distribution function (probability density function) of the impulse train, unlike that illustrated in Figure 8b, is a unit impulse at zero, as shown in Figure 8d. Thus, the effect of taking the derivative of the original square wave 2:~

input and then evaluating the probability density function of the derivative is to shift peaks to X=0.
The same principle as explained in the ideal case of Figure 8 is applicable to the filtered ECG shown in Figure 9a. It will be recalled that the trace of Figure 9a represents the curve of Figure 2a after filtering. As can be seen, the peak "A" corresponding to the ST segment which appears in Figure 2b has been eliminated from the probability density function wave of Figure 9b. Furthermore, the Figure 9b zero peak is considerbly larger than that of Figure 2b. Thus, the filter improves the detection accuracy by enhancing the zero peak of the probability density function and thereby emphasizing the measure of the differences between VF and normal cardiac rhythm.
As mentioned previously, the distribution need not only be sampled at zero as in the embodiment of the VF detector described above. If two sampling points, say Xl and X0 are defined as shown in Figures 3b and 9b, more discrimination resolution becomes available by taking a ratio. As illustrated, approximate measurement would show the value of the pro-bability density function at these two points on the waveform for the two examples to be:
For Normal Rhythm Kx(Xl) = .012 x 0) Kx(Xl) .012 m Kx(X ~ 2.5 C = .0048 m For Ventricular Fibrillation Kx(Xl) = 08 KX(Xo) = .11 KX(Xl) m KX(Xo) .72 It can readily be seen that this measure yields over two orders of magnitude difference between normal cardiac rhythm and fibrillation. Sensing a value of C near 1.0 corresponds to the detection of a severe arrhythmia.
In view of the high degree of reliability necessary for the successful application of an implantable automatic defibrillator, it may become desirable to improve the accuracy of the detection system even relative to that described im-mediately above. This can be done by adding stages of sensing devices responsive to other parameters. One such parameter which can aid in the discrimination of very severe tachy-arrhythmias and fibrillation, is the variability in the R-to-R wave interval. As noted above, even during extremely high rate tachyarrhythmias, R-waves can be identified and generally occur at a stable rate. During fibrillation, on the other hand, all regularity in the autput of an R-wave detector is lost. It is therefore possible to discriminate between fibrillation and tachyarrhythmias by measuring the variability of the R-wave intervals by means of an R-wave detector. By combining the probability density function de-tector and an R-wave interval detector, it becomes a practi-cality to discriminate between fibrillation and even severe tachyarrhthmias with an accuracy never before attained.

A technique of ascertaining R-to-R wave interval ;

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variability by way of a phase lock loop will now be described.
The phase lock loop circuit has the capability of "locking"
onto periodic input signals and providing an AC output voltage which is at a constant phase and an integral multiple fre-quency with respect to the input. If the input is not periodic, however, the loop cannot "lock", and this condition is easily detected. By utilizing the probability density function de-tector as a first detector stage and a phase lock loop de-tector as a second detector stage, the absence o a locked state in the phase lock loop detector, coupled with the condi-tion of the first detector stage having issued a fibrillation output, verifies the presence of VF with an exceedingly high degree of accuracy. Phase lock loop circuits are well described in the literature, and an example of a low power version with lock indication, directly applicable to fibrillation detection, can be found in Application Note ICAN-6101, RAC COS/MOS Inte-grated Circuits, 1975 Databook Series, pp. 471-478. Accordingly, the phase lock loop circuitry is shown only in block form in Figures 11 and 12. Its application to a fibrillation de-tector is, however, a novel concept.
With reference now to Figures 10 and 11, the useof a phase lock loop in a fibrillation detection circuit will be described. The previously discussed probability density function fibrillation detector is an integral part of a first stage detector shown at 76. The input to the first stage detector 76 reaches an ECG amplifier 78, and is processed by the probability density function detector 80. If fibrilla-tion is sensed by the detector 80, then a signal is issued at line 82, and is fed to one terminal of an AND gate 84.
The second input terminal 86 of AND gate 84 is associated 2:~

with the second stage of the detector combination, and in particular, the phase lock loop circuit shown generally at 92.
The signal issued by the ECG amplifier 78 also serves as an input to a filter 88 which feeds filtered signals to an R-wave detector 90, each being of conventional design.
The ECG signal to filter 88 is illustrated in Figure lla, while the filtered signal serving as the input to the R-wave detector 90 is shown in Figure llb.
The R-wave detector 90 senses the presence of R-waves, and for each R-wave, issues a pulse of finite period.
If the R-waves are regular in interval, the output of the R-wave detector 90 is a periodic train of pulses. Figure llc illustrates the output of the R-wave detector 90 based upon the input of the filtered ECG shown in Figure llb. It will be noted that the first three pulses of R-wave detector 90 are periodic.
The phase lock loop 92 includes a phase detector 94, the output of which is filtered by a low-pass filter 96, in turn feeding signals to the control side of a voltage con-trolled oscillator 98. The oscillator 98 issues a regular train of square wave pulses and feeds the same to the phase detector 94, which then compares the phase of these regular pulses with the input from the R-wave detector 90.
The pase lock loop 92 may be of numerous designs, as these circuits are well known. In any event, it is contem-plated that the phase detector 94 provide output information for a lock detector 100 which is indicative of the phase rela-tionship between the R-wave detector pulses and the oscillator pulses, and in turn, which indicates whether the phase lock Z'~

loop 92 is able to lock upon the input from the R-wave de-tector 90.
Upon the lock detector 100 receiving an indication from the phase detector that the loop is locked, the detector 100 will, for example, issue a logical "0". Under this set of condition, the AND gate 84 will remain idle, even if the probability density detector 80 has indicated on lead 82 that fibrillation is present. On the other hand, if the lock de-tector 100 receives a signal indicating that the loop 92 is not locked, then a logical "l" will be issued to AND gate 84, and if this logical "l" occurs simultaneously with a similar signal from the probability density detector 80, then the AND gate 84 will issue a signal on line 102 which will trigger the de~ibrillating electronics. Still, however, if the phase lock loop 92 cannot lock, and yet the probability density detector 80 has sensed no irregularity, then the AND gate 84 will remain idle. This is illustrated in Figures lld through llf. Figure lld shows the output of the probability density function detector 80, Figure lle represents the output of the lock detector 100, and Figure llf represents the output of the AND gate 84 at lead 102.
Now, the impedance sensor VF detection circuit will be described. It has been found that the impedance between cardiac electrodes varies in accordance with the volume of blood in the heart. When in normal rhythm, the heart regularly contracts and fills, and hence the impedance change is periodic.
During fibrillation, however, stroke volume essentially goes to zero, and a severe drop in pulsatile impedance change can be seen. The circuit illustrated in block form in Figure 12 is able to detect the absence of pulsatile impedance changes, analogous to a drop in stroke volume and hence ventricular pressure. The traces of Figures 13a through 13f relate to the circuit illustrated in Figure ~
With reference then to Figures 12 and 13, the impedance VF detector is shown generally at 104. The detector 104 is powered by a power supply on line 106, actuated by a gate 108 which is, in turn, controlled by the probability density function detector shown at 110. As noted previously, the impedance VF detector requires a substantial amount of power from the implanted battery source. Therefore, so as not to drain the battery, the detector 104 is designed to remain idle until the probability density function detector 110 senses an abnormality, and triggers the impedance detector 104 by actuating its power supply. In this way, the circuit of Figure 12 provides an implied "AND" function. That is, the second-stage circuit 104, which triggers the defibrillating electronics, is only actuated upon command from the first-stage probability density function detector 110. Therefore both circuits must agree that fibrillation is present before a fibrillation output is generated.
The basic element in the impedance VF detector 104 is illustrated schematically as impedance 112. The impedance 112 is, for example, related to the impedance of the blood and tissue measured across intracardiac electrodes spaced apart on a catheter. A current source 114 associates with the impedance 112 and provides a current input of constant value. An oscillator 116 feeds the current source 114 so that source 114 generates an AC current to the impedance 112.
In this manner, the voltage across the impedance 112 will be proportional to the current multiplied by the impedance 2~

value. As typical values, the oscillator 116 is set to lOOKHz, with the current source 114 supplying lOO~a. The impedance 112 is typically on the order of 50 ohms, and therefore ap-proximately 5 mV appears across impedance 112. The voltage across the impedance 112 is then amplified by means of a voltage amplifier 118, and the amplified voltage from amplifier 118 is then demodulated by means of a synchronous demodulator 120.
The amplified and demodulated output of demodulator 120 is fed to a bandpass amplifier 122, and then to a trigger network 124, a ramp generator 126, and a threshold detector 128. The output of the threshold detector 128, if present, appears at terminal 120, and serves to trigger the defibrillation circuitry into operation.
Figure 13a represents an ECG which is at first normal, and then indicates fibrillation. Figure 13b shows, in an exaggerated form so as to appear on the same time scale, the output of oscillator 116, and Figure 13c represents a trace of the voltage across impedance 112 after amplification by -~ 20 amplifier 118 and corresponding to the ECG in Figure 13a.
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r~ It can be seen in Figure 13c that the voltage across impedance 112 increases for each normal beat of the heart as blood is ejected from the heart.
The output of demodulator 120, after amplification by amplifier 122, is illustrated in Figure 13d where a negative-going signal is indicated for each reduction in voltage, or pulse, across impedance 112. Ramp generator 126 develops a ramp which is shown in Figure 13e. It will be noted that the ramp returns to its baseline each time the demodulated and amplified output of amplifier 122 represented in Figure 13c 2:~ :

crosses a set threshold level. Accordingly, during normal cardiac rhythm, the threshold detector 128 remains inactive.
However, once fibrillation commences, where indicated in Figure 13a, the curve of Figure 13d smoothes out, without the threshold being reached, and therefore the ramp of Figure 13e continues to elevate until it exceeds the threshold of detector 128.
At this occurrence, detector 128 is triggered, and a fibrilla-tion output is issued on line 130.

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Claims (3)

The embodiments of the invention in which an exclusive property or privilege is claimed are defined as follows:

1. A two-stage apparatus for cardioverting a mal-functioning heart in such a manner that power is conserved, the apparatus comprising: first-stage detector means for discriminating between a normal cardiac state and a state requiring cardioversion; powering means for continually power-ing said first-stage detector means; second-stage detector means for further discriminating between a normal cardiac state and a state requiring cardioversion, normally in a stand-by condition and going to an active full power sensing con-dition upon initiation by said first-stage detector means;
means for enabling said second-stage detector means and for thereby bringing said second-stage detector means from its stand-by condition to its full power active sensing condition;
control means for powering said second-stage detector only upon said first-stage detector means sensing a cardiac state requiring cardioversion; delivery means for delivering cardio-verting energy to the heart; and means for actuating said delivery means upon said second-stage detector means sensing a cardiac state requiring cardioversion.

2. The apparatus recited in claim 1, wherein said second-stage detector means is an impedance detector which senses the pulsatile changes of impedance between two cardiac electrodes; and further comprising means for discriminating between normalcy in the pulsatile changes and a cardiac con-dition requiring cardioversion.

3. The apparatus recited in claim 1 or 2, wherein said second-stage detector is normally in a low power stand-by condition.