CA1087691A - Fibrillation monitor and defibrillator - Google Patents

Abstract

ABSTRACT OF THE DISCLOSURE

A circuit detects the state of a heart by monitoring the continuous time average of the ratio of high slope to low slope ECG segments and effects cardioversion if a malfunction is indicated by such time average ratio exceeding a predetermined threshold. The circuit comprises an ECG
monitor for sensing ECG signals from the heart, signal shaping circuitry for generating the slope of the sensed ECG signals by providing an approximation of the derivative of the input ECG, circuitry for discriminating between high slope and low slope segments, averaging circuitry for continuously time averaging the ratio of high slope to low slope segments, a threshold detector for determining whether such time average ratio of high slope to low slope segments is within predetermined threshold limits indicative of normalcy and circuitry for effecting cardioversion if said ratio is outside such threshold.

Description

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 suryival 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 identification of high risk VF patients, anti-arrhythmic drugs which can be pro-phylactically administered to these patients, programs for wide-spread cardio-pulmonary resuscitation training and implantable devices which can automatically detect VF and deliver cardioverting countershocks.
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 typically 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 impos-sible or impractical. There is accordingly a great need for an electronic device able to accurately 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 EC~, signal from the paddles.
Such a defibrillator could safely be used by even an untrained ~, operator. ...
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10876~1 With regard to the automatic implantable de~ibrillator, techniques have been developed which are generally acceptable for detecting VF and discriminating between life threatening arrhyth-mias and other cardiac malfunctions. Yet there is considerable room for improvement with re~ard 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, par-ticularly 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 specifically to whether failures should occur in a passive or an active manner.
Obviously, failures must be minimi~ed, 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 ~F or other life threatening arrhythmias, so that failure m~des may be passive.
The present invention is directed generally to the development of an accurate simple, VF detector which at least partly mitigates at least some of the drawbacks of known VF

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The present invention 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 ventricular fibril- -lation. Though presented as a part of an automatic implantable defibrillator, it should be appreciated that the present in-vention is not limited to this specific application. For ex-ample, certain other arrhythmias, or tachyarrhythmias can easily be identified by utilizing the teachings of the present invention.
Customarily, the term electrocardiogram (ECG) implies the use of electrodes on the body surface to obtain electrical signals 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 invention, VF may be detected with a degree of accuracy never before possible, and hence inherent passive failure modes can be afforded. The inventive detector enjoys operation independent from the concepts of QRS detection and heart rate calculations to maximize accuracy. As is known, these concepts are particularly difficult to define during ven- ; -tricular fibrillation. Furthermore, high-amplitude P and T-waves can inaccurately be sensed as R-waves, leading to false VF diag-nosis. The inventive VF detector has simple circuitry to minimize component count and therefore the possibility of electronic component failure~ ~nd, the circuitry o~ the inyentive VF detector is easily adaptable to low power operation.

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~: ~ . ' ' :; , 376~1 According to the invention, there is provided a cir-cuit for detecting the state of a heart by monitoring the con-tinuous time average of the ratio of high slope to low slope ECG segments and for effecting cardioversion if a malfunction is indicated by such time average ratio exceeding a predeter-mined threshold, the circuit comprising ECG monitor means for sensing ECG signals from a heart; signal shaping means for generating the slope of the sensed ECG signals by providing an approximation of the derivative of the input ECG; means for discriminating between high slope and low slope segments;
averaging means for continuously time averaging the ratio of high slope to low slope segments; threshold means for deter-mining whether such time average ratio of high slope to low slope segments is within predetermined threshold limits in-dicative of normalcy; and means for effecting cardioversion if the ratio is outside the threshold.
The invention further provides a circuit for detecting the state of a heart by monitoring the continuous time average of the ratio of high slope to low slope ECG segments and for effecting cardioversion if a malfunction is indicated by such time average ratio exceeding a predetermined threshold, the circuit comprising ECG monitor means for sensing ECG signals from a heart; signal shaping means for generating the slope ~ . ~
of the sensed ECG signals by providing an approximation of ~ -the derivative of the input ECG; automatic gain control means .
for normalizing the height of the derivative peaks, the auto-matic gain control means having a pick-off point at a location after the input ECG signals are shaped; means for discriminating between high slope and low slope segments; averaging means for continuously time averaging the ratio of high slope to : ---... . .

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and means for effecting cardioversion if the ratio is outside the threshold.
Reference is 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 density 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 are used interchangeably.
As will become apparent from the following description, VF
may be detected by monitoring a sampled probability density.
The probability density function can be monitored at any number of levels, but in a simple arrangement monitoring is ac~
complished 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.
There is also described hereinafter a second-stage VF detector which senses the regularity of the R-to-R interval.
It has been observed that during high rate tachyarrhythmias (on the order of 250 beats per minute), R-waves can still be identified, and almost always occur at a stable rate. During fibrillation, on the other hand, there are no such regular - 5a -~.~

R-waves. A novel second-stage detector described hereinafter utilizes a phase lock loop circuit stage which monitors the variability in the R-to-R interval. The loop locks onto regu-larly occuring R-waves, but if the R-to-R interval becomes irregular, as in VF, the loop cannot loek.
A second-stage detector in the form of an impedance sensor which measures impedance between cardiac electrodes is ~lso described hereinafter. It has been found that the impedance due to cardiac contractions is related to stroke volume. The impedance sensor requires a relatively ]arge input power to perform its sensing function, and hence a circuit may be pro-vided by which the impedance sensor remains idle for the greater majority of time, and is aetuated only upon the preceding de-tector stage sensing what is diagnosed to be VF.
It is accordingly the main objeet of the present invention to provide an aceurate deteetor of cardiae aetivity.
Embodiments of the invention will now be more par-tieularly deseribed with referenee to the aeeompanying drawings, -given by way of example, in whieh: -Figure l(a) is a traeing of a square wave given for exemplary purposes;
Figure l(b) is a plot of the probability density funetion of the wave illustrated in Figure l(a);
Figure 2(a) is a typieal eatheter sensed ECG traee;
Figure 2(b) is a plot of the probability density funetion of the ECG traee illustrated in Figure 2(a);

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Figure 3(a) is ~n ~CG trace representing ventricular fibrillation;
Figure 3(b) is a curve representing the probability density function of the ECG trace illustrated in Figure 3(a~;
Figure 4 is a block di~gram of a probability density function detector;
Figure 5 is a detailed circuit schematic of the detector illustrated in Figure 4; ~:~
Figure 6(a) is a curve of an exemplary input ECG signal to the detector circuit of Figures 4 and 5, showing both normal cardiac rhythm and fibrillation;
Figures 6(b) through 6(e) are curves representing signals at select locations in the circuit ].llustrated in :
Figures 4 and 5 based upon the ECG input illustrated in Figure 6(a);
` Figure 7 is a block diagram of a circuit for developing probability density function traces for the input of an oscilloscope Figures 8(a) through 8(d) are curves illustrating an ideal example of filtering an ECG trace to move the probability density function to zero;
Figure 9(a) is a curve similar to that illustrated in Figure 2(a), ~ut representing the ECG trace after filtering;
Figure 9(b) is a probability density function similar to that shown in Figure 2(b), but illustrating the function of the filtered ECG of Figure 9(a), - Figure 10 is a block diagram of a phase lock loop second-stage detector ~-.

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Figures ll(a) through ll(f) represent signals at select locations in the circuit illustrated in Figure 10;
Figure 12 is a block diagram of a second-stage impedance sensor for detecting ventricular fibrillation; and Figures 13(a~ through 13(f) 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 explana-tion of the circuit, there follows a brief discussionof the theory of probability density.
The detector system described hereinafter, -is based upon a series of measurements on the ECG. The meas-urements are known in the literature as the probability density function, denoted as K (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 Xl and X2. Looking at ~ -the simplified example illustrated in Figure l(a), 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 l(b), 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 probability density function of an ECG changes markedly between normal cardiac rhythm and ventricular fibrillation.
In this regard, the attention of the reader is directed to Figure ~

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2(a~ which illustrates a typical ECG trace, Fi~ure 2(b~ which shows the probability density ~unction of the ECG illustrated in Figure 2(a), Figure 3(a) which illustrates an exemplary ECG trace representing ventricular fibrillation, and Figure 3(b) which is the probability density function of the trace illustrated in - Figure 3(a). It will be noted that when comparing normal cardiac rhythm with ventricular fibrillation, the greatest 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 comparing Figures 2(b) and 3(b).
In a most simplified arrangement the probability density is sampled at one value of x, namely X=O 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. Phy-siologically then, sampling the probability density of the fil-tered ECG at X=O 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 form in ~ Figure 5. A representative set of waveforms is illustrated in - Figure 6.

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', lU~7~i91 The detector circuit is shown generally at 10, having a first stage ECG section 12, followed by a window stage 14, inte-grator 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 preamplifier 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 characteristic.
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 pr~vide an approximation 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 facilitate 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 capa-citors 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 would appear at terminal 20 and the corresponding ;-output of the ECG section 12 which would appear at terminal 30 are illustrated in Figures 6 (a~ and 6(b~, 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.
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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 com-- parator 32 lies inside or outside a band centered about a given window level introduced at terminal 60. In the simplified embodi-ment 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 6(b), and the resultant digital signal devel-oped by window comparator 32 and appearing at terminal 34 can be seen in Figure 6(c). 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 6(c) that upon the onset of fibrillation, very little time is spent by the fil-tered ECG signal inside the designated band, corresponding to the lower value of the probability density function at X=O as shown in Figures l(b) and 3(b).
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 illustrated in Figure 6(d). As can be seen, this output signal takes the form - of a ramp when fibrillation begins. This output signal at terminal 38 in turn becomes an input to the threshold detector, or compar-ator 40. Detector 40 then switches when the ramp signal at terminal 38 reaches a given threshold level. Hysteresis is pro-vided in the threshold detector stage 40 for a latching function so that the ramp must fall past level Vt to-Vh (shown on the transfer characteristic beneath detector 40~ for fibrillation 1~8'769:~l detection to cease. This is indicated in ~iguxe 6 during the period of inactive fibrillation shown in Figure 6~a~ wherein the trace of Figure 6(d) 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 simplified detector configuration provides inherent passive failure mode behavior 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 den-sity 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 retained, but are displayed in a different format. Thus information of general diagnostic significance is 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 probability density provide a great deal of information in the detection and study of tachyarrythmias.
In Figure 7, an input signal is introduced at input terminal 62, to be then passed through an automatic gain control circuit 64. In this manner, input signals of dlfferent amplitude ; can be handled by the overall circuit. On the probability density display, sign~l amplitude appears on the abscissa, and therefore, AGC will normalize the width of the display.

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108769'1 A digital stora~e element 66 follows AGC 64, and serves to proyide a repetitive source 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 precisely analogous to the definition of the distribution as defined 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 of Figures 2, 3 and 9 were de-veloped from the circuit illustrated in Figure 7. T~e trace of Figure 2(a), as noted above, represents an electrogram recorded from an intracardiac catheter. The corresponding density function is shown in Figure 2(b). Several regions have been identified on the respective traces, representing the same cardiac events but on the two different display patterns. For example, the region "B" of Figure 2(a~ is the most negative peak of the R-wave. The signal spends very little time at this yalue, thus the corresponding peak of the probability density curve of Figure 2(b) is small. The peak at "A" is representative of the ST segment, and as is readily seen, the ECG signal spends more time on the aveXage at this level than .

at region "s". Accordingl~, the peak is higher at "A" in Fi~ure ~ -2(b). The ECG dwells longest at the baseline identified at "C" in Figure 2(a), and the zero peak is the largest in Figure 2(b).
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 orignal 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 l, reference should be made to Figure 8. In Figure 8(a), there is illustrated a square wave alternating between "+A" and "-A"~
The probability density function of this square wave is given in Figure 8(b) and is similar to that shown in Figure l(b). Since the square wave spends no time at X=O, the probability density function has no peak at X=O. Figure 8(c) represents an impulse train which is developed by taking the derivative of the square wave illus-trated in Figure 8(a). The distribution function (probability density function) of the impulse train, unlike that illustrated in Figure 8(b), is a unit impulse at zero, as shown in Figure 8(d).
Thus, the effect of taking the derivative of the original square wave input and then evaluating the probability density function of the derivative is to shift peaks to X=O. ~:
The same principle as explained in the ideal case of Figure 8 is applicable to the filtered ECG shown in Figure 9(a). It will be recalled that the trace of Figure 9(a) represents the curve -~of Figure 2(a) after filtering. As can be seen, the peak "A"
corresponding to the ST segment which appears in Figure 2(b) has been eliminated from the probability density function wave of Figure 9(b). Furthermore, the Figure 9(b) zero peak is consider-:~., C

87~91 ably larger than that of Figure 2(b). Thus, the filter improyes the detection accuracy by enhancing the zero peak of the probability density function and thereby emphasizing the measure of the dif-ferences 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 de-scribed above. If two sampling points, say Xl and XO are defined as shown in Figures 3(b) and 9(b~, more discrimination resolution becomes available by taking a ratio. As illustrated, approximate measurement would show the value of the probability density func-tion at these two points on the waveform for the two examples to be:

For Normal Rhythm Kx(Xl) = .012 KX(Xo~ = 2.5 Cm = Kx(Xl) = .012 KX(Xo) 2.5 Cm = .0048 For Ventricular Fibrillation Kx(Xl) = .08 KX(Xo) = .11 C = Kx(Xl~ = .72 - `
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10~7691 It can readily be seen that this measure yields over two orders of magnitude difference between normal cardiac rhythm and fibril-lation. Sensing a value of Cm near 1.0 corresponds to the de-tection of a severe arrhythmia.
In view of the high degree of reliability necessary for - the successful application of an implantable automatic defibril-lator, it may become desirable to improve the accuracy of the detection system even relative to that described immediately above. This can be done by adding stages of sensing devices res-ponsive to other parameters. One such parameter which can aid in the discrimination of very severe tachyarrhythmias and fibril-lation, 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 ; fibril-lation, on the other hand, all regularity in the output of an R-wave detector is lost. It is therefore possible to discrimin-ate between fibrillation and tachyarrythmias by measuring the ~-variability of the R-wave intervals by means of an R-wave detector.
By combining the probability density function detector and an R-wave interval detector, it becomes a practicality to discriminate be- -tween fibrillation and even severe tachyarrythmias with an accuracy never before attained.
A technique of ascertaining R-to-R wave interval 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 frequency 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 , :-: :, . , . , 108769~

detector as a first detector stage and a phase lock loop detector as a second detector stage, the absence of a locked state in the phase lock loop detector, coupled with the condition 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, RCA COS/MOS Integrated 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 detector is, how-ever, a novel concept.
With reference now to Figures 10 and 11, the use of 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 prob-ability density function detector 80. If fibrillation is sensedby 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 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 ~f con~entional design. The ECG signal to filter 88 is illustrated in Figure ll(a), while the filtered signal serving as the input to the R-wave detector 90 is shown in ~ :
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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 ll(c) illustrates the output of the R-wave detector 90 based upon the input of the - filtered ECG shown in Figure ll(b). 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 controlled 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 phase lock loop 92 may be of numerous designs, as these circuits are well known. In any event, it is contemplated that the phase detector 94 provide output information for a lock detector 100 which is indicative of the phase relationship between the R-wave detector pulses and the oscillator pulses, and in turn, which indicates whether the phase lock loop 92 is able to lock upon the input from the R-wave detec-tor 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 "O". Under this set of conditions, 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 detector 100 receives a signal indicating that the loop 92 is not locked, then a logical "1" will be issued to AND gate 84, and if this logical "1" occurs simul-~ -18-1~87~91 taneously with a similar signal ~rom the probability density detector 80, then the AND gate 84 will issue a signal on line 102 which will trigger the defibrillating electronics. Still, how-ever, if the phase lock loop 92 cannot lock, and yet the prob-ability density detector 80 h~s sensed no irregulaxity, then the AND gate 84 will remain idle. This is illustrated in Figures ll(d) through ll(f~. Figure 11(d~ shows the output of the prob-ability density function detector 80, Figure ll(e) represents the output of the lock detector 100, and Figure ll(f) 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 fib-rillation, 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 13la) through 13(f) relate to the circuit illustrated in Figure 11.
With reference then to Figures 12 and 13, the impedance VF detector is shown generally at 104. The detector 104 is powered b~ 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 C "-- : - - : . :.

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7~91 density function detector llO 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 defib-rillating 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.

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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 pro-vides a current input of constant value. An oscillator 116 feeds the current source 114 so that source 114 generates an AC current Ih A to the impedance 112. .I~ this manner, the voltage across the impedance 112 will be proportional to the current multiplied by the impedance value. As typical values, the oscillator 116 is set to lOOKHz, with the current source 114 supplying 10~ a. The impedance 112 is typically on the order of 50 ohms, and therefore approximately 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 amplifiex 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 tri~ger the defibrillation circuitry into operation.

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Figure 13(a) represents an ECG which is at first normal, and then indicates fibrillation. ~igure 13(b~ shows, in an exag-gerated form so as to appear on the same ti~lle scale, the output of oscillator 116, and Figure 13~cl represents a trace of the voltage across impedance 112 after amplification by amplifier 118 and - corresponding to the ECG in Fi~ure 13(a). It can be seen in Figure 13(c) 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 13(d) 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 13(e~. It will be noted that the ramp returns to its baseline each time the demodulated and amplified output of amplifier 122 represented in Figure 13(c~ crosses a set threshold level. Accordingly, during normal cardiac rhythm, the threshold detector 128 remains inactive. However, once fibrillation commences, where indicated in Figure 13(a), the curve of Figure 13(d) smoothes out, without the threshold being reached, and therefore the ramp of Figure 13(e) continues to elevate until it exceeds the threshold of detector 128. At this occurrence, -~
detector 128 is triggered, and a fibrillation output is issued on line 130.

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Claims (16)

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

1. A circuit for detecting the state of a heart by monitoring the continuous time average of the ratio of high slope to low slope ECG segments and for effecting cardio-version if a malfunction is indicated by such time average ratio exceeding a predetermined threshold, the circuit compris-ing: ECG monitor means for sensing ECG signals from a heart;
signal shaping means for generating the slope of the sensed ECG signals by providing an approximation of the derivative of the input ECG; means for discriminating between high slope and low slope segments; averaging means for continuously time averaging the ratio of high slope to low slope segments;
threshold means for determining whether such time average ratio of high slope to low slope segments is within predetermined threshold limits indicative of normalcy; and means for effecting cardioversion if said ratio is outside said threshold.

2. The apparatus recited in claim 1, and further comprising automatic gain control means having a pick-off point at a location after said ECG signals are shaped.

3. The apparatus recited in claim 1, wherein said signal shaping means includes a high pass filter.

4. The apparatus recited in claim 1 or 2, wherein said signal shaping means is a filter having bandpass filter characteristics.

5. The apparatus recited in claim 1, wherein said discriminating means is a window comparator having a threshold band, and wherein said window comparator issues digital output signals, the sense of which is dependent upon whether shaped ECG signals lie inside or outside said threshold band.

6. The apparatus recited in claim 5, wherein said window comparator has a window level centered about the base line of said filtered ECG signals corresponding to zero slope.

7. The apparatus recited in claim 1, 2 or 3, wherein said discriminating means includes an absolute value circuit and a level comparator.

8. The circuit recited in claim 1, 2 or 3, wherein said means for effecting cardioversion is a pulse generator.

9. A circuit for detecting the state of a heart by monitoring the continuous time average of the ratio of high slope to low slope ECG segments and for effecting cardio-version if a malfunction is indicated by such time average ratio exceeding a predetermined threshold, the circuit compris-ing: ECG monitor means for sensing ECG signals from a heart;

signal shaping means for generating the slope of the sensed ECG signals by providing an approximation of the derivative of the input ECG; automatic gain control means for normalizing the height of the derivative peaks, said automatic gain control means having a pick-off point at a location after the input ECG signals are shaped; means for discriminating between high slope and low slope segments; averaging means for continu-ously time averaging the ratio of high slope to low slope segments; threshold means for determining whether such time average ratio of high slope to low slope segments is within predetermined threshold limits indicative of normalcy; and means for effecting cardioversion if said ratio is outside said threshold.

10. The circuit recited in claim 9, wherein said signal shaping means is a high pass filter.

11. The circuit recited in claim 9, wherein said automatic gain control means includes a peak-to-peak detector.

12. The circuit recited in claim 9, 10 or 11, wherein said discriminating means includes an absolute value circuit and a level comparator.

13. The circuit recited in claim 1, 10 or 11, wherein said averaging means includes an integrator.

14. The circuit recited in claim 1, 10 or 11, wherein said averaging means includes a capacitor.

15. The circuit recited in claim 9, wherein said threshold means is a comparator.

16. The circuit recited in claim 15, wherein said comparator has threshold hysteresis.