Note: Descriptions are shown in the official language in which they were submitted.
1087691
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. ...
,i :'
-2-
) ,
qF~ '
.,
-;
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
_3_
C
' ~ :
1087~91
detectors.
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.
f~ -4-
.
.. . .
~: ~ . ' ' :; ,
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
: -
--
... . .
, - ~. '
1087691
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 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 -
~.~
1087691
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);
C ~ ~
lV8769~;
., . : . .
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 ~-.
-7- -
' : ' '
'
. .
.
1~)876!~1 ~
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 ~
;',:, :"-'
--8--
S ~
.. . . . - : : :
,: ~, .~ : ,
7~
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.
,.
' ~
.
',
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.
.
lQ
.
~ . , - . , : .
- . . . . . . . .. ..
1(3~769~L`
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.
'.,' ''~,.....
-12-
.... ...
- , , . ~ . ' . . : . ,
. .
. . . .
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
-13-
.
1087691
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-
:~.,
-14-
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 - `
m
KX(X
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
-16-
,
:-: :, . , . ,
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 ~ :
Figure ll(b).
.
-17-
': ' . '" '- . , ~ .
1087691
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
--19--
C "-
- : - - : . :.
.: . .
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.
, . . .
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.
-20-
'~
.: ': ''', . .. :
1~8~9~
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.
-21-
.
