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Description  |
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BACKGROUND OF THE INVENTION
This invention relates to methods and apparatus for applying cardioversion
and defibrillation shocks to a patient's heart, and more particularly, to
methods and apparatus for applying charge-balanced shocks (waveforms) to
the heart.
Cardiac stimulating devices such as pacemakers and
cardioverter-defibrillators are well known. Typically, cardiac stimulating
devices contain sensing circuitry for monitoring the various heartbeat
signals produced by a patient's heart. Cardiac stimulating devices with
sensing circuitry can analyze the patient's heartbeat signals to determine
when and at what energy level any electrical pulses should be applied to
the heart.
Some cardiac stimulating devices can determine whether the patient is
suffering from an arrhythmia such an episode of tachycardia or a
fibrillation event. When an arrhythmia is detected, a cardiac stimulating
devices may attempt to terminate the arrhythmia by applying electrical
pulses to the patient's heart. These pulses may be in the form of high
energy cardioversion or defibrillation shocks. Cardioversion pulses have
energies in the range of about 2-5 J. Typical defibrillation shocks have
energies in the range of about 30-40 J.
Shocks are applied to the patient's heart via leads and electrodes. It is
well known that when current is passed through an electrode submerged in
an electrolyte such as interstitial fluid or blood, the
electrode/electrolyte interface polarizes. The shock-induced polarization
produces a voltage rise that can obscure the patient's heartbeat signals.
This effect, which is commonly known as "post shock block," is most
dramatic in systems in which the electrodes used to apply the shock are
adjacent to, or the same as, the electrodes used to sense the heartbeat
signals.
As a result of post shock block, whenever a cardioversion or defibrillation
shock is applied to the heart it becomes impossible for the sensing
circuitry to detect the heartbeat signals until the blocking voltage
decays. During the 5-30 second period before the shock induced voltage
decays, the cardiac stimulating device is not able to monitor the
patient's condition. The cardiac stimulating device is therefore unable to
determine whether or not the shock that was just applied was successful at
terminating the arrhythmia. A failure to successfully terminate a
tachycardia or a fibrillation might require that a more aggressive therapy
be applied in a further attempt to terminate the arrhythmia.
Unfortunately, due to post shock block, it is not possible to determine
the appropriate course of action until the polarization voltage induced by
the shock decays.
What is therefore needed is a way in which to reduce the effects of post
shock block.
SUMMARY OF THE INVENTION
In accordance with the principles of the present invention, methods and
apparatus are provided for applying charge-balanced cardioversion and
defibrillation shocks to a patient's heart.
When conventional cardioversion and defibrillation shocks are applied to
the heart to terminate arrhythmia episodes, the electrode/electrolyte
interface used to apply such shocks becomes polarized. The polarized
interface creates an effect known as post shock block, which obscures
heartbeat signals from the sensing circuitry used to monitor these
signals. Charge-balanced shocks have at least one positive shock phase and
at least one negative shock phase. The sum of the time-integrated currents
of the positive phases of the shock is equal to the sum of the
time-integrated currents of the negative phases of the shock. As a result,
the charge delivered to the electrode/electrolyte interface during the
positive phases is balanced by the charge removed during the negative
phases. Shocks of this type do not polarize the interface as conventional
shocks, thus significantly reducing the effects of post shock block.
Two particularly suitable charge-balanced shock waveforms are the biphasic
shock waveform and the triphasic shock waveform. The biphasic
charge-balanced shock waveform has a single positive shock phase and a
single negative shock phase. The time-integrated current of the positive
shock phase is equal to the time-integrated current of the negative shock
phase. When the charge-balanced biphasic shock waveform is applied to the
patient's heart to terminate an arrhythmia, the sensing
electrode/electrolyte interface is polarized less than when a conventional
shock waveform is used, thereby reducing the effects of post shock block.
A suitable triphasic charge-balanced shock waveform has either two positive
shock phases and a single negative shock phase or two negative phases and
a single positive phase. The sum of the time-integrated currents of the
two phases is equal to the time-integrated current of the shock phase with
the opposite polarity. When the charge-balanced triphasic shock waveform
is applied to the patient's heart to terminate an arrhythmia, the sensing
electrode/electrolyte interface is polarized less than with a conventional
shock, so that the effects of post shock block are significantly reduced.
The defibrillation efficacy of biphasic, and particularly triphasic,
charge-balanced shocks is comparable to or superior to that of
conventional shocks.
Charge-balanced shocks can be applied to the patient's heart using any
suitable lead arrangement. For example, a bipolar or a tripolar lead can
be used. A suitable storage capacitor within a cardiac stimulating device
can be charged when it is desired to apply the charge-balanced shock.
Switching circuitry is preferably used to alternately apply the positive
and negative shock phases to the heart by discharging the capacitor
through the blood and heart tissue of the patient. Control circuitry
ensures that the durations of the positive and negative shock phases are
such that the total magnitude of the time-integrated current of the
positive phases is equal to that of the negative phases.
BRIEF DESCRIPTION OF THE DRAWINGS
The above and other advantages of the invention will be apparent upon
consideration of the following detailed description, taken in conjunction
with the accompanying drawings, in which like reference numerals refer to
like parts throughout, and in which:
FIG. 1 is a perspective view of a bipolar lead arrangement connected to a
cardiac stimulating device;
FIG. 2 is a schematic diagram illustrating the effective circuit of the
can, coil, and tip electrodes of FIG. 1 in contact with the blood and
tissue of a heart;
FIG. 3 is a perspective view of a tripolar lead arrangement connected to a
cardiac stimulating device;
FIG. 4 is a simplified schematic diagram of the experimental arrangement
used to determine the magnitude of the post shock block effect for various
shock waveforms;
FIG. 5 is a conventional monophasic shock waveform;
FIG. 6 is a dual trace graph showing the effects of post shock block
following the application of the conventional monophasic shock of FIG. 5;
FIG. 7 is a conventional equal-phase duration biphasic shock waveform;
FIG. 8 is a dual trace graph showing the effects of post shock block
following the application of the conventional equal-phase biphasic shock
of FIG. 7;
FIG. 9 is an illustrative biphasic charge-balanced shock waveform;
FIG. 10 is a dual trace graph showing the effects of post shock block
following the application of the biphasic charge-balanced shock of FIG. 9;
FIG. 11 is an illustrative triphasic charge-balanced shock waveform;
FIG. 12 is a dual trace graph showing the effects of post shock block
following the application of the triphasic charge-balanced shock of FIG.
11;
FIG. 13 is an illustrative cardiac stimulating device with which
charge-balanced antiarrhythmia shocks can be applied to the heart; and
FIG. 14 is a flow chart depicting an illustrative process for applying the
positive and negative phases of a charge-balanced shock to the heart.
DETAILED DESCRIPTION OF PREFERRED EMBODIMENTS
Cardiac stimulating devices such as pacemakers and
cardioverter-defibrillators are well known. A variety of devices are
presently available that apply electrical pulses to a patient's heart in
order to maintain a healthy heart rhythm. Some cardiac stimulating devices
simply apply pacing pulses to the patient's heart at regular predetermined
intervals. More typically, cardiac stimulating devices contain sensing
circuitry for monitoring the various heartbeat signals produced by a
patient's heart. Cardiac stimulating devices with sensing circuitry can
analyze the patient's heartbeat signals to determine when and at what
energy level any electrical pulses should be applied to the heart.
Some cardiac stimulating devices can determine whether the patient is
suffering from an arrhythmia such an episode of tachycardia (a condition
in which the heart beats too quickly) or a fibrillation event (a condition
in which the heart quivers chaotically). When an arrhythmia is detected,
appropriate corrective therapies can be applied to the heart. Some cardiac
stimulating devices attempt to terminate tachycardia episodes by applying
bursts of fairly weak electrical pulses to the patient's heart. Other
cardiac stimulating devices can apply high energy shocks to the heart. For
example, some devices contain cardioversion circuitry, which allows
individual pulses having energies in the range of about 2-5 J to be
applied to the heart to terminate an arrhythmia. Cardiac stimulating
devices with defibrillation capabilities can apply still higher energy
pulses to terminate heart fibrillation. Typical defibrillation shocks have
energies in the range of about 30-40 J.
Shocks are applied to the patient's heart using leads containing various
electrodes. It is well known that when current is passed (e.g.,
defibrillation shock) through an electrode submerged in an electrolyte
such as interstitial fluid or blood, the electrode/electrolyte interface
polarizes (i.e., it will maintain a potential after the current flow
stops). The shock-induced polarization produces a voltage rise that can
obscure the patient's heartbeat signals. This effect, which is commonly
known as "post shock block," is most dramatic in systems in which the
electrodes used to apply the shock are adjacent to the electrodes used to
sense the heartbeat signals.
As a result of post shock block, whenever a cardioversion or defibrillation
shock is applied to the heart it becomes impossible for the sensing
circuitry to detect the heartbeat signals until the blocking voltage
decays. However, about 5-30 seconds may elapse before the voltage induced
by the shock decays sufficiently to allow the sensing circuitry to detect
the patient's heartbeat signals. During this period, the cardiac
stimulating device is not able to monitor the patient's condition. The
cardiac stimulating device is therefore unable to determine whether or not
the shock that was just applied was successful at terminating the
arrhythmia.
A successful termination of a tachycardia or fibrillation episode might
require that no pacing pulses be applied, or might require that pacing
pulses be applied only if the patient's heart does not beat on its own. A
failure to successfully terminate a tachycardia or a fibrillation might
require that a more aggressive therapy be applied in a further attempt to
terminate the arrhythmia. But because of post shock block, it is not
possible to determine the appropriate course of action until the
polarization voltage induced by the shock decays.
One approach for alleviating the effects of post shock block is to locate
the electrodes used to apply cardioversion and defibrillation shocks ("the
shock electrodes") at a position remote from the electrodes that are used
to sense the heartbeat signals ("the sense electrodes"). Although this is
theoretically possible, in practice a more compact lead arrangement having
at least one of the shock electrodes on the same lead as a sense electrode
is often preferred. Two commonly used lead arrangements are the bipolar
lead and the tripolar lead.
A typical bipolar lead 20 is shown in FIG. 1. The bipolar lead 20 has two
electrodes: a tip electrode 22 and a coil electrode 24. The bipolar lead
20 is connected to a cardiac stimulating device 28. The cardiac
stimulating device 28 contains circuitry for monitoring the patient's
heartbeat signals and for applying high energy pulses to the heart in
response to detected arrhythmias. The cardiac stimulating device 28 is
typically also capable of providing pacing pulses to the heart when
needed.
In operation, the bipolar lead 20 is surgically implanted in a patient's
heart (e.g., in the patient's right ventricle). The tip electrode 22
typically is capable of being attached to the heart wall. For example, the
tip electrode 22 may have tines 26, which hold the tip into place or the
tip electrode 22 may have a portion that allows it to be screwed into
place. The patient's heartbeat signals are monitored by measuring the
voltage between the tip electrode 22 and a coil electrode 24 using the
sensing circuitry contained within the cardiac stimulating device 28.
Thus, the sense electrodes in a bipolar lead are the tip electrode 22 and
the coil electrode 24.
If the cardiac stimulating device 28 has a pacing capability, pacing pulses
can be supplied to the heart via the tip electrode 22 and the coil
electrode 24. Because the coil electrode is much larger than the tip
electrode 22, the current density is greatest near the tip electrode 22,
so that most of the electrical stimulation of the pacing pulse is applied
to the heart in the vicinity of the tip electrode 22. In this
configuration, the coil electrode 24 is known as the indifferent electrode
for pacing and sensing.
High energy antiarrhythmia shocks are applied to the heart using the coil
electrode 24 and a can electrode 30 (the metallic housing of the cardiac
stimulating device 28). Thus, for a typical bipolar lead arrangement, the
two shock electrodes are formed by the coil electrode 24 (also the
indifferent pacing and sensing electrode) and the can electrode 30.
Because the can electrode 30 is much larger than the coil electrode 24 the
can 30 is the indifferent shock electrode.
When the bipolar lead configuration of FIG. 1 is used, shocks applied with
the coil electrode 24 and the can electrode 30 polarize the
electrode/electrolyte interfaces of these two electrodes. The equivalent
circuit of the arrangement of FIG. 1 is shown in FIG. 2. As is well known,
in cardiac stimulating devices the interface between the blood and heart
tissue of the patient and each electrode can be modelled as a capacitance
(the Warburg capacitance) in series with a resistance (the Warburg
resistance) both of which are in parallel with a resistance known as the
Faradic resistance. As shown in FIG. 2, for the lead configuration of FIG.
1, the Warburg capacitances of the can 30, the coil 24, and the tip 22 are
C.sub.CAN, C.sub.COIL, and C.sub.TIP. The Warburg resistances for the FIG.
1 system are R.sub.W(CAN), R.sub.W(COIL), and R.sub.W(TIP) and the Faradic
resistances are R.sub.F(CAN), R.sub.F(COIL), and R.sub.F(TIP). The
voltages V.sub.CAN, V.sub.COIL, and V.sub.TIP are the half-cell potentials
that develop between the respective electrodes and the surrounding blood
and heart tissue. The resistances R.sub.1, R.sub.2 and R.sub.3 represent
the bulk resistance of the blood and tissue between the electrodes.
When a shock is applied between the can electrode 30 and the coil electrode
24 (FIG. 1), the capacitances C.sub.CAN, C.sub.COIL, are charged by the
current that flows between these electrodes through the blood and heart
tissue. The capacitance C.sub.TIP is charged by leakage currents in the
sensing circuitry due to the defibrillation shock. Illustrative leakage
current paths between CAN, COIL and TIP, that pass through C.sub.TIP,
P.sub.1 and P.sub.2 are shown in FIG. 2. Leakage elements P.sub.1 and
P.sub.2 are parasitic pathways within the device's electronics and can be
resistive, capacitive and/or nonlinear elements, such as zener diodes,
used to protect the sensing electronics against transient high voltages.
The voltages on capacitors C.sub.COIL and C.sub.TIP cause the post shock
block effect. The various capacitances and resistances of FIG. 2
continually fluctuate, so it is not possible to determine what the
voltages induced on the capacitors will be when a given shock is applied
to the heart. As a result, when the sensing circuitry in the cardiac
stimulating device 28 (FIG. 1) attempts to monitor the patient's heartbeat
by measuring the voltage difference between TIP and COIL, the voltage
difference on C.sub.COIL and C.sub.TIP obscure the heartbeat signals.
Although capacitors C.sub.CAN, C.sub.COIL, and C.sub.TIP eventually
discharge through the various resistances, until the capacitors discharge
sufficiently to reduce the voltages on the capacitors below the level of
the patient's heartbeat signals, the cardiac stimulating device 28 (FIG.
1) is unable to monitor the patient's condition.
The effects of post-shock block are somewhat less severe when a tripolar
rather than a bipolar lead arrangement is used. As shown in FIG. 3, a
tripolar lead 32 may be connected to a cardiac stimulating device 34, in
much the same way as a bipolar lead. In contrast to the bipolar
arrangement, however, there is no common sense and shock electrode.
Sensing (and pacing) is performed between tip electrode 36 and a ring
electrode 38, which is the indifferent electrode for pacing and sensing.
Shocks are applied with a coil electrode 40 and a can electrode 42. The
can electrode 42 is the indifferent shock electrode.
When a defibrillation shock is applied to the patient's heart using the
tripolar lead configuration of FIG. 3, current flows primarily between the
coil electrode 40 and the can electrode 42. As a result, the capacitances
associated with the coil electrode 40 and the can electrode 42 are charged
and the electrode/electrolyte interface polarizes. However, because
sensing takes place between the tip electrode 36 and the ring electrode
38, which are isolated to some degree from the coil electrode 40 and the
can electrode 42, the post shock blocking effect is not as extreme as when
a bipolar lead arrangement is used. Nevertheless, the post shock blocking
effect is not eliminated, because leakage currents within the device's
electronic circuitry will still charge the capacitances associated with
the tip electrode 36 and the ring electrode 38.
Moreover, tripolar leads are generally not as attractive as bipolar leads
for applying cardioversion and defibrillation shocks. Tripolar leads are
inferior to bipolar leads in this respect because, in order to allow for
the presence of the ring electrodes, the shocking coils in tripolar leads
are shorter than those in bipolar leads and shorter coils have been found
to reduce the efficacy with which the coil electrode applies shocks to the
heart. The farthest distance along the leads that the coil electrodes in
both type of leads may extend is restricted by the size of the heart
chamber into which the leads are placed. However, in tripolar leads the
coil electrodes must accommodate the ring electrode 38. Thus, the coil
electrode 40 in tripolar lead 32 (FIG. 3) cannot extend as far along the
lead toward the tip electrode 36 (FIG. 3) as the coil electrode 24 (FIG.
2) can extend toward tip 22 (FIG. 2). Moreover, the construction of a
tripolar lead is more complex, due to the presence of an extra electrode
and the associate conductor and connections to it in the lead body. In
addition, tripolar leads have a larger lead diameter and a stiffer lead
body than bipolar leads, due to the presence of the extra conductor and
insulating materials.
Thus, for all lead systems it would be desirable to be able to reduce the
effects of post shock block. Reducing the effects of post shock block
would allow cardiac stimulating devices to monitor a patient's heartbeat
signals without becoming obscured by the blocking voltages induced by
during the application of conventional cardioversion and defibrillation
shocks. In accordance with the present invention, it has been determined
that the effects of post shock block can be significantly reduced if the
shocks that are applied to the heart are charge-balanced (i.e., contain no
net DC component). The experimental set-up shown schematically in FIG. 4
was used to confirm that shocks of this type significantly reduce the
effects of post shock block. A tank 44 of normal saline solution was used
to represent the blood and heart tissue of a patient. Single 1 V
peak-to-peak sine wave cycles of 10 ms in duration were generated at a
repetition rate of two pulses per second by voltage source 46. The sine
wave cycles produced signals in the saline solution comparable to the
normal heartbeat signals that occur in the heart. The signals in the
saline solution were monitored using a tip electrode 48 and a coil
electrode 50 of a lead 52.
The voltage between the tip electrode 48 and the coil electrode 50 was
measured using sensing circuitry 54. The measured voltage was provided to
a processing unit 58. When it was desired to simulate the effects of post
shock block, a high energy pulse from the pulse generating circuitry 54
was delivered to the saline solution via a metal plate electrode 56 and
the coil electrode 50.
In a cardiac stimulating device, pacing pulses are typically applied to the
heart through a blocking capacitor that filters out undesirable DC
currents. Because pacing pulses typically have magnitudes on the order of
several volts, it is feasible to generate pacing pulses directly from the
cardiac stimulating device battery using digital switching circuitry. In
order to generate cardioversion and defibrillation shocks, however, it is
necessary to produce significantly higher voltages (e.g., 50-1000 V). To
generate shocks with voltages of this magnitude, cardiac stimulating
devices contain a storage capacitor that is charged by the battery when it
is desired to apply a shock. Once the storage capacitor is charged to the
appropriate level, switching circuitry within the cardiac stimulating
device is used to discharge the capacitor through the shock electrodes
connected to the heart.
A commonly used cardioversion and defibrillation shock waveform is the
monophasic shock. A typical monophasic shock waveform is shown in FIG. 5.
The monophasic shock of FIG. 5 exhibits an exponential decay. The time
constant of the decay is equal to the product of the capacitance of the
storage capacitor used to generate the shock and the resistance of the
blood and heart tissue though which the capacitor is discharged.
The monophasic shock of FIG. 5 was applied to the saline solution using the
set-up of FIG. 4. The resulting voltage measured between the tip electrode
48 and the coil electrode 54 is shown in the top trace of FIG. 6. Whenever
the processing unit 58 determined that the voltage between the tip
electrode 48 and the coil electrode 50 exceeded a predetermined threshold,
the processing unit 58 confirmed that a heartbeat signal was detected.
Comparable algorithms are used in typical cardiac stimulating devices to
determine whether or not measured signals correspond to cardiac events.
As shown in the upper trace of FIG. 6, during the period from T0 to T1, the
peak-to-peak magnitude of the measured heartbeat signals, such as an
illustrative heartbeat signal 60, is approximately 3.5 units. The lower
trace in FIG. 6 shows how each of these heartbeat signals was successfully
detected. At time T1, the monophasic shock of FIG. 5 was applied to the
saline solution with the coil electrode 50 and the plate electrode 56
(FIG. 4). The signals subsequently measured between the tip electrode 48
and the coil electrode 50, such as illustrative heartbeat signal 62, have
a reduced magnitude (of approximately 2-2.5 units), and are therefore not
detected by the processing circuitry 58, as shown in the lower trace of
FIG. 6. Although a heartbeat signal is successfully detected at time T2
(nine seconds after the shock was applied), considerable undersensing of
the heartbeat signals persists long after the shock was applied. For
example, at T3, well over 30 seconds after the shock was applied, a
heartbeat signal was not properly detected.
Similar results were obtained when the conventional equal-phase duration
biphasic shock of FIG. 7 was applied. As shown in FIG. 8, after the
biphasic shock of FIG. 7 was applied at time T1, the heartbeat signals
were significantly undersensed for at least 30 seconds due to the effects
of post shock block. Thus, as shown in FIGS. 6 and 8, conventional
monophasic and biphasic shocks create substantial post shock block effects
that interfere with the ability of the processing unit 58 to detect the
heartbeat signals.
In contrast, when charge-balanced shocks are used, the effects of post
shock block are substantially reduced. A biphasic charge-balanced shock
waveform suitable for use as a cardioversion or defibrillation shock is
shown in FIG. 9. The shock of FIG. 9 is called a "charge-balanced" shock,
because the time-integrated current applied to the heart during the
positive phase of the shock (t=0 ms to approximately t=3 ms) is equal in
magnitude to the time-integrated current applied to the heart during the
negative phase of the shock (approximately t=3 ms to t=10 ms). Thus the
charge delivered to the blood and tissue by the tip electrode 48 (FIG. 4)
during the positive phase is balanced by the charge removed from this area
during the negative phase of the shock. By balancing the charge delivered
through the electrodes, the polarization of the electrode/electrolyte
interface--which is the cause of post shock block--is significantly
reduced.
Any suitable arrangement for applying charge-balanced shocks to the heart
may be used. One convenient approach for generating a charge-balanced
shock is to apply the shock to the heart from a charged storage capacitor.
With this approach, the shock will exhibit the same characteristic
exponential decay exhibited by the conventional shocks of FIGS. 5 and 7.
For charge-balanced biphasic shocks of the type shown in FIG. 9, a good
approximation of the current during the positive and negative phases of
the shock is given by Equation 1.
##EQU1##
The resistance of the blood and heart tissue through which the shock is
delivered is R. The storage capacitor used to deliver the shock has a
capacitance of C. The initial voltage across the capacitor is V.sub.0. In
Equation 1, t represents time. The duration of the positive shock phase is
t.sub.1. The total duration of the shock is t.sub.2. From t=0 (the
beginning of the shock) to t.sub.1, Shock.sub.-- Current is positive. From
t.sub.1 to t.sub.2, Shock.sub.-- Current is negative. From Equation 1 it
follows that the magnitude of the time-integrated current for the positive
phase of the shock is given by Equation 2.
##EQU2##
Performing the integration in Equation 2 one obtains the relationship in
Equation 3.
Pos.sub.-- Time.sub.-- Integrated.sub.-- Current=V.sub.0
C(1-e.sup.-t.sbsp.1.sup./RC) (3)
The magnitude of the time-integrated current for the negative phase of the
shock is given by Equation 4.
##EQU3##
Carrying out the integration in Equation 4 results in the expression of
Equation 5.
Neg.sub.-- Time.sub.-- Integrated.sub.-- Current=V.sub.0
C(e.sup.-t.sbsp.1.sup./RC -e.sup.-t.sbsp.2.sup./RC) (5)
For charge-balanced biphasic shocks, the magnitude of the time-integrated
current of the positive shock phase is equal to the magnitude of the
time-integrated current of the negative shock phase. Equations 3 and 5 can
therefore be equated to one another and the resulting expression solved
for the duration of the positive shock phase (t.sub.1) as a function of
the total shock duration (t.sub.2), as shown in Equation 6.
t.sub.1 =RC ln[2/(1+e.sup.(-t.sbsp.2.sup./RC))] (6)
For a chosen total shock duration (e.g., 10 ms), Equation 6 can be used to
determine the duration of the first (positive) phase to ensure that the
shock is charge-balanced, as a function of the resistance of the patient's
blood and heart tissue and the capacitance of the cardiac stimulating
device storage capacitor. In addition to using Equation 6 to select
appropriate relative durations for the positive and negative shock phases,
a value for the initial shock voltage, V.sub.0, must be selected (see
Equation 1). The voltage V.sub.0 is preferably selected using conventional
methods for determining the appropriate value for the strength of
cardioversion and defibrillation shocks.
The magnitude of the post shock block effect produced by the
charge-balanced biphasic shock of FIG. 9 was evaluated using the system of
FIG. 4. As shown in the upper trace in FIG. 10, during the period from T0
to T1, the peak-to-peak magnitude of the measured heartbeat signals, such
as an illustrative heartbeat signal 64, is approximately 3.5 units. The
lower trace in FIG. 10 shows how each of these heartbeat signals was
successfully detected. At time T1, the charge-balanced biphasic shock of
FIG. 9 was applied to the saline solution with the coil electrode 50 and
the plate electrode 56 (FIG. 4). At time T2, less than three seconds after
the charge-balanced shock was applied, the heartbeat signals measured
between the tip electrode 48 and the coil electrode 50, such as heartbeat
signal 66, were of sufficient magnitude (approximately 3.5 units) to be
successfully detected by the processing unit 58 (FIG. 4). Thus, as shown
on the lower trace in FIG. 10, the only interruption in the detection of
heartbeat signals was the period from T1 to T2, a delay of less than three
seconds.
Significant reductions in post shock block were also accomplished using the
triphasic charge-balanced shock shown in FIG. 11. The charge-balanced
shock of FIG. 11 is made up of three phases: a positive phase from
approximately 0 ms to 2.7 ms, a negative phase from approximately 2.7 ms
to 6.3 ms, and a positive phase from 6.3 ms to 10 ms. The triphasic shock
of FIG. 11 is charge-balanced because the time-integrated current applied
to the heart during the two positive phases of the shock is equal in
magnitude to the time-integrated current applied to the heart during the
negative phase of the shock.
A suitable approach for generating a charge-balanced triphasic shock is to
apply the shock | | |
