Cellular Action Potential and Excitable Gap During Fibrillation

In the past few years, knowledge of the characteristics of the action potentials during fibrillation has increased greatly. This is a direct result of the introduction of techniques for recording action potentials in whole hearts, either in vivo³¹ or in perfused, isolated hearts.^4–6,10,31 During fibrillation, the action potentials are altered: the APD is decreased; the action potential upstroke is slowed (decreased first-order derivative, dV/dt) and of decreased magnitude; the plateau phase is abbreviated; and DIs are abbreviated or absent (Fig. 2-3). During the first few seconds of VF (or atrial fibrillation), the activation rate is quite rapid; the mean cycle length of VF in patients undergoing defibrillator implantation was 213 ± 27 msec.⁶ DIs are rarely seen during early fibrillation, and the upstroke of most action potentials occurs before the transmembrane potential has returned to baseline from the previous action potential. The demonstration of an excitable gap in fibrillating atrial tissue,³⁴ as well as evidence of an excitable gap in fibrillating ventricular tissue,³⁵ suggests that there are periods late in the action potential in the fibrillating myocardium during which an electrical stimulus can capture a portion of the fibrillating myocardium. Knowledge of an excitable gap provides an opportunity to stimulate the tissue just in front of a fibrillating wavefront, to cause wavefront block.

Figure 2-3 Recording during ventricular fibrillation in a human.

Leads I, II, and III are body surface electrocardiograms. Note that there is no period of diastole between action potentials. MAP, Right ventricular monophasic action potentials; ECM, local bipolar electrogram.

(From Swartz JF, Jones JL, Fletcher RD: Characterization of ventricular fibrillation based on monophasic action potential morphology in the human heart. Circulation 87:1907-1914, 1993.)

As described in Chapter 1, the electrical activity of the heart is controlled ultimately by ion channels located in the cell membrane of the myocyte. It has been established that both the voltage-gated fast channels (sodium [Na⁺]) and slow channels (Na⁺ and calcium [Ca²⁺]) are active during the first few seconds of VF.⁸ The fast-channel activity is indicated by the rapidity of the upstroke of the action potential (phase 0) during early fibrillation and its sensitivity to administration of the Na⁺ channel blocker, tetrodotoxin. As fibrillation proceeds, the upstrokes of the action potentials become increasingly slower, with a decreased dV/dt_max, but the action potentials remain sensitive to tetrodotoxin until 1 to 5 minutes after initiation of fibrillation. A transition then occurs in which the action potential upstrokes become insensitive to tetrodotoxin. This suggests that the propagation of the action potential is no longer mediated primarily by the fast voltage-gated Na⁺ channels and may be mediated by slow voltage-gated Ca²⁺ channel activity in the later stages of fibrillation.⁴ The activation complexes recorded from the ventricular myocardium remain active only as long as the coronary arteries are perfused with oxygenated blood, suggesting that ischemia may be responsible for the loss of fast-channel activation during prolonged VF.²⁶

Waveforms, Current Strength, and Distribution During Defibrillation

The two most common waveform shapes used clinically are the monophasic and biphasic waveforms. In monophasic waveforms, the polarity of the shock is unchanged at each electrode for the entire duration of the electrical shock. In biphasic waveforms, the polarity of the shock reverses at each electrode partway through the defibrillation waveform. Many studies, in both animals and humans, have shown that biphasic waveforms can defibrillate with less current and energy than monophasic waveforms, in both internal and transthoracic defibrillation configurations.^37–40 Within each type, waveforms can be described as truncated exponential or damped sinusoidal shapes. ICDs use truncated exponential biphasic waveforms. Most external defibrillators have used damped sinusoidal monophasic waveforms, but because of the inductor necessary to shape the waveform, these defibrillators tend to be large and heavy. More recently, smaller, lighter external defibrillators have been developed that use truncated exponential biphasic waveforms similar to those used in ICDs. Damped sinusoidal biphasic waveforms are used in external defibrillators in Russia; similar to truncated exponential biphasic waveforms, these show improved efficacy over monophasic waveforms.^41,42

However, not all biphasic waveforms are superior to monophasic waveforms, For example, if the second phase of the biphasic waveform becomes much longer than the first phase, the energy required for defibrillation increases and can eventually rise to a level greater than the energy required to defibrillate with a monophasic waveform (with duration equal to the first phase of the biphasic waveform).^40,43,44 The optimum duration of the two phases of the biphasic waveform depends on the electrode impedance and the defibrillator capacitance.^45–48

Several groups have shown that defibrillation efficacy for square waveforms follows a strength-duration relationship similar to that for cardiac stimulation;^49,50 as the waveform becomes longer, the average current at the 50% success point (the current when one half of delivered shocks will succeed) becomes progressively less, approaching an asymptote called the rheobase.⁵¹ On the basis of this observation, some suggest that cardiac defibrillation can be mathematically modeled using a resistor-capacitor (RC) network to represent the heart^45,51–53 (Fig. 2-4). As empirically determined, the time constant for the parallel RC network is 2.5 to 5 msec.^45,47,53 In one version of the model,⁵³ a current waveform is applied to the RC network. The voltage across the network is then calculated for each time point during the defibrillation pulse. The relative efficacies of different waveform shapes and durations can be compared by determining the current strength that is necessary to make the voltage across the RC network reach a particular value, called the defibrillation threshold.

Figure 2-4 Parallel resistor-capacitor (RC) network representation of heart.

Responses of RC network to monophasic and biphasic truncated exponential waveforms with a time constant of 7 msec. The parallel RC network has a time constant of 2.8 msec. A, Input monophasic waveforms. The leading-edge current of the input waveform was 10 A. The waveforms were truncated at 1, 2, 3, 4, 5, 6, 8, and 10 msec. B, Model response, V(t). Initially, as the waveform lengthens, V(t) increases until it reaches a maximum at about 4 msec, after which it begins to decrease. C, Input biphasic waveforms. The leading-edge current was 10 A. Phase 1 was truncated at 6 msec. Phase 2 was truncated after 1, 2, 3, 4, 5, 6, 7, and 8 msec. D, Model response does not change polarity until the phase 2 duration is longer than 2 msec.

(From Walcott GP, Walker RG, Cates AW, et al: Choosing the optimal monophasic and biphasic waveforms for ventricular defibrillation. J Cardiovasc Electrophysiol 6:737-750, 1995.)

Several observations can be made from this model. First, for square waves, as the waveform duration becomes longer, the voltage across the network increases, approaching an asymptote, or rheobase. For truncated exponential waveforms, however, the model voltage rises, reaches a peak, and then, if the waveform is long enough, begins to decrease (see Fig. 2-4). Therefore, the model predicts that monophasic exponential waveforms should be truncated at a time when the peak voltage across the RC network is reached. Current or energy delivered after that point is wasted. In support of this prediction, strength-duration relationships measured in both animals⁵³ and humans⁵⁴ do not approach an asymptote but rather reach a minimum and remain constant as the waveform lengthens. This minimum occurs over a range of waveform durations and does not extend indefinitely. Schuder et al.⁵⁵ showed that, if the duration of a waveform becomes too long, defibrillation efficacy decreases.

Second, the model predicts that the heart acts as a low-pass filter.⁵² Therefore, waveforms that rise gradually should have better efficacy than waveforms that turn on immediately. This prediction has been shown to hold true for external defibrillation,⁵⁶ internal atrial defibrillation,⁵⁷ and internal ventricular defibrillation.⁵⁸ Ascending ramps defibrillate with a greater efficacy than do descending ramps.^58,59

Third, several groups have suggested that the optimal first phase of a biphasic waveform is the optimal monophasic waveform.^46,53 If so, what does the model predict as the “best” second phase of a biphasic waveform? Empirically, the role of the second phase apparently is to return the model voltage response back to zero as quickly as possible, thereby maximizing the increased efficacy of the biphasic waveform over that of the monophasic waveform with the same duration as phase 1 of the biphasic waveform. If the network voltage does not reach zero, or if it overshoots zero, efficacy is lost. Swerdlow et al.⁴⁷ showed in humans that the “best” phase 2 of a biphasic waveform returns the model response close to zero.

Together, these models allow the clinician to choose optimal capacitor sizes for truncated exponential biphasic waveforms, the waveforms most often used in ICDs. The capacitor must be large enough to be able to raise the network voltage to its threshold value and still hold enough charge to drive the network voltage back to zero. For a 40-Ω interelectrode impedance and a network time constant of 2.8 msec, the minimum capacitor that can accomplish this has a capacitance of 75 microfarads (µF).

These models can also help in choosing shock parameters, including phase duration for biphasic waveforms. Shock duration should be varied depending on defibrillation lead impedance; patients with low impedance should receive short-duration shocks, whereas patients with high impedance should receive longer-duration shocks. Initially, the concept of “tilt” was used to vary shock phase durations with impedance. Tilt is defined as the quotient of the leading-edge voltage minus the trailing voltage at the end of the shock divided by the voltage at the beginning of the shock, expressed as a percentage. For example, a shock with starting voltage of 500 V and ending voltage of 100 V has a tilt of 80%. If tilt is held constant, shock duration varies linearly with shock impedance. Compared with model predictions (Fig. 2-5), fixed-tilt waveforms are too short for patients with low impedance and too long for patients with higher impedance. One defibrillator manufacturer has developed a reference table that gives waveform phase durations as a function of patient impedance using a model similar to the one previously presented.^60,61 Their model uses “model time constants” of 2.5, 3.5, and 4.5 msec. It is recommended that the 3.5-msec constant be used first. If the defibrillation threshold for the 3.5-msec waveform is unsatisfactory, using the 4.5-msec constant is likely to be more effective than 2.5 msec. Waveforms that are too short tend to fail; waveforms that are too long tend to waste shock energy yet still defibrillate.

Figure 2-5 Waveform durations.

Optimal duration of first phase of a biphasic waveform as a function of shock time constant is shown for model responses of 2.0, 3.5, and 4.5 msec. Also shown are the first-phase durations for 50%-tilt and 65%-tilt waveforms.

The location of the defibrillation electrodes affects the magnitude of the shock necessary to defibrillate the heart. Typically, 200 to 360 J of energy are necessary for successful defibrillation, with the defibrillation electrodes located on the body surface, during transthoracic defibrillation with a damped sinusoidal monophasic waveform. Although less energy is required for a truncated exponential biphasic waveform,⁶² only about 4% to 20% of the current that is delivered to transthoracic defibrillation electrodes ever reaches the heart.^63,64 Indeed, when the defibrillation electrodes are placed in the heart itself, usually only 20 to 34 J of energy is required, and the requirement may be as low as only a few joules when large, contoured epicardial electrodes are used.^40,65 The strength of the shock also varies for different locations on or in the heart; epicardial patches defibrillate with a lower shock energy than transvenous electrode configurations.⁶⁶

Although defibrillation efficacy is usually described by some measure of defibrillation shock “strength” (energy, voltage, or current), little insight into the mechanisms of defibrillation can be obtained from these measures. Knowing how the current (or voltage) of a defibrillation shock is distributed over the heart allows greater understanding of how defibrillation occurs. Several studies have measured the potential gradient distribution throughout the heart during a defibrillation shock.^14,67,68 The potential gradient is a measure of the spatial variation of shock voltage across the heart. The potential gradient is measured in volts per centimeter (V/cm) of tissue. In a region with a high potential gradient, the difference in voltage between a given point and an area 1 cm distant from that point is high. Regions of low potential gradient have a measured voltage that is similar to that of nearby points. These studies show an uneven potential distribution for most electrode configurations, with areas of high potential gradient near the defibrillation electrodes and areas of low potential gradient in regions distant from the defibrillating electrodes. It has been hypothesized that a minimum potential gradient must be attained for successful defibrillation to occur, and that this requirement is independent of the current applied or the electrode configuration.^15,16 After a shock that fails to defibrillate VF, the site of earliest activation immediately after the shock can be mapped and related to the electric field that was produced by the shock. For shocks near the defibrillation threshold, the sites of earliest activation after a failed shock occur in the areas of lowest potential gradient.

The minimum potential gradient required for defibrillation is lower for biphasic than for monophasic waveforms (4 vs. 6 V/cm). A minimum potential gradient of 6 V/cm was required for successful defibrillation using a 10-msec truncated exponential monophasic waveform in the open-chest dog model.¹⁶ Similar findings were observed using a 14-msec truncated exponential monophasic waveform and multiple electrode configurations.¹⁵ In contrast, a minimum potential gradient of 4 V/cm was required for successful defibrillation using a truncated exponential biphasic waveform. Because higher shock strengths are required to induce a higher potential gradient, biphasic shocks successfully defibrillate with lower energy than monophasic shocks (i.e., a lower-voltage gradient is required).

The requirement for a minimum potential gradient may reflect the need for a shock to prevent the generation of new activation fronts that can result in reinitiation of fibrillation.⁶⁹ Examination of activation patterns after failed defibrillation for progressively larger shock strengths indicate that postshock activation occurs at numerous sites throughout the ventricle, and that reentry is common when the shock strength is much lower than that needed for defibrillation⁷⁰ (Fig. 2-6). At shock strengths just lower than those required for defibrillation, postshock activation arises in a limited number of myocardial regions. The activation fronts then propagate to activate other regions of the myocardium for a few cycles before reentry occurs; activation becomes disorganized; and fibrillation is reinitiated. Although postshock activation sites can still arise in regions of lowest potential gradient after a shock slightly greater than that required for defibrillation, the cycles of activation that originate from these sites are slower. These activations terminate after a few cycles without reinitiating fibrillation.^69–71

Figure 2-6 Phase maps of single rabbit heart.

Maps show the last cycle before a shock (left side) and the first cycle after a shock (right side) of 100 V (A), 200 V (B), 600 V (C), and 800 V (D) that failed to defibrillate. The defibrillation threshold was 800 ± 200 V in this series. Colors represent phase, and the symbols + and − indicate phase singularities or centers of reentrant circuits of opposite direction. Phase singularities as a marker of reentry were observed during ventricular fibrillation just before the shock in all cases. A and B, Postshock phase singularities were observed after failed 100-V and 200-V shocks. Visual analysis of animations of the optical recordings indicated that many of the phase singularities represented reentrant activations occurring immediately after the shock, so that the postshock interval was 0 msec. C and D, No phase singularities were observed after the 600-V and 800-V shocks. For the 600-V shock, activation propagated away in all directions from two early sites, at the apex and at the lateral base of the left ventricle, both of which appeared after a postshock interval of 42 msec. For the 800-V shock, a single wavefront of activation appeared at the apex and propagated away in all directions in a focal pattern after a postshock interval of 72 msec.

(From Chattipakorn N, Banville I, Gray RA, Ideker RE: Effects of shock strengths on ventricular defibrillation failure. Cardiovasc Res 61:39-44, 2004.)

At least two inferences can be drawn from the model that the lowest potential gradient region determines whether or not a shock will halt fibrillation. First, different shock electrode configurations with the same defibrillation threshold (DFT) have the same lowest potential gradient value, but in different regions of the heart. Single-coil versus dual-coil defibrillator systems have similar DFTs. Mapping of postshock activations shows that the earliest recorded activity after the shock is the anterior left ventricle for dual-coil shocks but the posterior area for single-coil shocks.⁷² Thus the low gradient region presumably is anterior for dual-coil shocks and posterior for single-coil shocks, although this inference needs to be confirmed experimentally. Crossley et al.⁷³ measured DFTs for shocks from either an electrode in the right ventricular (RV) apex or RV outflow tract to the pulse generator can and showed that DFTs were not significantly different for the two configurations. We can infer that the lowest potential gradient level was the same for each configuration, but likely in a different location.

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