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1 Department of Cardiology, Cleveland Clinic Foundation, Cleveland, Ohio 44195; and 2 Department of Biomedical Engineering, Tulane University, New Orleans, Louisiana 70112
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ABSTRACT |
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 TOP
 ABSTRACT
 INTRODUCTION
METHODS
RESULTS
DISCUSSION
REFERENCES
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We recently suggested that failure of implantable defibrillation therapy may be explained by the virtual electrode-induced phase singularity mechanism. The goal of this study was to identify possible mechanisms of vulnerability and defibrillation by externally applied shocks in vitro. We used bidomain simulations of realistic rabbit heart fibrous geometry to predict the passive polarization throughout the heart induced by external shocks. We also used optical mapping to assess anterior epicardium electrical activity during shocks in Langendorff-perfused rabbit hearts (n = 7). Monophasic shocks of either polarity (10-260 V, 8 ms, 150 µF) were applied during the T wave from a pair of mesh electrodes. Postshock epicardial virtual electrode polarization was observed after all 162 applied shocks, with positive polarization facing the cathode and negative polarization facing the anode, as predicted by the bidomain simulations. During arrhythmogenesis, a new wave front was induced at the boundary between the two regions near the apex but not at the base. It spread across the negatively polarized area toward the base of the heart and reentered on the other side while simultaneously spreading into the depth of the wall. Thus a scroll wave with a ribbon-shaped filament was formed during external shock-induced arrhythmia. Fluorescent imaging and passive bidomain simulations demonstrated that virtual electrode polarization-induced scroll waves underlie mechanisms of shock-induced vulnerability and failure of external defibrillation.
sudden cardiac death; ventricular fibrillation; external shock; optical mapping; bidomain simulations
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INTRODUCTION |
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TOP
ABSTRACT
INTRODUCTION
METHODS
RESULTS
DISCUSSION
REFERENCES
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THE ABILITY OF ELECTRIC STIMULI to induce cardiac arrest presumably due to ventricular fibrillation was established by Hoffa and Ludwig (19) 150 years ago. On the other hand, Prevost and Battelli (23) demonstrated that ventricular fibrillation can be terminated by an electric discharge delivered directly into the heart. Further development of this idea carried out by Mirowski et al. (21) and Schuder et al. (27) resulted in a new therapy, implantable cardioverter defibrillator (ICD), which has been recognized as one of the most effective means against sudden cardiac death (1). Unfortunately, the cost of the ICD therapy remains a limiting factor of its wider application. Alternatively, external defibrillation (17, 37) has become a common therapy in literally every emergency room in the developed world. Recent development of semiautomatic external defibrillators, which do not require the presence of a trained health professional, may significantly extend the area of application of the therapy, including public places and private homes. This potential wider application raises additional concerns regarding the safety and optimization of external defibrillation therapy. However, these issues cannot be addressed without a better understanding of the mechanisms of external defibrillation.
We recently suggested a mechanism that might be responsible for the failure of implantable defibrillation therapy (6). Our hypothesis is based on the finding that ICD shocks produce areas of positive and negative polarization of various amplitudes next to each other, known as virtual electrode polarizations (VEP) (28, 30). A new wave front may be formed in a region where strong positive and negative polarizations meet. Such wave fronts have been shown to rapidly excite negatively polarized areas after the shock withdrawal, thus eliminating all shock-induced excitable regions and, ultimately, completing successful defibrillation (2). On the other hand, areas where both strong polarizations are adjacent to an area of no polarization meet the criteria of a phase singularity (6) and are responsible for the formation of reentrant circuits. Thus the virtual electrode-induced phase singularity may result in a new arrhythmia and, consequently, in failure of defibrillation therapy.
In this report we suggest that VEP is a common mechanism by which the shock affects cardiac tissue and thus may also be responsible for the success and failure of external defibrillation therapy. We chose two complementary research approaches to elucidate the mechanisms of vulnerability to external shocks. We used a bidomain simulation approach, which has been critically important in predicting the VEP effect during shocks (24, 25, 32). It has the ability to provide insights into the shock-induced transmembrane polarization in the myocardium. Because of issues of computational tractability, only the passive version of the bidomain model can be used in a three-dimensional geometry. It, however, lacks the power to predict the postshock response and requires experimental confirmation. Fluorescent imaging of defibrillation has proven to be the only technique capable of faithfully recording the electrical activity during defibrillation shocks (4, 7). However, this method lacks the ability to map voltage in three dimensions. Thus the combination of the two methodologies provides a unique array of tools capable of predicting the three-dimensional voltage distribution and the postshock active response.
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METHODS |
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TOP
ABSTRACT
INTRODUCTION
METHODS
RESULTS
DISCUSSION
REFERENCES
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Computational methods. The finite-element method was employed to model external defibrillation. A uniform field shock was delivered via a conductive bath to an anatomically precise representation of the rabbit ventricles, including the fibrous structure. The myocardium was modeled as a bidomain, and the transmembrane potential (Vm) distribution induced by the shock was calculated. The combination of realistic fiber architecture, realistic geometry, and the bidomain model make possible the accurate prediction of the location and shape of virtual electrodes induced by electric field shocks.
Bidomain model.
The bidomain model is commonly used (18, 33) to accurately
reproduce the electrical activity of excitable tissue. It is a system
of two reaction-diffusion equations, one for the extracellular space
and the other for the intracellular space, coupled by the transmembrane
current. For computational tractability and because the initial
response of myocardial tissue to external shocks is predominantly the
formation of shock-induced virtual electrodes, assuming a passive
membrane behavior is a good approximation (18, 32). Thus
the membrane is modeled as a parallel combination of a conductance and
a capacitance. Furthermore, because we are interested in the
Vm distribution induced by a uniform electric field shock and not in the evolution of such a distribution, the time
dependence of the transmembrane current was eliminated, resulting in a
significant computational acceleration. The model was thus simplified
to a system of two differential equations coupled by a
constant-resistance membrane
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(1) |
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(2) |
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(3) |
i, e, and
Vm are the intracellular, extracellular, and
transmembrane potentials, respectively. The fiber architecture is
incorporated into the model via the global intracellular and extracellular conductivity tensors, 
i and
e (3)
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(4) |
· T is
the outer product of the unit vector parallel to the local fiber
direction with itself, I is the 3 × 3 identity matrix,
and 
i,e represents the longitudinal (l) or transverse
(t) component of the local intracellular and extracellular
conductivities of a cardiac fiber.
The bath inside the ventricular cavities and outside the heart
satisfies Laplace's equation. At the interface between myocardium and
bath, the intracellular current satisfies a no-flux condition and the
extracellular current satisfies a conservation-of-flux condition.
All parameters used in the simulations and their values are listed in
Table 1.
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Geometry and finite-element grid.
Because the shape of virtual electrodes is greatly affected by tissue
geometry and fibrous architecture, it was critical to use anatomically
accurate geometry and measured fibrous architecture. The rabbit
ventricle geometry and fiber structure were provided by Vetter and
McCulloch (34) as a regularly sampled set of data. A
regular mesh was constructed using these data. Model dimensions, space
constants, and finite-element mesh statistics are listed in Table 1.
The finite-element mesh discretization was smaller than the smallest
length constant. Simulated shock was applied from a pair of flat
electrodes, which reproduced the conditions in the rabbit experiments.
Figure 1 shows the anterior and posterior view of the model.
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Experimental methods.
Experiments were performed in vitro on Langendorff-perfused rabbit
hearts (n = 7; Fig.
2A). Detailed protocols have
been published (5, 7). Isolated rabbit hearts were removed
and placed onto a Langendorff apparatus, where the hearts were perfused
with modified Tyrode solution containing 15 mM 2,3-butanedione monoxime
(BDM; Sigma Chemical), the excitation-contraction uncoupler. In two experiments, shocks were applied without and with BDM. The temperature and pH were maintained at 36 ± 0.5°C and 7.35 ± 0.05, respectively.
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Experimental protocol and data analysis. A previously described (6, 7) optical mapping system schematically shown in Fig. 2B was used in our experiments. The heart was stained with the voltage-sensitive dye di-4-ANEPPS, as previously described (7). Fluorescence was excited at 520 ± 45 nm and collected at >610 nm by the 16 × 16 photodiode array (model C4675, Hamamatsu). The magnification was adjusted to focus on an area from 0.5 × 0.5 to 1.0 × 1.0 mm per diode. The entire field of view varied between 8 × 8 and 16 × 16 mm. After amplification, the signals were sampled at 1,894 frames/s. Each frame included 256 optical channels and 8 instrumentation channels.
The heart was continuously paced at a cycle length of 300 ms. The pacing stimulus strength was adjusted to twice the diastolic threshold of excitation. Shocks (150 µF, 8 ms) were applied during the T wave from a clinical defibrillator (model VHS-02, Ventritex). A total of 162 shocks was delivered with an average of 23.1 ± 16.2 shocks/heart. The stimulator was paused for 2 s after the shock. Each scan contained 1.5-2 s of data, including the last basic beat action potential, the action potential altered by a shock, and two or more subsequent action potentials (see Fig. 5D). The signal analysis software used in this study was described previously (6, 7). This program automatically calculated the maps of activation (26), repolarization (9), action potential duration (9), and calibrated Vm (11). Maps of activation and repolarization were built using (dVm/dt)max and (d2Vm/dt2)max methods, respectively. Maps of Vm were calculated assuming a normal resting potential of
85 mV, and an
action potential amplitude of 100 mV was present at all recording
sites. This approach is similar to a technique used in the atrium
(16). Contour maps were automatically built using Origin
5.0 (Microcal Software).
We estimated the upper and lower limits of vulnerability (ULV and LLV)
in five hearts. Accurate measurements of these two parameters were not
conducted to reduce the number of shocks applied to the same heart.
Estimation was done by identifying arrhythmic responses to shocks of
various strengths. A response was considered arrhythmic if at least one
extra beat was induced. The lowest shock voltage at which the
arrhythmia was induced was considered to be the LLV; the highest shock
voltage was the ULV. The two parameters were averaged for both
polarities. We recognize that such an estimate is likely to
overestimate ULV and underestimate LLV. Despite this limitation, we
could correlate the two parameters with observed VEP and postshock
response, allowing us to address vulnerability to arrhythmias.
Values are means ± SD. Comparison between variables was analyzed
by Student's two-tailed t-test for paired and unpaired
samples. Differences were considered significant if P < 0.05.
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RESULTS |
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TOP
ABSTRACT
INTRODUCTION
METHODS
RESULTS
DISCUSSION
REFERENCES
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Epicardial VEP in the bidomain model.
Figure 3 shows a steady-state epicardial
distribution of shock-induced Vm in the passive
bidomain model. Preshock Vm was set at 0 mV.
Figure 3 shows that the RV epicardium was depolarized by the shock,
whereas the left ventricular (LV) epicardium was hyperpolarized by the
shock. As seen from the width of white and lightly colored areas, the
Vm gradient between the two polarizations was
not uniform throughout the epicardium. The apical view reveals that the
gradient between positive and negative polarization is the steepest at
the apex and decreases toward the base. We recently showed that the
amplitude of the VEP gradient is the main predictor of sites of origin
of a VEP-induced wave front (2). This observation suggests
that the wave front is more likely to originate at the apex than at the
base. Therefore, conditions for virtual electrode-induced phase
singularity might be met and reentry might ensue (6).
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Epicardial VEP in the rabbit heart.
Figure 4 shows transmembrane polarization
measured in one experiment. Subsequent Figures (5-10) show
similar observations from three more experiments.
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Virtual electrode-induced phase singularity.
We estimated LLV and ULV in five hearts, which were 22.0 ± 21.7 V
and 161.0 ± 25.1 V, respectively. Table
2 summarizes these estimates. To identify
the mechanisms of shock-induced arrhythmia, we analyzed only responses
to shocks within the vulnerability limits.
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Site of origin of the wave front of reexcitation. On the basis of our observation that the wave fronts originated from the apical part of the field of view, we hypothesized that the earliest activation can be seen directly at the apex. To prove this hypothesis, we mapped the electrical activity directly at the apex by 1) moving our field of view down and mapping anterior apical epicardium and 2) turning the heart to a horizontal position with the apex facing our mapping system while keeping the same heart orientation with respect to the shocking electrodes.
Figure 8 shows an example of mapping at the apex of the anterior epicardium. The heart is different from that used in Fig. 4. A monophasic shock (100 V, 8 ms) was delivered by the cathode facing the RV and the anode facing the LV (Fig. 8A). As seen in the raw traces shown in Fig. 8C, the shock depolarized the RV epicardium and negatively polarized the LV epicardium. Figure 8B shows a 10-ms isochrone map of activation. The wave front of excitation originated at the apex of the LV and spread toward the base, then turned around a pivoting point and invaded the RV epicardium. Thus a sustained ventricular tachycardia was induced.
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Three-dimensional pattern of VEP predicted by the bidomain model.
Because of the absorption of the excitation and emission light,
fluorescent imaging is limited in its ability to assess midmyocardial electrical activity. Therefore, we used the bidomain model to predict
the polarization throughout the entire heart, including the right and
left midmyocardium, the endocardium, and the septum. Figure
11 shows the results. As evident from
the ventricular cross sections presented in Fig. 11, VEP was produced
throughout the entire heart in a complex fashion. The exact pattern is
strongly influenced by the orientation of the heart with respect to
electrodes (not shown). Yet, several features of VEP are common to any
external stimulation with homogeneous electric field. Every surface of the myocardium is positively polarized if it faces the cathode and
negatively polarized if it faces the anode. The surface polarizations are stronger in amplitude than any bulk polarization. As previously shown by Wiedmann (35) and Trayanova (31),
such surface polarization decays exponentially with distance from the
surface. In some areas, the polarization extends deeper than the
surface myocardial layers because of the fiber curvature effect
(29). Furthermore, the gradient between the positive and
the negative polarizations is distributed unevenly, with strong
gradient in some areas (e.g., RV free wall in slice 4) and
weak gradients in others (e.g., LV free wall in slice 4).
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Evidence of a three-dimensional scroll wave produced by VEP. As shown in Figs. 5-9, the line of steep gradient between the positive and negative polarizations at the epicardium is the line of block of the induced vortex. Similarly, there is a theoretically predicted surface between the opposite-in-sign epicardial and bulk polarizations that may serve as a filament of a scroll wave in three dimensions. However, such a surface may or may not be visible to direct epicardial mapping, depending on its depth. If such a filament is located within 1-2 mm from the epicardium, it may be detected by the optical system as "dual-humped" signals (8, 10, 11).
The data already presented here support such a hypothesis. Indeed, careful examination of Fig. 6 reveals that nearly all signals in the upper right corner show typical "dual-humped" morphology carrying the signature of a deeper wave front (8, 10). Similar dual-humped signals are evident in Fig. 8C. Figure 12 provides yet another example of dual-humped signals. In this case, we chose a reentry produced by a 120-V shock in the same heart as in Figs. 5 and 6. Signals shown in Fig. 12C demonstrate a strong dual-humped morphology. The epicardial activation map (Fig. 12A) was reconstructed using only the largest peaks of (dV/dt)max. It shows that a wave of excitation was generated at the apex within the first 10 ms after shock withdrawal, which occurred at 520 ms. As illustrated in Fig. 12C, this wave front propagated upward. After reaching the upper boundary of the field of view, the wave turned around in a fashion similar to that in Fig. 5C but at a higher vertical location in the field of view, invaded the already recovered RV epicardium, and spread toward the apex (Fig. 12, A and B). At the same time, the LV signals show second components (Fig. 12C), which were ignored during the construction of the upper map. These were used to construct the map of activation at the endocardium or midmyocardium. After its arrival at the top boundary of the field of view, the wave front made a turn to the left. Figure 12B shows that the recording sites were sequentially activated. Simultaneously, another wave front originated there and spread backward toward the apex along the right side of the field of view (Fig. 12C). As evident from the endocardial or midmyocardial recordings, there was an uninterrupted wave front propagating from the base to the apex. Such a peculiar activation pattern cannot be explained within the two-dimensional paradigm. Indeed, how is it possible that the same sites in the upper right corner of Fig. 12A were reactivated within 30-40 ms if normal action potential duration in this area was 175 ms? Only a three-dimensional scroll wave of the type described above can easily explain such propagation. Figure 12A illustrates our reconstruction of the scroll wave and its ribbon-shaped filament. At first, the wave shown with red isochrones propagated within the space limited by the epicardial surface and the filament. After recovery of the adjacent tissue, the scroll wave invaded it below and to the left of the filament.
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DISCUSSION |
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TOP
ABSTRACT
INTRODUCTION
METHODS
RESULTS
DISCUSSION
REFERENCES
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This study presents theoretical and in vitro experimental data obtained during externally applied electric shocks. The data show that, similar to internally applied shocks investigated in our previous studies (2, 6, 7, 11), external shocks evoke a VEP pattern. It provides the basis for a virtual electrode-induced phase singularity and the resulting reentrant scroll wave, which underlie shock-induced arrhythmogenesis. Yet there are important differences between these two cases. The pattern of epicardial polarization produced by externally applied shocks is different from that induced by internal shocks. Only two areas of opposite polarization are present on the epicardium: negative facing the cathode and positive facing the anode. As a result, only one wave front is induced; it begins at the epicardium and ends at the endocardium while propagating through the apex of the heart. This wave front has two wave breaks, which could result in two reentrant circuits: one at the anterior and another at the posterior epicardium. This is different from the effect of internal shocks, which could potentially induce four reentrant circuits (quatrefoil reentry) (20).
Recent progress in theoretical and experimental approaches to defibrillation research has resulted in formulation of the VEP theory (32). A growing body of evidence suggests that VEP is perhaps the most important component of the interaction between externally applied electric field and heterogeneous myocardium. A number of structural heterogeneities of different spatial scales have been considered as a substrate of shock-induced stimulation and defibrillation. These heterogeneities include, in order of increasing spatial scale, cell-to-cell junctions (22), syncitial heterogeneities (14, 15), unequal anisotropy between intra- and extracellular domains (28), tissue-bath interface (12, 31, 35), and fiber curvature (29). In addition, heterogeneity of the external field itself may contribute to VEP (28). The larger the spatial heterogeneity in the external field, the stronger the shock-induced polarization (32). VEP described by Sepulveda et al. (28) arise around small-sized electrodes that generate strongly nonuniform fields. Such VEP has been observed during ICD shocks and may play an important role in internal defibrillation (6, 12). However, external defibrillation is clearly driven by different VEP mechanism(s). It is possible that fiber curvature and tissue-bath interface play the major role in this type of defibrillation mostly because of their spatial scale. Our data provide the first experimental and theoretical evidence supporting this prediction (13, 32).
An earlier study by Zhou et al. (36) did not present any evidence of VEP during externally applied shocks. Only a single transmembrane polarization polarity along the line connecting the two opposite electrodes was observed during any given shock polarity. This observation contradicts our experimental and theoretical findings, according to which positive and negative polarizations are present during shocks of any polarity. Comparison of our results with those of Zhou et al. is difficult because of the differences in recording methodology. Zhou et al. recorded electrical activity from only nine posterior epicardium spots near the base, whereas we imaged nearly the entire anterior epicardium. Most importantly, Zhou et al. used a bipolar electrode to estimate extracellular voltage gradient "immediately adjacent to each laser recording spot as shocks were given" (36). Such an electrode pair and a piece of silicone rubber that held it may have altered the transmembrane polarization because of tissue-bath interface or secondary sources near the recording stainless steel electrodes. As in our previous experiments (12), we verified such a possibility by mapping when the heart was touching the glass window and when it was placed at a distance from it. Contact with the glass window somewhat reduced but did not eliminate the opposite polarization observed without the contact.
Our data show that the steep gradients between oppositely polarized areas are sites of wave front origination. Careful three-dimensional mapping of such regions may help identify these sites and, most importantly, the locations of the phase singularities. As is evident from the fluorescent imaging data, the bidomain simulation provided accurate prediction of such sites at the epicardium. Indirect evidence of scroll waves developed after the shock supports the endocardial and midmyocardial distribution of shock-induced Vm, as predicted by the bidomain model. Thus it appears feasible in the future to be able to theoretically predict the areas of potential phase singularities on the basis of the specific ventricular geometry and defibrillation lead configuration. Yet, careful mapping of three-dimensional VEP is required to fully support the theory. This is especially important in hearts with structural disease. Areas of infarct, fibroses, or ischemia would change VEP and might provide additional substrate for wave front initiation and phase singularities.
Limitations. Our study is limited because of the inability to experimentally assess electrical activity in the three-dimensional myocardium. Furthermore, stand-alone passive bidomain simulations have limited predictive power because of the lack of representation of the ionic currents in the computer model. Yet, the combination of optical imaging of the epicardial surface of the ventricles with three-dimensional passive bidomain model simulations provides guidance as to what the deep myocardial activity could be, as well as an assurance of the correct interpretation of our experimental findings.
The passive bidomain simulations predict only the shock-induced Vm changes throughout the three-dimensional myocardium and not the postshock activity. Although highly desirable, inclusion of ionic membrane kinetics in the rabbit heart model remains an insurmountable task. First, the ionic model would increase the memory requirement by over an order of magnitude. In addition, we estimate that to obtain 200-ms of data, the CPU time requirements would increase four to seven orders of magnitude. Before we are able to approach a problem of such magnitude, more efficient numerical algorithms need to be implemented. Our study is also limited because of the use of BDM as excitation-contraction uncoupler. Such treatment may have an effect on ionic channel conductance and arrhythmogenesis. We verified the extent of this limitation by mapping Vm during external shocks in two hearts with and without BDM. These measurements showed that positive and negative polarizations are present in both cases. Unfortunately, strong movement artifacts did not allow reconstructing the pattern of activation during arrhythmogenesis.| |
ACKNOWLEDGEMENTS |
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The authors thank Dr. McCulloch (University of California, San Diego) and his group for providing the rabbit heart fiber orientation and geometry and Dr. Eason (University of Vermont) for invaluable help with the simulations.
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FOOTNOTES |
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This study was supported by National Heart, Lung, and Blood Institute Grants R01-HL-58808 (I. R. Efimov), R01-HL-59464 (I. R. Efimov), and R01-HL-63195 (N. Trayanova), National Science Foundation Grants DMF-9709754 (N. Trayanova) and BES-9809132 (N. Trayanova), and American Heart Association Ohio Valley Affiliate Grant-in-Aid 9806201 (I. R. Efimov).
Address for reprint requests and other correspondence: I. R. Efimov, Biomedical Engineering, Case Western Reserve University, 10900 Euclid Ave., Cleveland, OH 44106 (E-mail: ire{at}po.cwru.edu).
The costs of publication of this article were defrayed in part by the payment of page charges. The article must therefore be hereby marked "advertisement" in accordance with 18 U.S.C. Section 1734 solely to indicate this fact.
Received 22 December 1999; accepted in final form 6 March 2000.
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TOP
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RESULTS
DISCUSSION
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Vm) in C and D were
calculated by subtracting the normal response from the shock-induced
response. Voltage was calibrated on the basis of the assumption that
normal action potential amplitude is 100 mV. C:
shock-induced polarization profiles recorded along the middle
horizontal row of recording sites at the last sample (528 µs) of the
shock. Different colors represent data recorded during shocks of
different intensity (±50 to ±250 V). D: maps of
polarization produced by shocks of different polarity and intensity.
