On the electrocardiogram, ventricular fibrillation presents as irregular and random waveforms with no clearly identifiable QRS complexes or P waves. Therefore, the electrocardiogram cannot be used for understanding the underlying mechanism of fibrillation. To study the re-entrant activation patterns, underlying local ventricular activation needs to be recorded. In the explanted perfused heart, multi-electrode grids enable the simultaneous recording of local extracellular electrograms at multiple locations. In a clean unipolar local electrogram the minimal dV/dt of the RS complex corresponds to the local time of myocardial activation under the electrode (Spach et al. 1972; Dower 1962). From these local times of activation a map can be generated, which reveals the activation sequence and the possible re-entrant pathway as shown in Fig. 14.2. Multi-electrode grids can be placed on the endocardium and the epicardium in order to simultaneously measure the activation pattern on both sides. Needle electrodes can be use to record the transmural activation patterns (Pogwizd and Corr 1987).

In the clinic, catheter mapping is used to record activation times (Shenassa et al. 2013). For monomorphic arrhythmias, a single roving catheter can be used, which will be sequentially placed on numerous sites (for instance the CARTO system (Biosense-Webster, Baldwin Park, CA, USA)). The position of the electrode is monitored using a heterogeneous magnetic field of which the strength can be detected by a sensor at the tip of the catheter. This enables the 3D reconstruction of the activation pattern. However, this technique is not suitable for mapping irregular arrhythmias because it is based on sequential recordings. Newer CARTO mapping systems use catheters with multiple electrodes and begin to circumvent this problem. Another often-used system is EnSite (St. Jude Medical, Inc., St. Paul, MN, USA), which is based on multi-electrode non-contact mapping. With this method the inverse solution is used to calculate the local moments of activation on the endocardium (Taccardi et al. 1987). The introduction of catheter non-contact and contact mapping has improved diagnosis and led to treatment of arrhythmias by ablation or antitachycardia pacing (Josephson et al. 1978; Durrer et al. 1967). However, electrical mapping techniques are not suitable for recording the tissue response directly after defibrillation because the overwhelming shock-induced artifacts mask the signal.

Electrical mapping has been a useful tool for studying the mechanisms of arrhythmia origin and perpetuation and even led to a variety of therapies. However, there are shortcomings of electrical mapping. The spatial resolution is limited and it is not possible to measure local electrical activity during or after a defibrillating shock (for reasons that will be discussed in detail later). Optical mapping provides an alternative method for measuring cardiac electrical activity and addresses the aforementioned shortcomings of electrical mapping (Boukens and Efimov 2014). Since optical mapping requires the use of fluorescent dyes the method is currently restricted to the experimental ex vivo settings.

Cardiac electrical activity can be transduced into fluorescent signals by using voltage sensitive dyes that bind to the membrane of cardiomyocytes. When a dye molecule binds to the membrane, emitted fluorescence can be recorded upon excitation of the dye. The wavelength of the emitted light depends on the membrane potential of the cardiomyocyte. When the membrane potential changes (i.e. during the action potential), the dynamic wavelength of the emitted light can be recorded as a loss in fluorescence with appropriate filter settings on the acquisition hardware. Filters are specific for particular dyes (e.g. excitation 520 ± 20 and emission >650 nm for di-4-ANEPPS). The change in emitted light is linearly related to the membrane potential, which enables the recording of “optical action potentials”. This mechanism is illustrated in Fig. 14.3. The number of myocytes that give rise to an optical action potential signal depends on the type of the dye, the optical magnification, the focal plane and the pixel size of the sensor that is used for recording. Optical action potentials are not only derived from fluorescence collected from the surface of tissue but also of fluorescence from deeper layers (Bishop et al. 2007). Therefore, the optical action potential is an underestimation of the transmembrane potential on the surface and should be compared with transmembrane potential averaged over depth (Janks and Roth 2002). Commonly used sensors for recording optical action potentials are photodiode arrays, CCD and CMOS cameras (Efimov and Salama 2012).

Although, optical mapping is often the method of choice in experimental cardiology there are some disadvantages of this technique that need to be considered. Excitation of the fluorescent dye has been shown to produce toxic free radicals and may cause death of cardiomyoctyes. This side effect is most prominent when single myocytes or monolayers are measured. Also, the excitation of deeper located layers cause light scattering and distortion of the optical signal (Bishop et al. 2007; Ramshesh and Knisley 2003). Another disadvantage is that contractions of the heart lead to changes in emission intensity related to motion artifacts not the membrane potential. Therefore, in order to accurately measure the shape of optical action potentials, the excitation and contraction must be pharmacologically uncoupled. A common uncoupler agent is blebbistatin (Fedorov et al. 2007), but cytochalasin D and 2,3-butane-dion monoxime have also been used.

Although optical mapping and, to a lesser extent, conventional multi-electrode grids, increases spatial resolution, neither can track conduction when the activation front propagates outside the limited field of view. This hampers the reconstruction, and thereby interpretation, of the activation sequence, especially during complex re-entrant patterns. Panoramic imaging refers to optical imaging using multiple cameras or mirrors to measure from at least three different angles. This technique allows reconstruction of complete epicardial activation patterns (Kay et al. 2004) and has given insight into complex activation patterns during ventricular fibrillation. Figure 14.4 is an example of a re-entrant activation pattern of a rabbit heart using panoramic optical mapping (Ripplinger et al. 2009; Rogers et al. 2007).

Another problem with conventional optical mapping is that a curved epicardial surface is projected onto a 2-dimensional sensor. This means that distance in the “Z” direction is not taken into account. This is a limitation, when conduction velocity is calculated, because it leads to underestimation of conduction velocity, especially near the edges of the field of view. Panoramic imaging enables the projection of a 3-dimensional conduction pattern on the surface of the heart (Lou et al. 2012). Conduction velocity can be more precisely measured thereby improve our understanding of conduction during normal and re-entrant activation patterns.

As mentioned above the optics of a set-up can be fined tuned to specific applications, providing control over the spatial resolution and depth of the recorded signals. The fine spatial resolution of optical mapping makes it an ideal platform for tracking the dynamics of ventricular fibrillation, pinpointing focal discharges and the critical points of reentrant wavefronts as well as the critical points that are created when energy is applied during electrotherapy. Chapter 13 (Gray et al.) outlines how optical mapping can be used to study the mechanisms that maintain ventricular fibrillation including phase singularity tracking.

Programmed electrical stimulation is essential for studying electrophysiology of the intact myocardium. Stimulation is established by generating a potential difference between two electrodes, causing electrons to flow between the negative (cathode) and the positive (anode) electrode. If the two electrodes both make contact with the cardiac tissue and are in close proximity, this is called bipolar stimulation. If one electrode makes contact with cardiac tissue and the other (indifferent) electrode makes contact elsewhere, ground or the bath solution, it is called unipolar stimulation. When a unipolar cathode is used for stimulation then the tissue under the electrodes depolarizes, which leads to excitation. Adjacent virtual anodes hyperpolarize tissue and transiently slow activation in lateral directions. Anodal stimulation causes hyperpolarization under the electrode and depolarization under two virtual cathodes, which stimulate the tissue at the lateral borders of the hyperpolarized tissue. The current flow during anodal and cathodal stimulation will be discussed in more detail in the Sect. 5.3.

The method of stimulation may influence stimulation threshold or refractoriness of the tissue. For instance, the repolarizing effect of anodal stimulation may shorten action potential duration and lead to short refractory periods or even completely de-excite the tissue restoring full excitability (Mehra et al. 1977). During bipolar stimulation refractory periods are shorter than during cathodal stimulation as well due to anodal stimulation at the proximal electrode (Mehra and Furman 1979). However, the incidence of inducing arrhythmias is not different between bipolar or unipolar stimulation, (Stevenson et al. 1986) unless anodal and cathodal shocks are considered, which do have significant differences in arrhythmogenic potential.

During cathodal stimulation the cardiac tissue right under the electrode depolarizes. However, the myocardium adjacent to the depolarized myocardium hyperpolarizes due to virtual anodes. Thus, the “real” cathode generates two “virtual” anodes. This phenomenon is referred to as “virtual electrodes” (Sepulveda et al. 1989; Furman et al. 1975). To avoid confusion in the nomenclature, often investigators refer to all areas of positive and negative polarizations as “virtual electrodes”, regardless of their proximity to real stimulation electrodes. We will follow this convention. Figure 14.6a shows a computed (passive bidomain model) potential field during central anodal stimulation in a two-dimensional cardiac tissue sheet. The center is highly hyperpolarized (red) whereas the lateral adjacent sides are highly depolarized (blue). In this example the fiber direction is diagonal. The depolarized field expands up and down and shows a “dog-bone-like” morphology. The morphology of the virtual electrode depends on the stimulus strength, fiber direction, and both extra and intracellular resistance (Wikswo et al. 1991). A higher stimulus current causes an increase in the amplitude of positive and negative polarization of corresponding regions. Inversion of the current (i.e. cathodal stimulation) leads to the inverse potential field and virtual electrode pattern. Bipolar stimulation generates virtual electrodes as well (Nikolski and Efimov 2000). The number of virtual electrodes that is generated depends on the location and spacing of the bipole with respect to the fiber direction under the electrodes. When the anode and the cathode are positioned perpendicular to the fiber direction a total of six virtual electrodes will be generated (Fig. 14.6b). If the electrodes are positioned in parallel with the fiber direction a total of four virtual electrodes are observed (Fig. 14.6c) (Nikolski et al. 2002). For comparison, the unipolar stimulation produces a total of three virtual electrodes.

Optical imaging enables experimental testing of the virtual electrode hypothesis in ex vivo heart preparations. Figure 14.7 shows the transmembrane potential during stimulation of rabbit epicardium. The measured potential field closely resembles the expected morphology based on the computer simulation.

During electrical stimulation not only the “real” electrode can cause excitation but the simultaneously generated surrounding virtual electrodes can produce wavefronts as well. The wavefront will propagate from the virtual cathode to the virtual anode. The direction and lifetime of the induced wavefront is directly dependent of the specific VEP pattern. In this section we will only discuss unipolar cathodal and anodal stimulation, because it is directly relevant for understanding the mechanism of defibrillation, which will be reviewed below. There are four mechanisms of unipolar electrical stimulation: cathode make, anode make, cathode break and anode break (Dekker 1970). The “make” excitation occurs just after the beginning of the stimulus pulse, whereas “break” excitation happens after the end of the pulse.

Cathodal make stimulation resembles normal activation of the cardiac tissue during direct injection of depolarizing current via a microelectrode. The negatively charged real cathode causes depolarization of the myocardium right under the electrode. This depolarization causes the membrane to reach excitation threshold and initiates an activation front. Depending on the stimulus current the activation front will start directly under the electrode or from a point farther from the electrode. Nearby negative virtual anodes will slow the conduction but ultimately will be depolarized by the actively propagating wavefront.

The mechanism of excitation during anodal make stimulation resembles that of cathode make stimulation. However, the positively charged real anode causes hyperpolarization of the myocardium right under the electrode. This means that excitation starts at the lateral sides near the virtual cathode. The amplitude of depolarization near virtual cathode during anodal stimulation is lower than near real cathode during cathodal stimulation. Therefore, the threshold of excitation is lower during cathode make than for anode make stimulation.

As mentioned above, break excitation happens at the end or break of the stimulus. Cathode break stimulation occurs when the myocardial tissue directly under the electrode is refractory and unexcitable. However, the shock does still de-excite the tissue at the virtual anodes, rapidly repolarizing the membrane potential even if the tissue was refractory. At the end of the stimulus the current diffuses to the hyperpolarized tissue, exiting to the myocardium through the virtual anodes. This leads to excitation. The path that is created by the shock is parallel with the fiber direction, and the propagation direction into excitable tissue is different than cathode make excitation. The virtual anode is usually much weaker than the real cathode and therefore cathode make excitation will occur over cathode break excitation.

During anodal break excitation the hyper excitable path is perpendicular to the fiber direction. Thus, after the stimulus pulse, the tissue at the virtual cathode repolarizes slowly allowing current to flow into the tissue directly under the anode, which was rapidly de-excited during the shock. This will lead to excitation and a wave front that will propagate transversely before conducting into the normal myocardium.