Local and Global Effects of Cardiac Electrical Stimulation

Local.

Bradycardia pacing requires that the stimulus capture fully excitable local myocardium during diastole. The stimulated wavefront then propagates to most or all of the myocardium, which is also fully excitable, thereby resulting in electrical depolarization of cells and resultant mechanical contraction. In contrast, ATP stimuli must interact with the specific reentrant circuit driving the tachycardia, which is generally remote from the site of pacing, and it must do so while most of the myocardium is refractory or relatively refractory. Thus the ATP stimulus must capture local myocardium during the relative refractory period, propagate to the reentry circuit through relatively refractory myocardium, enter the circuit during an excitable gap in refractoriness, and terminate the tachycardia by causing a bidirectional block (see Chapter 33). Stimulus strength for local capture by ATP is higher than that for bradycardia pacing because ATP pulses are usually delivered to myocardium that is relatively refractory rather than fully excitable.

Principles of Bioelectrical Stimulation

Thresholds for Pacing and Defibrillation.

A threshold stimulus is the minimum stimulus required to evoke a response. Stimuli weaker than the threshold never evoke a response, and stimuli stronger than the threshold always evoke a response. Thus the threshold for pacing is the minimum stimulus strength needed to depolarize local myocardium and to initiate a propagated response. Defibrillation is best described by a probability-of-success curve (Fig. 36-1A) rather than by a threshold.¹ Shock strength is plotted on the abscissa and probability of successful defibrillation on the ordinate. Because defibrillation is probabilistic, the same clinically relevant shock strength may either succeed or fail on successive attempts. Nevertheless, the term defibrillation threshold (DFT) is used as the minimum shock strength that results in defibrillation during testing. DFT testing refers to various methods that assess the efficacy of defibrillation by calculating a shock strength on the sloping portion of the curve based on successes and failures of a few shocks at different strengths. Thus these methods are based on limited, discrete sampling of a continuous statistical distribution (Fig. 36-1B). Because repeated sampling of a probability distribution is likely to result in variations, repeated measurements of the DFT result in variation in measured values.

Waveforms.

The waveform of an electrical pulse is the temporal pattern of its amplitude, measured by voltage (or current). Voltage is a critical parameter for pacing or defibrillation because it determines the electrical field that interacts with the heart. In general, current is linearly related to voltage by Ohm’s law (V = IR, where V is voltage, I is current, and R is resistance). Waveform duration is critical because a pacing pulse or shock interacts with the heart for the duration of the waveform. Furthermore, the time course of the heart’s response to a pacing or defibrillation pulse depends on time-dependent passive and active ion channel processes, collectively referred to as the membrane time constant (τ_m) of cardiac tissue (see Chapter 33). Thus the most easily measured electrical parameter relevant to pacing or defib­rillation is voltage (or current) as a function of time. Although implantable cardioverter-defibrillator (ICD) shocks are often specified in terms of energy (joules), energy is not a direct determinant of defibrillation.

All pacing and defibrillation in implantable devices result from discharge of a capacitor. Thus they have a fixed leading-edge voltage and a trailing-edge voltage determined by the waveform duration and waveform time constant τ_w, which is defined as the product of the capacitance (C) and electrical resistance (R) of the pacing or defibrillation pathway (electrodes and tissue, τ_w = RC). τ_w is the duration in which the capacitor delivers 86% of its stored energy. Because pacing waveforms are short pulses delivered through high-resistance pathways, they approximate constant-voltage pulses with a lower amplitude, longer duration after potentials of opposite polarity (Fig. e36-1A). Defibrillation pulses are high-voltage, capacitive-discharge, truncated exponential waveforms as shown in Figure e36-1B. Biphasic waveforms defibrillate more efficiently (lower voltage) than monophasic waveforms do (Fig. e36-1C). ICDs deliver “single-capacitor” biphasic waveforms in which the initial voltage of the second phase equals the final voltage of the first phase because they can be generated by reversing the polarity of a single capacitor after the first phase is truncated and then continuing the discharge.

Strength-Duration Relationship.

A plot of the stimulus strength required for pacing or the shock strength required for defibrillation as a function of pulse duration is known as a strength-duration curve (Fig. 36-2). The strength-duration curve can be approximated by an inverse exponential or hyperbolic function. The strength-duration curve is characterized by two parameters. The rheobase is the long-duration asymptote (essentially the lowest value), which is determined by properties of the lead system and electrode-myocardial interface.

The chronaxie is the duration at which the threshold is twice the rheobase amplitude. It may be considered an approximation of an aggregate membrane time constant for the myocardium. Clinically, the chronaxie is important for design of efficient pacemakers and ICDs because it relates to the waveform that paces or defibrillates with the lowest energy, and minimizing the energy required is an important consideration for the longevity and size of pacemaker and ICD generators. A waveform with a duration approximately equal to the chronaxie paces with the lowest energy. Presently, no comprehensive theory permits determining the waveform duration that defibrillates with lowest energy from first principles, but approximations and empiric data relate it to the chronaxie. For capacitive-discharge defib­rillation waveforms, the duration (of the first phase of a biphasic waveform) that defibrillates with minimum energy may be considered intermediate between the optimal duration for response of the cell membrane (chronaxie or τ_m) and the optimal duration for the capacitor to deliver its charge (τ_w). Usually, τ_w exceeds τ_m, so phase 1 of transvenous biphasic waveforms exceeds the chronaxie by 25% to 75%.

Unipolar EGMs are recorded between a small tip electrode in the heart and a large remote (indifferent) electrode, typically the metal housing of the pulse generator device, or the can. The location of the remote electrode has little effect on the cardiac EGM, but it may record noncardiac potentials, such as pectoral myopotentials. Bipolar (true bipolar) EGMs are recorded between the tip and ring electrodes on a lead. Integrated bipolar EGMs are recorded between the tip of a right ventricular defibrillation lead and the large right ventricular coil. When compared with true bipolar electrodes (tip to ring), integrated bipolar EGMs have a wider field of view and thus are more likely to oversense nonphysiologic signals or physiologic signals that do not reflect local myocardial depolarization (Fig. 36-3A). Signals that do not originate in the local myocardium are called far-field signals. They include signals originating in a different cardiac chamber.

The typical amplitude of transvenous atrial and ventricular EGMs is in the range of 1.0 to 5 mV and 5 to 20 mV, respectively. The frequency content of ventricular and atrial EGMs is similar (5 to 50 Hz). T waves have a lower frequency (1 to 10 Hz), whereas most noncardiac myopotentials and electromagnetic interference have higher frequencies. This permits the use of electronic band pass filters to reduce sensing of signals that do not represent myocardial depolarization (oversensing) (Fig. 36-4).

Leads for pacing and defibrillation have in common a cable structure that connects the terminal pins, which are inserted into device receptacles called the “header,” as well as the electrodes used for sensing, pacing, or defibrillation (see Fig. 36-3). Pacing leads may have either bipolar or unipolar structures. With unipolar pacing leads, only one electrode is used for both sensing and pacing; the other electrode in the circuit is the pacemaker generator housing (can) itself. Unipolar pacing at high output may result in pectoral muscle stimulation because energy is dissipated from the pacemaker can. Unipolar sensing is much more likely to result in ventricular oversensing of noncardiac signals such as pectoral muscle myopotentials and electromagnetic interference because the can is part of the sensing circuit and the sensing dipole is large. Bipolar pacing leads have two electrodes between which sensing and pacing occur. Because the sensing dipole is smaller in bipolar sensing circuits than in unipolar circuits, oversensing of noncardiac signals is less common with bipolar sensing. Oversensing of pectoral myopotentials with bipolar sensing often indicates a breach in lead insulation in the pacemaker pocket.

Defibrillation leads may have one or two sensing and pacing electrodes in addition to one or two defibrillation coils. If the defibrillation lead has two electrodes—a tip and a ring electrode—“true bipolar” sensing and pacing occur between these two electrodes. If the defib­rillator lead has only a tip electrode, “integrated bipolar” sensing and pacing occur between the tip electrode and the distal defibrillator coil. The integrated bipolar design simplifies lead design by reducing the number of electrodes in the lead, but it incorporates a defibrillation electrode in the pacing circuit, which increases the likelihood of postshock oversensing or undersensing.

All right ventricular defibrillation leads have a distal defibrillation coil (Fig. 36-3B). Dual-coil leads also have a proximal coil that is usually located in the superior vena cava or high right atrium. Some clinicians prefer single-coil leads because if extraction of the lead is required, proximal coil fibrosis presents the highest risk for serious injury to the superior vena cava. In right-sided implants, the use of dual-coil leads permits removing the can from the defibrillation circuit, which may improve the efficacy of defibrillation. Left pectoral implants are preferred over right pectoral implants for ICDs because the defibrillation vector to the can includes more of the left ventricle.

A subcutaneous ICD system has recently been approved (Fig. e36-2).⁸ This system consists of a single defibrillation coil implanted parallel to the sternum and tunneled to the ICD generator pocket located near the left anterior axillary line. It avoids the implant and intravascular infection risks associated with a transvenous lead, but the energy requirement for defibrillation is considerably higher than that needed for transvenous ICD systems, and it cannot deliver painless endocardial pacing. Thus it is not suitable if bradycardia pacing is indicated; patients’ tolerance of subcutaneous ATP has not yet been evaluated. Follow-up for subcutaneous ICDs differs from that for transvenous ICDs because patients who have multiple episodes of shocked VT that might benefit from ATP or those who require dual-chamber or biventricular pacing may need revision to a transvenous system.

Pacemaker and ICD pulse generators (Fig. e36-3) have a clear plastic header, to which the leads are attached, and a titanium casing, or “can,” that houses the electronic components. The volume of the can is in the range of 10 to 15 cm³ for pacemakers and 30 to 35 cm³ for ICDs. Components common to both include the battery, voltage supply/control unit, microprocessor, ROM and RAM memory, telemetry control, system controller, rate-adaptive sensors, filters, sensing amplifier, and pacing output circuit/control unit. ICDs have additional high-voltage components, including a transformer, capacitor, and output circuitry.

The High-Voltage Capacitor.

A capacitor consists of two conductors separated by an insulator (dielectric). They store electrical charge on the surface of the conductors, store electrical energy in the field between the two conductors, and determine the duration required to deliver the shock defibrillation waveform. The energy stored in a capacitor is derived by

Estd=12CV2

This equation links stored energy—a key determinant of ICD size—to the voltage stored on the capacitor, which except for minimal voltage loss in the output circuit, is equal to the initial shock waveform voltage (V_i). ICD capacitors have an energy density that ranges from 3 J/cm³ for aluminum electrolytic capacitors to 5 J/cm³ for tantalum powder capacitors, almost 1000 times less than that for batteries.