Ablation energy sources

Radiofrequency energy delivery

Delivery of RF energy depends on the establishment of an electrical circuit involving the human body as one of its in-series elements. The RF current is applied to the tissue via a metal electrode at the tip of the ablation catheter and is generally delivered in a unipolar fashion between the tip electrode and a ground pad applied to the patient’s skin. Bipolar RF systems also exist, in which the current flows between two closely apposed small electrodes, thus limiting the current flow to small tissue volumes interposed between the metal conductors. Bipolar systems, partly because of their relative safety, are now the preferred tools in electrosurgery (oncology, plastic surgery, and ophthalmology) and have been increasingly utilized for catheter ablation of cardiac arrhythmias.

Unipolar radiofrequency systems

Unipolar systems apply RF energy between the ablation catheter tip electrode and a large dispersive electrode (indifferent electrode, or ground pad) applied to the patient’s skin. The polarity of connections from the electrodes to the generator is not important because the RF current is an alternating current. The system impedance comprises the impedance of the generator, transmission lines, catheter, electrode-tissue interface, dispersive electrode-skin interface, and interposed tissues. As electricity flows through a circuit, every point of that circuit represents a drop in voltage, and some energy is dissipated as heat. The point of greatest drop in line voltage represents the area of highest impedance and is where most of the electrical energy becomes dissipated as heat. Therefore, with excessive electrical resistance in the transmission line, the line actually warms up and power is lost. Currently used electrical conductors from the generator all the way through to the patient and from the dispersive electrode back to the generator have low impedance, to minimize power loss.

Flow of the RF current from the ablation electrode through the myocardium to the indifferent electrode results in resistive tissue heating and lesion formation. With normal electrode-tissue contact, only a fraction of all power is effectively applied to the tissue. The rest is dissipated in the blood pool and elsewhere in the patient. With an ablation electrode in contact with the endocardial wall, part of the electrode contacts tissue and the rest contacts blood, and the RF current flows through both the myocardium and the blood pool surrounding the electrode. The distribution between both depends on the impedance of both routes and also on how much electrode surface contacts blood versus cardiac tissue. Whereas tissue heating is the goal of power delivery, the blood pool is the most attractive route for RF current because blood is a better conductor and has significantly lower impedance than tissue and because the contact between electrode and blood is often better than with tissue. Therefore, with normal electrode-tissue contact, much more power is generally delivered to blood than to cardiac tissue.

After leaving the electrode-blood-tissue interface, the current flows through the thorax to the ground pad. Part of the RF power is lost in the patient’s body, including the area near the ground pad. However, because the surface area of the ablation electrode (approximately 28 mm ² ) is much smaller than that of the dispersive electrode (approximately 100–250 cm ² ), the current density (current in amperes divided by surface area) is higher in the immediate vicinity of the ablation electrode, and heating occurs preferentially at that site, with no significant heating occurring at the dispersive electrode. Nonetheless, when RF energy delivery is power-limited, dissipation of energy can occur at the dispersive electrode site (at the contact point between the ground pad and the skin) to a degree that can limit lesion formation at the ablation electrode.

The dispersive electrode may be placed on any convenient skin surface. The geometry of the RF current field is defined by the geometry of the ablation electrode and is relatively uniform in the region of volume heating. Thus, the position of the dispersive electrode (on the patient’s back or thigh) has little effect on impedance, voltage, current delivery, catheter tip temperature, or geometry of the resulting lesion. The size of the dispersive electrode, however, is important. Sometimes it is advantageous to increase the surface area of the dispersive electrode. This increase leads to lower impedance, higher current delivery, increased catheter tip temperatures, and more effective tissue heating. This is especially true in patients with baseline system impedance greater than 100 Ω. Moreover, when the system is power limited, as with a 50-W generator, heat production at the catheter tip varies with the proportion of the local electrode-tissue interface impedance to the overall system impedance. If the impedance at the skin-dispersive electrode interface is high, then a smaller amount of energy is available for tissue heating at the electrode tip. Therefore, when ablating certain sites, adding a second dispersive electrode or optimizing the contact between the dispersive electrode and skin should result in relatively more power delivery to the target tissue. Additionally, meticulous skin preparation is imperative to optimize skin-patch contact and minimize impedance at the skin interface with the dispersive electrode.

A large surface area, and good skin contact, is required at the indifferent electrode patch not only to effectively dissipate heat but also to prevent skin burns. The skin temperature beneath the patch is inversely proportional to patch surface area in contact with the skin and the distance between the patch and the active electrode. If ablation is performed with a high-amplitude current (more than 50 W) and skin contact by the dispersive electrode is poor, skin temperatures can potentially rise above 45°C which, even when short lasting, can result in significant skin burns, especially the “leading edge” of the patch (the side closest to the RF catheter electrode) ( eFig. 8.1 ).

Bipolar radiofrequency systems

Bipolar RF ablation uses two adjacent ablation electrodes between which the RF flows, leading to lesion formation between electrodes. Although used in cardiac surgery, bipolar RF ablation has not been widely adopted as a catheter-based treatment for clinical arrhythmias.

A number of in vitro, in vivo, and small-cohort clinical studies have demonstrated that bipolar RF ablation using two ablation catheter tips (at opposite sides of the myocardial wall) as the active and ground poles, can potentially create deeper, transmural lesions as compared to sequential or simultaneous unipolar ablation with the same settings ( Fig. 8.1 ). This can be accomplished by configuring an existing unipolar RF circuit to designate a second catheter as the grounding connection in lieu of an impedance patch (using custom-designed cables).

In bipolar ablation configurations, a higher impedance circuit is created (given the smaller surface area of the ground catheter bipolar ablation in comparison to the ground pad in unipolar ablation), leading to higher current density in synergy with increased resistive heating. Additionally, simultaneous tissue heating results in the concentration of thermal injury within the zone between the two closely spaced ablation electrodes, increased conductive heating, and more efficient lesion formation. The ability to achieve transmurality with bipolar RF ablation has been reported with tissue thickness up to 25 mm, in contrast to sequential unipolar RF, which was unable to achieve transmurality when thicknesses are greater than 15 mm. This can offer a significant advantage for ablation of septal atrial tachycardia or ventricular tachycardia (VT) circuits (with one ablation catheter positioned at one side of the septum and the second catheter positioned at the opposite side) or intramural free wall VT circuits (with one ablation catheter positioned endocardially and the second catheter positioned epicardially).

In another design, bipolar RF ablation is used to create contiguous lesions between adjacent electrodes mounted on a multielectrode ablation catheter positioned at one side of the myocardial wall. A multichannel duty-cycled RF ablation generator (GENius™ MultiChannel RF Generator, Medtronic, Minneapolis, MN, USA) contains 12 independently controlled RF generators and is capable of delivering RF energy to each electrode independently. RF energy can be delivered in either bipolar (between the electrodes) and unipolar (from the electrode to the ground pad) current by a phase difference between the channels (thus the term “phased RF”). During unipolar ablation, current flow is provided between the electrodes and an indifferent electrode attached to the patient’s skin, while in the bipolar energy mode, current flows between adjacent electrodes. In this design, bipolar RF ablation lesions are longer but shallower than unipolar RF lesions. To improve lesion depth, RF energy can also be delivered in a unipolar mode from each electrode to a dispersive electrode.

A second atrial fibrillation (AF) ablation system (nMARQ, Biosense Webster, Diamond Bar, CA, USA) employs a multichannel RF generator capable of independently delivering unipolar or bipolar RF energy to a maximum of 10 electrodes simultaneously. The nMARQ catheter is a decapolar mapping and ablation catheter with individually irrigated platinum electrodes. Each electrode possesses a thermocouple and holes for irrigation. RF energy is applied in a temperature-controlled mode and energy delivery can be individually arranged over each combination of the 10 electrodes in unipolar mode or bipolar mode between two adjacent electrodes.

Tissue heating

Tissue heating resulting from RF ablation depends on two aspects: (1) resistive heating directly related to electrical current flow at the catheter-tissue interface; and (2) conductive heating extending passively from resistive heating to adjacent tissue.

Whereas free electrons serve as charge carriers inside the RF generator, cables, and RF electrode, the electric current inside tissue is carried by charged ions (such as Na ⁺ , K ⁺ , and Cl ^– ). During alternating current flow, the ions attempt to follow the changes in the direction of the alternating current. For a typical frequency of 500 kHz, the direction of the current (and ion movement) changes a million times per second. Ion oscillations generate heat due to friction, thus converting electromagnetic (current) energy into molecular mechanical energy or heat. This type of electric current-mediated heating is known as ohmic (resistive) heating.

Using Ohm’s law, with resistive heating, the amount of power (= heat) per unit volume equals the square of current density times the specific impedance of the tissue. Hence, local resistive heating is proportional to the square of the current density. The current density for an electrode can be approximated by dividing the current delivered to the ablation electrode by the effective surface area over which that current is delivered. With a spherical ablation electrode, the current distributes radially from the source, and current density therefore decreases with the square of distance from the center of the electrode. As a consequence, power dissipation per unit volume decreases with the fourth power of distance. The thickness of the electrode eliminates the first steepest part of this curve, however, and the decrease in dissipated power with distance is therefore somewhat less dramatic. Approximately 90% of all RF power that is delivered to the tissue is absorbed within the first 1 to 1.5 mm from the ablation electrode surface. Therefore, only a thin rim of tissue in immediate contact with the RF electrode is directly heated (within the first 2 mm of depth from the electrode). The remainder of tissue heating occurs passively as a result of heat conduction from this rim to the surrounding tissues.

Resistive heating is rapid, beginning and ending immediately with the initiation and cessation of RF energy application. On initiation of fixed-level energy application, the temperature at the electrode-tissue interface rises monoexponentially to reach steady state within 7 to 10 seconds, and the steady state is usually maintained between 80°C and 90°C. On the other hand, conduction of heat to deeper tissue is a passive, relatively slow process and requires 30 seconds to 2 minutes to equilibrate (thermal equilibrium). Therefore, the rate of tissue temperature rise beyond the immediate vicinity of the RF electrode is much slower, resulting in a steep radial temperature gradient as tissue temperature decreases radially in proportion to the distance from the ablation electrode, and deep tissue temperatures continue to rise for several seconds after interruption of RF delivery (the so-called thermal latency phenomenon). Therefore, RF ablation requires at least 30 to 60 seconds to create full-grown lesions. In addition, when temperature differences between adjacent areas develop because of differences in local current density or local heat capacity, heat conducts from hotter to colder areas, thus causing the temperature of the former to decrease and that of the latter to increase. Furthermore, heat loss to the blood pool at the surface and to intramyocardial blood vessels determines the temperature profile within the tissue.

At steady state, the lesion size is proportional to the temperature measured at the interface between the tissue and the electrode, as well as to the RF power amplitude. By using higher powers and achieving higher tissue temperatures, the lesion size can be increased. Tissue temperatures of up to 110°C can be achieved. Above this temperature, tissue vaporization ensues and prevents any further heating by RF current due to the electrically insulating properties of vapor. Also, boiling of the plasma at the electrode-tissue interface can cause tissue carbonization (“charring”) at the site of very high RF current densities (typically very close to the RF electrode), producing an electrically insulating coagulum, which is accompanied by a sudden increase in electrical impedance that prevents further current flow into the tissue and further heating.

The range of desired tissue temperatures during RF ablation is 50°C to 90°C. Within this range, relatively uniform desiccation of tissue can be expected. If the temperature is lower than 50°C, no or only minimal tissue necrosis results. Temperatures above 100°C are associated with tissue vaporization, including steam pops, and charring. Because the rate of temperature rise at deeper sites within the myocardium is slow, a continuous energy delivery of at least 60 seconds is often warranted to maximize depth of lesion formation.

Ablation using high-power (50–90 W) short-duration (for less than 15 seconds) RF applications seeks to change the balance between resistive and conductive tissue heating, whereby the size of the thermal lesion depends almost exclusively on the electrical energy deposited, instead of on the thermal conduction phenomenon. Application of high RF current allows for increased local resistive heating; however, because the ablation duration is very brief, temperature decay is significantly shorter, resulting in restriction of heat conduction to deeper tissues. Thus, the thermal lesion size would depend on the electrical conductivity of the tissue (which is responsible for resistive heating) rather than its thermal conductivity; and the lesion would be created instantly, for as long as RF application continues. This is in contrast to conventional RF power settings, which rely on passive heat conduction into deeper tissue for transmural lesion formation (thus requiring longer RF application at lower power).

High-power, short-duration RF ablation produces wider, shallower lesions and much less reversible injury ( Figs. 8.2 and 8.3 ). This strategy has been proposed in cardiac structures with limited wall thickness (e.g., during pulmonary vein [PV] electrical isolation for AF), whereby lesion transmurality is not dependent on substantial conductive heating. This has the potential of avoiding passive heating of deeper tissue beyond the thin left atrial (LA) wall, with the possibility of reducing damage to extracardiac structures, resulting in an improved safety profile. Conversely, increased local resistive heating could increase the risk of tissue vaporization (“steam pops”) or coagulum formation.

Convective cooling

At the site of endocardial ablation, tissues heated directly by RF energy start losing heat into deeper layers of myocardium as well as into the circulating blood pool and the metal ablation electrode ( Fig. 8.4 ). Heat conduction into deeper myocardial layers helps increase the size of the ablation lesion. Heat conduction to the ablation electrode helps indirect monitoring of tissue temperature via the embedded temperature sensors. On the other hand, convective heat loss from the endocardium at the site of ablation into the circulating blood pool results in cooling of the endocardial surface and constitutes the dominant factor opposing effective myocardial heating and lesion formation. Because the tissue surface is cooled by the blood flow, the highest temperature during RF delivery occurs slightly below the endocardial surface. Consequently, the width of the endocardial lesion matures earlier than the intramural lesion width (20 seconds versus 90–120 seconds). Hence, the maximum lesion width is usually located intramurally, and the resultant lesion is usually teardrop shaped, with less necrosis of the superficial tissue.

As the magnitude of convective cooling increases (e.g., unstable catheter position, poor catheter-tissue contact, or high blood flow in the region of catheter position), there is reduced efficiency of heating as more energy is carried away in the blood and less energy is delivered to the tissue. When RF power is limited and insufficient to overcome the heat lost by convection, lesion size is reduced. On the other hand, when RF power delivery is not limited, convective cooling allows for more power to be delivered into the tissue (as long as adequate electrode-tissue contact is maintained), increasing the depth of direct resistive heating and intramural tissue temperatures (and hence larger ablation lesion size), while avoiding very high endocardial surface temperatures (and temperatures measured by the catheter tip sensors) that can otherwise result in blood boiling and coagulum formation at the electrode tip, with consequent sudden rise in electrical impedance and reduction of further power delivery.

The concept of convective cooling can explain why there are few coronary arterial complications with conventional RF ablation. Coronary arteries act as a heat sink; substantive heating of vascular endothelium is generally prevented by heat dissipation in the high-velocity coronary blood flow, even when the catheter is positioned close to the vessel. Although this is advantageous, because coronary arteries are being protected, it can limit success of the ablation lesion if a large perforating artery is close to the ablation target and can protect the target tissue from ablation by transferring heat away from it.

The effects of convective cooling have been exploited to increase the size of catheter ablative lesions. Active electrode cooling by ablation electrode irrigation is currently used to largely eliminate the risk of overheating at the electrode-tissue contact point and increase the magnitude of power delivery and the depth of volume heating.

Of note, standard RF ablation is not as effective when delivered to the epicardium compared with lesions delivered to the endocardium for a number of reasons, one of which is the lack of convective cooling (by the circulating blood) of the ablation electrode in the epicardial space. This results in high electrode temperatures at low-power settings (≤10 W), limiting power delivery in the pericardial space and hindering lesion formation. Therefore, active electrode cooling is required in this setting.

Catheter tip temperature

Ablation catheter tip temperature depends on tissue temperature, tissue contact of the ablation electrode, convective cooling by the surrounding blood, heat capacity of the electrode material, and type and location of the temperature sensor.

Catheter tip temperature is measured by a sensor located in the ablation electrode. There are two different types of temperature sensors: thermistors and thermocouples. Thermistors require a driving current, and the electrical resistance changes as the temperature of the electrical conductor changes. More frequently used are thermocouples, which consist of copper and constantan wires and are incorporated in the center of the ablation electrode. Thermocouples are based on the so-called “Seebeck effect”; when two different metals are connected (sensing junction), a voltage can be measured at the reference junction that is proportional to the temperature difference between the two metals.

The electrode temperature rise is an indirect process—the ablation electrode is not heated by RF energy, but it heats up because it happens to touch heated tissue. Consequently, the catheter tip temperature is always lower than, or ideally equal to, the superficial tissue temperature. Conventional electrode catheters with temperature monitoring report the temperature only from the center of the electrode mass with one design, or from the apex of the tip of the catheter with another design. It is likely that the measured temperature underestimates the peak tissue temperature (which occurs slightly below the endocardial surface).

Several other factors can increase the disparity between catheter tip temperature and tissue temperature, including catheter tip irrigation, large ablation electrode size, and poor electrode-tissue contact. With a large electrode tip, a larger area of the electrode tip is exposed to the cooling effects of the blood flow than with standard tip lengths, thus resulting in lower electrode temperatures. Similarly, with poor electrode-tissue contact, less electrode material is in contact with the tissue, and heating of the tip by the tissue occurs at a lower rate, resulting in relatively low tip temperatures. Catheter tip irrigation increases the disparity between tissue temperature and electrode temperature because it results in cooling of the ablation electrode, but not the tissue. Of note, a novel irrigated-tip catheter design (the DiamondTemp ablation catheter [Medtronic, Minneapolis, MN, USA]) allows for improved tissue surface temperature monitoring (see below).

Cellular effects of radiofrequency ablation

The primary mechanism of tissue injury by RF ablation is likely to be thermally mediated. Hyperthermic injury to the myocyte is both time and temperature dependent, and it is likely the result of changes in the cell membrane, protein inactivation, cytoskeletal disruption, and nuclear degeneration, as well as other potential mechanisms. The cell membrane, in particular, is very sensitive to thermal injury. Hyperthermia results in potential phase change in membrane fluidity, kinetic and structural changes in membrane ion channels and ion pumps, inhibition of transport proteins, and formation of nonspecific ionic membrane pores.

Experimentally, the resting membrane depolarization is related to temperature. In the low hyperthermic range (37–45°C), little tissue injury occurs, and a minor change can be observed in the resting membrane potential and action potential amplitude. However, action potential duration shortens significantly, and conduction velocity becomes greater than at baseline. In the intermediate hyperthermic range (45–50°C), progressive depolarization of the resting membrane potential occurs, and action potential amplitude and rate of phase I depolarization (dV/dt) decrease. Additionally, abnormal automaticity is observed, reversible loss of excitability develops, and conduction velocity progressively decreases. In the high-temperature ranges (higher than 50°C), marked depolarization of the resting membrane potential occurs, and permanent loss of excitability is observed. Temporary (at temperatures of 49.5–51.5°C) and then permanent (at 51.7–54.4°C) conduction block develops, and fairly reliable irreversible myocardial injury occurs with a short hyperthermic exposure.

RF ablation typically results in high temperatures (70–90°C) for a short time (up to 60 seconds) at the electrode-tissue interface but significantly lower temperatures at deeper tissue sites. This results in rapid tissue injury within the immediate vicinity of the RF electrode but relatively delayed myocardial injury with increasing distance from the RF electrode. Therefore, although irreversible loss of electrophysiological (EP) function can usually be demonstrated immediately after successful RF ablation, this finding can be delayed because tissue temperatures continue to rise somewhat after cessation of RF energy delivery (thermal latency phenomenon). This effect can account for the observation that patients undergoing atrioventricular nodal (AVN) modification procedures who demonstrate transient heart block during RF energy delivery can progress to persisting complete heart block, even if RF energy delivery is terminated immediately. Reversible loss of conduction can be demonstrated within seconds of initiating the RF application, which can be caused by an acute electrotonic effect. On the other hand, there can be late recovery of EP function after an initial successful ablation, likely related to reversible injury to the myocardium by inadequate RF energy delivery; initial loss of EP function due to tissue heating resolves after healing of the damaged, but surviving, myocardium.

In addition to the dominant thermal effects of RF ablation, some of the cellular injury has been hypothesized to be caused by a direct electrical effect, which can result in dielectric breakdown of the sarcolemmal membrane with creation of transmembrane pores (electroporation), resulting in nonspecific ion transit, cellular depolarization, calcium overload, and cell death. Such an effect has been demonstrated with the use of high-voltage electrical current. However, it is difficult to examine the purely electrical effects in isolation of the dominant thermal injury.

Tissue effects of radiofrequency ablation

Changes in myocardial tissue are apparent immediately upon completion of the RF lesion. Pallor of the central zone of the lesion is attributable to denaturation of myocyte proteins (principally myoglobin) and subsequent loss of the red pigmentation. Slight deformation, indicating volume loss, occurs at the point of catheter contact in the central region of lesion formation. The endocardial surface is usually covered with a thin fibrin layer and, occasionally, if a temperature of 100°C has been exceeded, with char and thrombus ( Fig. 8.7 ). In addition, a coagulum (an accumulation of fibrin, platelets, and other blood and tissue components) can form at the ablation electrode because of the boiling of blood and tissue serum.

On sectioning, the central portion of the RF ablation lesion shows desiccation, with a surrounding region of hemorrhagic tissue and then normal-appearing tissue. Histologic examination of an acute lesion shows typical coagulation necrosis with basophilic stippling consistent with intracellular calcium overload. Immediately surrounding the central lesion is a region of hemorrhage and acute monocellular and neutrophilic inflammation. The effect on extracellular connective tissue is maximal when closer to the heat source, and the damage extended less deeply in comparison to the cardiomyocytes, which is likely related to the higher denaturation threshold of collagen (∼70°C) in comparison to cardiomyocytes (∼55°C).

The progressive changes seen in the evolution of an RF lesion are typical of healing after any acute injury. Within 2 months of the ablation, the lesion shows fibrosis, granulation tissue, chronic inflammatory infiltrates, and significant volume contraction. The lesion border is well demarcated from the surrounding viable myocardium without evidence of a transitional zone. This likely accounts for the absence of proarrhythmic side effects of RF catheter ablation. As noted, because of the high-velocity blood flow within the epicardial coronary arteries, these vessels are continuously cooled and are typically spared from injury, despite nearby delivery of RF energy. Nonetheless, high RF power delivery in small hearts, such as in pediatric patients, or in direct contact with the vessel can potentially cause coronary arterial injury.

The border zone around the acute pathological RF lesion accounts for several phenomena observed clinically. The border zone is characterized by marked ultrastructural abnormalities of the microvasculature and myocytes acutely, as well as a typical inflammatory response later. The most thermally sensitive structures appear to be the plasma membrane and gap junctions, which show morphological changes as far as 6 mm from the edge of the pathological lesion. The border zone accounts for documented effects of RF lesion formation well beyond the acute pathological lesion. The progression of the EP effects after completion of the ablation procedure can be caused by further inflammatory injury and necrosis in the border zone region that result in late progression of physiological block and a delayed cure in some cases. On the other hand, initial stunning and then early or late recovery of function can be demonstrated in the border zone, thus accounting for the recovery of EP function after successful catheter ablation in the clinical setting, which can be caused by healing of the damaged, but surviving, myocardium.

The effects of RF ablation in abnormal myocardium are unpredictable and poorly understood. Experimental studies found significant differences between the effect of RF ablation in normal myocardium and those in scarred myocardium. In contrast to homogeneous thermal injury with uniform lesion shape resulting from RF ablation in healthy myocardium, ablation lesions in heterogenous scar demonstrate variability in shape with indistinct borders and irregular architecture, likely related to the heterogeneous thermodynamic properties of scar tissue components. Microscopically, thermal injury causes an erratic pattern, resulting in cardiomyocytes exhibiting features of coagulation or contraction band necrosis interspersed between islands of surviving myocardial bundles surrounded by collagen. Acute RF ablation lesion volumes in scar are significantly smaller than those in normal myocardium. Scarred myocardium (such as postinfarction ventricular myocardium) can exhibit complex nonhomogeneous strands and islands of cardiomyocytes partially encapsulated by collagen and interspersed by adipose tissue. Electrical resistance of scar tissue is approximately 20% to 40% lower than that in healthy myocardium; hence, RF energy passing through collagen produces a smaller degree of heating as compared to normal myocardium. Additionally, the higher denaturation temperature threshold of collagen can potentially protect cardiomyocyte bundles from RF-induced thermal injury. Furthermore, adipose tissue, often present in remodeled scar, has higher electrical resistivity (25-fold increased resistivity compared to that of cardiomyocytes); hence, it serves as a heat capacitor, diverting current from cardiomyocytes to fat, which can potentially protect the surviving cardiomyocytes from thermal injury.

Coagulum formation

Excessive tissue heating can lead to high temperatures at the electrode-tissue interface. Once the peak tissue temperature exceeds the threshold of 100°C, boiling of blood and tissue serum at the electrode-tissue interface can ensue. When boiling occurs, denatured serum proteins and charred tissue form a thin film that adheres to the electrode, thus producing an electrically insulating coagulum (an accumulation of fibrin, platelets, and other blood and tissue components). The coagulum can detach and embolize. Of note, unlike typical thrombi, a coagulum is not formed by activation of clotting factors, and its formation is not prevented by anticoagulation.

In temperature-controlled RF ablation, electrode temperature does not reach the boiling point. Nonetheless, the true tissue temperature can be significantly higher than the measured electrode temperature. Additionally, serum proteins denature at temperatures well below the interfacial boiling temperature. Therefore, coagulum can still form even when the electrode temperature is limited to 65°C with a 4-mm ablation electrode, and 55°C with an 8-mm electrode. Open irrigation cools the electrode and its direct environment (blood and catheter-tissue interface), reducing (but not eliminating) the risk of interfacial boiling and coagulum formation. Also, the irrigant helps wash away denatured proteins as it forms.

Usually, coagulum adheres to the electrode tip, which is accompanied by a sudden increase in electrical impedance that prevents further current flow into the tissue and further heating, at which point the ablation catheter needs to be withdrawn from the body and cleaned to remove the adhering coagulum. However, when the coagulum attaches to the tissue surface rather than to the electrode tip, electrode temperature or impedance may not be affected, and coagulum formation can continue unnoticed. Therefore, the absence of impedance rise during ablation does not guarantee the absence of coagulum formation on the tissue contact site.

Steam pop

High RF power application (especially in the setting of cooled RF ablation) can cause superheating within the tissue (with subendocardial tissue temperatures exceeding 100°C), which can result in boiling of water within the tissue under the electrode. Consequently, evaporation and rapid steam expansion can occur intramurally, and a gas bubble can develop in the tissue under the ablation electrode. Continued application of RF energy causes the bubble to expand and its pressure to increase, which can lead to eruption of the gas bubble (causing a popping sound) through the path with the least mechanical resistance that can leave behind a gaping hole (the so-called “steam pop”). This often occurs toward the heat-damaged endocardial surface (leading to crater formation) or, infrequently, across the myocardial wall (resulting in myocardial perforation).

The consequence of a steam pop depends on the area of the heart being ablated. Rupture of the steam bubble toward the endocardium can produce a crater composed of necrotic debris and exposed muscle fibers, which can lead to cardioembolic sequela from cellular debris and release of tissue factors that can be prothrombogenic. The risk of cardiac perforation is low in areas of dense ventricular scar. The risk of perforation and cardiac tamponade is likely to be higher for ablation in the thin-walled right ventricular (RV) outflow tract and in the atria. Therefore, it is prudent to take a more conservative approach to power application in these areas.

Steam pops are often associated with a sudden, although small (<10 Ω), impedance rise and a sudden drop in electrode temperature. When steam pops burst toward the endocardial surface, a shower of microbubbles is often detected on intracardiac echocardiography (ICE).

The risk of steam pops is relatively low with temperature-controlled standard 4-mm-tip RF catheter ablation, since the electrode temperature, which approximates tissue temperature, is limited to a safe level. However, when there is significant discrepancy between tissue and electrode temperature, due to passive or active electrode cooling, tissue temperature can reach the boiling point without being detected by the monitoring electrode temperature.

Electrode orientation also seems to affect the significance of steam pops; pops that occur when the electrode tip is perpendicular to the tissue are more likely to cause cardiac perforation than those that occur when the electrode is lying horizontally on the tissue. Therefore, one should try to avoid high-pressure perpendicular tissue contact, especially at higher RF power outputs. RF applications preceding steam pops have a greater and more rapid decrease in impedance and occur at a higher maximum RF power than applications without pops. Because of the considerable overlap of impedance changes in lesions with and without steam pops, it is not possible to advocate a general limit for impedance decrease for all RF lesions. However, when ablation is performed in areas at risk for perforation (i.e., especially in thin-walled structures), reducing power to achieve an impedance decrease of less than 18 Ω is a reasonable strategy to reduce the chance of a pop.

Ablation electrode temperature

The relationship between lesion size and the recorded catheter tip temperature is complex and subject to several confounding factors. When electrode tip temperature closely resembles tissue temperature (the main determinant of lesion formation), lesion size appears to increase directly with electrode temperature up until the point of coagulum formation at the electrode-tissue interface, when further power delivery becomes limited. However, with good contact between catheter tip and tissue and low cooling of the catheter tip, the target temperature can be reached with little power, thus resulting in fairly small lesions, although a high catheter tip temperature is being measured.

Importantly, catheter tip temperature is not a reliable measure of tissue temperature, as it is also influenced by convective cooling, electrode-tissue contact, and type and location of the temperature sensor. A low catheter tip temperature can be caused by a high level of convective cooling, which allows a higher amount of RF power to be delivered to the tissue (because it is no longer limited by temperature rise of the ablation electrode) and yields relatively large lesions. This is best illustrated with active cooling of the ablation electrode using irrigation during RF energy delivery; the tip temperature is usually less than 40°C, which allows the application of high-power output for longer durations.

Radiofrequency power amplitude

Lesion size is proportional to the amount of RF power delivered effectively into the tissue. Higher RF power delivery increases the amount of directly heated tissue as well as tissue temperatures, resulting in greater depth of thermal injury and larger lesion size. However, the mere amplitude of RF power application does not necessarily translate to larger amount of power delivered into the tissue, since RF power can be wasted into the surrounding blood pool due to poor electrode-tissue contact.

Duration of radiofrequency application

The RF lesion is predominantly generated within the first 10 seconds of target energy delivery and tissue temperatures, and it reaches a maximum after 30 seconds. Extension of RF application beyond 45 to 60 seconds during power-controlled RF delivery does not generally seem to increase lesion size further. However, lesions created using large or irrigated RF electrodes have a longer half-time of lesion growth, and extended ablation (60–120 seconds) can potentially help create larger lesions.

Of note, a recent report found that, compared to lesions resulting from continuous ablation, lesions resulting from interrupted ablation of the same total duration (whether the interruption is intentional or caused by poor catheter stability and intermittent tissue contact due to respiratory or cardiac motion) have a smaller overall lesion volume, predominantly due to reduced lesion width. The smaller lesion size with RF interruption is likely related to the amount of time necessary for energy to penetrate deeply into the myocardium. The heating of deeper tissues occurs gradually so that temperatures in these deeper regions continue to increase throughout a 60-second ablation. Additionally, changes in tissue caused by prior ablation (such as edema) can potentially limit subsequent ablation lesions.

These findings have important clinical implications. Interruption of RF energy delivery is likely to produce wide-based, shallow lesions, which may limit effectiveness in ablating deep intramyocardial targets. Strategies aimed at reducing the variability of catheter-tissue contact and minimizing disruption of the RF energy delivery can improve efficacy of ablation. Also, if energy delivery is disrupted, applications may need closer spacing to avoid gaps.

Electrode-tissue contact

Catheter electrode-tissue contact pressure is a major determinant of lesion size. Data indicate that increasing the firmness of contact (contact force, CF) is comparable to increasing RF power; that is, the same tissue temperatures and lesion size can be reached at a much lower power level when tissue contact is optimized.

The efficiency of energy transfer to the myocardium (i.e., temperature rise per watt of applied power) largely depends on the electrode-tissue contact. Increasing CF can improve the efficiency of RF lesion formation by several mechanisms: (1) expanding the area of the contact footprint of the electrode on myocardial tissue by embedding it deeper within the soft myocardium, which leads to a higher amount of RF power that can be effectively delivered to the tissue; (2) reducing the electrode surface exposed to surrounding blood pool and, hence, reducing RF current shunting into the low-impedance blood pool; (3) improving stability of the electrode-tissue contact and reducing catheter sliding with cardiac motion, which improves the efficiency of lesion formation; (4) stretching of the myocardium, bringing the epicardium closer to the ablation electrode and, hence, improving transmurality of the RF lesion.

However, at a certain moderate CF, further increase in contact firmness results in progressively smaller lesions because a lesser amount of RF power is required to reach target temperature. The temperature rise of the ablation electrode signals excessive tissue heating and limits power delivery in a temperature-controlled system. Reduced convective cooling of the ablation electrode (because of lesser interface with the blood pool) results in higher electrode temperatures and can further reduce RF current delivery.

Electrode length

Ablation catheters have tip electrodes that are conventionally 4 mm long and are available in sizes up to 10 mm long. With small ablation electrodes, the current density at the electrode-tissue interface will be high and tissue heating is more efficient. An increased electrode size reduces the efficiency of power transfer to the tissue (due to reduced delivery of the same power over a larger electrode surface area). Therefore, with the same total power, lesions created with a larger electrode are always smaller than lesions created with a smaller electrode ( Fig. 8.8 ). A larger electrode size also creates a greater variability in power transfer to the tissue because of greater variability of tissue contact, and tissue contact becomes much more dependent on catheter orientation with longer electrodes. Consequently, an 8-mm electrode may require a 1.5 to 4 times higher power level than a 4-mm electrode to create the same lesion size. On the other hand, in the absence of irrigation, the more efficient tissue heating with smaller electrodes results in rapid increase in the electrode-tissue interface, causing char formation and steam pops. Also, in temperature-controlled RF ablation, the quick rise in electrode temperature limits further RF power delivery into the tissues.

On the other hand, when the power is not limited, catheters with large distal electrodes allow for greater power delivery to create deeper lesions both by increasing the ablation electrode surface area in contact with the bloodstream (resulting in an augmented convective cooling effect) and by increasing the volume of tissue directly heated because of an increased surface area at the electrode-tissue interface ( Fig. 8.8 ). However, this assumes that the electrode-tissue contact, tissue heat dissipation, and blood flow are uniform throughout the electrode-tissue interface. As the electrode size increases, the likelihood that these assumptions are true diminishes because of variability in cardiac chamber trabeculations and curvature, tissue perfusion, and intracardiac blood flow, which affect the heat dissipation and tissue contact. These factors result in unpredictable lesion size and uniformity for electrodes more than 8 mm long.

There is a potential safety concern with the use of long ablation electrodes because of nonuniform heating, with maximal heating occurring at the electrode edges. Thus, large electrode-tipped catheters with only a single thermistor can underestimate maximal temperature and allow char formation and potential thromboembolic complications. Catheter tips with multiple temperature sensors at the electrode edges are preferable for temperature feedback. In addition, the greater variation in power delivered to the tissue and the greater discrepancy between electrode and tissue temperature make it difficult to avoid steam pops and char formation. Another point of concern is that the formation of char may only minimally affect electrode impedance by covering a much smaller part of the electrode surface. Therefore, the lower electrode temperature and the absence of any impedance rise may erroneously suggest a safer ablation process.

Other limitations of a large ablation electrode (8–10 mm in length) include the reduction in mobility and the flexibility of the catheter (which can impair positioning of the ablation electrode) and a reduction in the resolution of recordings from the ablation electrode. A larger electrode dampens the local electrogram, especially that of the distal electrode. In contrast, a smaller electrode improves mapping accuracy and feedback of tissue heating; its only drawback is the limited power level that can be applied to the tissue.

Electrode orientation

The effect of catheter orientation on lesion size depends on the presence of passive or active cooling, whether or not power delivery is restricted, the length of the ablation electrode, and the placement of the temperature sensors relative to the portion of the electrode in contact with tissue.

Lesion size is only slightly affected by catheter tip orientation using 4- or 5-mm-long tip catheters; the effects of catheter orientation become more pronounced for larger ablation electrodes. When the catheter tip is perpendicular to the tissue surface, a much smaller surface area is in contact with the tissue (resulting in increased current density at the electrode-tissue interface) and larger area exposed to the circulating blood pool (resulting in augmented convective cooling) than when the catheter electrode tip is lying on its side. On the other hand, parallel tip orientation provides larger electrode-tissue contact area (resulting in less power waste into the blood stream and less current density at the electrode-tissue interface). Therefore, with the same total power, perpendicular electrode orientation yields larger lesion volumes than parallel electrode orientation. However, if the RF power level is unrestricted and is increased to maintain a constant current density, lesion size (particularly superficial lesion length) will increase proportionally to the electrode-tissue contact area, which is larger with parallel tip orientation.

Furthermore, the character of the lesion created with temperature control depends on the placement of the temperature sensors relative to the portion of the electrode in contact with tissue. Thus, the orientation of the electrode and its temperature sensors determines the appropriate target temperature required to create maximal lesions while avoiding coagulum formation caused by overheating at any location within the electrode-tissue interface.

Electrode material

Although platinum-iridium electrodes have been the standard for most RF ablation catheters, gold exhibits excellent electrical conductive properties as well as more than four times greater thermal conductivity than platinum (300 versus 70 W/m°K), although both materials have similar heat capacities (130 and 135 J/kg °K). The higher thermal conductivity of gold can potentially lead to greater power delivery because of better heat conduction at the tissue-electrode interface and to enhanced cooling as a result of heat loss to the surrounding blood with this electrode material. However, the higher thermal conductivity of gold electrodes is of no advantage in areas of low blood flow (e.g., between myocardial trabeculae), where convective cooling at the electrode tip is minimal. Under these circumstances, electrode materials with a low thermal conductivity can produce larger lesions.

In-vitro experiments have demonstrated that gold ablation electrodes create wider and deeper lesions than platinum-iridium, potentially improving ablation efficacy. However, superiority of gold-tip catheters in clinical practice remains unproven. Conflicting results were observed in clinical studies comparing 8-mm gold-tip with platinum-iridium–tip catheters for ablation of the CTI. Additionally, during catheter ablation of the slow AVN pathway in patients with AVN reentrant tachycardia (AVNRT), no significant differences were observed between 4-mm gold-tip and platinum-iridium–tip catheters in the primary endpoint or in the increases of power or temperature at any of the measured time points. Interestingly, a significant reduction of charring on gold tips was observed, compared with platinum-iridium material, a finding suggesting a possible advantage of this material beyond its better conduction properties. An irrigated gold-tip ablation catheter (AlCath Flux eXtra Gold, Biotronik, Berlin, Germany) was recently tested and found to allow for improved energy delivery at lower catheter-tip temperatures and at lower irrigation flow rates compared to irrigated platinum–iridium-tip catheters.

Convective cooling

The ablation electrode temperature is dependent on the opposing effects of heating from the tissue and cooling by the blood flowing around the electrode. Because ablation lesion size is primarily dependent on the RF power delivered to the tissue, lesion size varies with the magnitude of convective cooling of the ablation electrode. Electrode cooling can be achieved passively (by local blood flow) or actively (by electrode irrigation).

At any given electrode temperature, the RF power delivered to the tissue is significantly reduced in areas of low local blood flow (e.g., deep pouch in the CTI, dilated and poorly contracting atria, between myocardial trabeculae). The reduced cooling associated with low blood flow causes the electrode to reach the target temperature at lower power levels, and if the ablation lesion is temperature controlled, power delivery will be limited. In these locations, increasing electrode temperature to 65°C or 70°C only minimally increases RF power but it does increase the risk of thrombus formation and impedance rise. Conversely, increasing local blood flow is associated with increased convective cooling of the ablation electrode. Passive electrode cooling can be promoted by using large electrode length or electrode material with high thermal conductivity. Convective cooling can also be achieved by internal or external saline irrigation of the ablation electrode. In the latter setting, the flow rate of the irrigant determines the degree of cooling; faster flow rates likely allow greater power application without impedance rises.

With enhanced convective cooling, more power is delivered to the tissue to reach and maintain target temperature, thus resulting in larger lesion volumes. However, when RF power output is limited, increased convective cooling curtails tissue heating due to increased heat loss into the blood pool. Therefore, at any given power output (with no temperature restriction), RF lesion size decreases with increasing convective cooling. Conversely, at any given electrode temperature (with no power restriction), RF lesion size increases with increasing convective cooling.

Irrigant fluid type

With normal electrode-tissue contact, RF current flows through both the myocardium and the blood pool surrounding the electrode. Since blood is a better conductor and has significantly lower impedance than tissue, most of the RF energy is diverted away from the myocardium. RF current flow into myocardial tissue can be increased by modulating the ratios of electric impedances between RF electrode, blood pool, and cardiac tissue. With open-irrigated RF ablation catheters, the ionic concentration of the irrigation fluid used can influence lesion formation. The standard method for cooling the RF electrode is by circulating normal saline (sodium chloride, NaCl 9 g/L) through the catheter tip during RF delivery. Normal saline has an electrical impedance that is lower than that of blood or tissue ionic and, therefore, RF can preferentially disperse to the saline cloud in contact with the ablation electrode. When using an irrigant with a lower ionic concentration and charge density (i.e., relatively higher electrical impedance) such as half-normal saline (NaCl 4.5 g/L), RF energy may preferentially flow to the targeted cardiac tissue that has a relatively lower electrical impedance.

Experimental studies found that open-irrigation ablation using half-normal saline can produce larger and deeper RF ablation lesions with the same power settings as compared with normal saline irrigation, albeit with an increased risk of steam pops. Irrigation with an isoosmolar nonionic solution (dextrose 5% in water, D5W) resulted in even larger lesions and more frequent steam pops.

Several clinical studies have reported the efficacy of open-irrigation ablation using half-normal saline in cases requiring deeper or more durable lesions, such as ventricular arrhythmias with deep intramural circuits refractory to standard ablation techniques.

System impedance

RF lesion creation is determined by RF current, which is determined by both RF power and impedance. The total impedance of an electrical circuit is determined by the impedances of its individual components within it, which are incorporated in series or in parallel. During RF ablation, circuit impedance (as measured by the RF generator) is influenced by several factors, including intrinsic tissue properties, electrode-tissue contact pressure, and dispersive electrode size and position, as well as patient body habitus. The magnitude of the current delivered by the RF generator is largely determined by the impedance between the ablation catheter and the dispersive electrode. Using Ohm’s law, the RF current (I) delivered to the tissue is dependent on the power administered (P) and the tissue resistance (R; I ² = P/R). Therefore, when ablation is performed in a power-controlled mode, resistance (impedance) has a direct and negative relationship to current; ablation at lower resistance yields higher current output (and tissue heating) compared with ablation at a similar power but higher resistance.

Increasing the size of the skin patch (or using two patches) and optimizing skin contact reduce electrical impedance at skin-patch interface (and minimize power dissipation) and provide for increased heating at the electrode-myocardial interface and thus improve ablation efficiency and increase lesion size. Additionally, the distribution of current density is skewed by the location of the grounding pad, and the geometric relationship between catheter, myocardium, and grounding pad can potentially affect lesion size. Experimental studies showed that RF applications performed at system low impedance values (80 Ω) resulted in significantly larger lesions when compared with higher impedance (120 Ω) lesions.

Radiofrequency system polarity

Most RF lesions are created by applying energy in a unipolar fashion between an ablating electrode touching the myocardium and a grounded reference patch electrode placed externally on the skin. As noted previously, bipolar energy delivery produces larger lesions than unipolar delivery. In general, the unipolar configuration creates a highly localized lesion, with the least amount of surface injury (i.e., narrower but deeper lesions), while bipolar RF ablation produces longer but shallower lesions.

Tissue temperature monitoring

RF lesion creation is thermally mediated. Maintaining tissue temperatures during ablation in a “safe and effective” zone (50–80°C) helps achieve adequate RF lesion and prevent steam pops. Therefore, directly measuring temperature at multiple points within the myocardium is the optimal method to monitor lesion formation during RF energy delivery. Direct monitoring of tissue temperature enables titration of RF power output, duration of energy application, and active cooling to optimize lesion formation.

Unfortunately, technologies to directly measure tissue temperature during RF ablation (which is the primary determinant of lesion formation) are currently not available for clinical use. Novel technologies are being evaluated, including microwave radiometry, near-field ultrasound thermal strain imaging, and magnetic resonance thermometry.

Electrode temperature monitoring

Irrigated radiofrequency ablation

The discrepancy between monitored electrode temperature and tissue temperature is significantly exaggerated by active cooling of the ablation electrode ( Fig. 8.3 ). The thermal effects on the electrode temperature are dependent on electrode heating from the tissue, internal cooling by the irrigation fluid, and external cooling from blood flow or open irrigation. With high irrigation flow rates, catheter tip temperature is no longer representative of tissue temperature and, therefore, feedback cannot be used to guide power output. The difference between the electrode temperature and tissue-electrode interface temperature is greater with the closed-loop electrode than with the open-irrigation electrode. The discrepancy is likely to be increased in areas of high blood flow, by increasing the irrigation flow rate, or by cooling the irrigant. Saline-irrigated catheters result in peak tissue heating several millimeters from the electrode-tissue interface. Because maximum tissue heating does not occur at the electrode-tissue interface, the value of temperature and impedance monitoring is limited with this type of catheter.

Therefore, monitoring lesion formation and optimizing power delivery during cooled RF ablation remain challenging. Appropriate energy titration is important to allow greater power application and to produce large lesions while avoiding overheating of tissue with steam formation leading to pops. Moreover, the inability to assess tissue heating, and hence to titrate power to an objective endpoint, prevents the operator from determining whether unsuccessful applications are caused by inadequate mapping or inadequate heating. In general, temperatures exceeding 42°C to 45°C with power greater than 30 W during open-irrigation RF ablation can be associated with a greater risk of steam pops and impedance rises, particularly during long RF applications, exceeding 60 seconds. Steam pops are often, but not always, audible. A sudden decrease in temperature, sudden catheter movement (as a consequence of the pop blowing the catheter out of position), and a sudden change in impedance are all potential indications that a pop has occurred.

Although the optimal method for adjusting power during saline-irrigated RF ablation is not yet clearly defined, some useful guidelines have emerged. The most commonly recommended approach is to perform ablation in a power-controlled mode, typically starting at 20 to 30 W and gradually increasing power to achieve evidence of tissue heating or damage. An impedance fall likely indicates tissue heating, similar to that observed with conventional RF ablation. When catheter temperature is between 28 and 31°C, power can be ramped up, watching for a 5- to 10-Ω impedance fall. Measured electrode temperature will generally increase, and electrode temperatures of 37°C to 40°C are commonly achieved. Power levels typically used during open-irrigation ablation depend on the site of ablation: 25 to 30 W in the left and right atrial free wall, 35 to 40 W for CTI and mitral isthmus ablation, 50 W in the left ventricle (LV), and 20 W in the coronary sinus (CS).

Temperatures exceeding 42°C to 45°C with power greater than 30 W during open-irrigation RF ablation can be associated with a greater risk of steam pops and impedance rises, particularly during long RF applications, exceeding 60 seconds. Steam pops are often, but not always, audible. A sudden decrease in temperature, a sudden catheter movement (as a consequence of the pop blowing the catheter out of position), and a sudden change in impedance are all potential indications that a pop has occurred. Whether the catheter is maintained in a stable position, as opposed to dragging it across the tissue during linear ablation, also likely influences tissue heating. High power can be applied continuously during dragging with little risk of excessive heating, although the duration of time to spend at each site to create an effective lesion may be difficult to ascertain. As a rule, the lowest effective power setting, shortest duration, and fewest applications should be employed whenever possible.

During open-irrigation RF ablation, initiation of irrigation results in a drop of electrode temperature by several degrees. Failure of electrode temperature to decrease indicates a lack of irrigant flow. When power delivery begins, catheter tip temperature should rise to 36°C to 42°C (the presence of rising temperature, not the magnitude, reflects tissue heating). Temperatures higher than 40°C achieved with low power (less than 20 W) can indicate that the electrode is in a location with little or no cooling from the surrounding circulating blood or that there is a failure of the catheter cooling system that requires attention. In contrast, the absence of any increase in tip temperature should raise the possibility of poor catheter contact.

With the internally irrigated system, the room temperature irrigant flowing at 36 mL/min typically cools the measured electrode temperature to 28°C to 30°C. During RF application, the electrode temperature increases; temperatures of 50°C can indicate that cooling is inadequate or has stopped, which warrants termination of RF application. The measured impedance typically decreases during cooled RF ablation by 5 to 10 Ω, in a manner similar to that observed during standard RF ablation.

New externally irrigated ablation catheter designs (Q-Dot MICRO catheter, and DiamondTemp ablation catheter) place multiple thermocouples at the surface of the ablation electrode tip, rather than internally, which are shielded from the irrigation fluid, allowing measurement of temperature at the catheter-tissue interface and, hence, enabling the use of irrigated ablation catheters in a temperature-controlled mode ( Fig. 8.9 ). However, as discussed above, even reliable recording of tissue surface temperature is unlikely to be sufficient for monitoring RF lesion formation, given the fact that maximal tissue heating occurs subendocardially.

Generator impedance monitoring

The magnitude of the current delivered by the RF generator used in ablation is largely determined by the impedance between the ablation catheter and the dispersive electrode ( Fig. 8.10 ). This impedance is influenced by several factors, including intrinsic tissue properties, catheter contact pressure, catheter electrode size, dispersive electrode size, presence of coagulum, and body surface area. Impedance measurement does not require any specific catheter-based sensor circuitry and can be performed with any catheter designed for RF ablation.

As tissue temperature rises during RF energy application, ions within the tissue being heated become more mobile, resulting in a decrease in impedance to current flow. Hence, a drop in impedance resulting from RF ablation can serve as a real-time marker of tissue heating. Although the impedance drop during RF ablation occurs mainly because of a reversible phenomenon, rather than from an irreversible myocardial tissue damage secondary to ablation, lesion diameter and depth have been shown to correlate well with impedance decrease, with an even more direct relationship than measured temperature. If the myocardial interface is not adequately heated—regardless of whether such is a result of poor tissue contact, lack of catheter stability, or inadequate power delivery—the characteristic decreases in impedance measurements for a given catheter location do not occur.

Typically, the impedance associated with firm catheter contact (before tissue heating has occurred) is 90 to 120 Ω. When catheter contact is poor, the initial impedance is 20% to 50% less, because of the lower resistivity of blood. Moreover, larger electrodes have larger contact area and consequently lower impedance. A 5- to 10-Ω reduction in impedance is usually observed in clinically successful RF applications, it correlates with a tissue temperature of 55°C to 60°C, and it is rarely associated with coagulum formation. Larger decrements in impedance reflect excessive tissue heating and are noted when coagulum formation is imminent. Once a coagulum is formed, an abrupt rise in impedance to more than 250 Ω is usually observed.

To titrate RF energy using impedance monitoring alone, the initial power output is set at 20 to 30 W and is then gradually increased to target a 5- to 10-Ω decrement in impedance. When target impedance is reached, power output should be manually adjusted throughout the RF application to maintain the impedance in the target range. A larger decrement of impedance should prompt reduction in power output. Lack of impedance drop during RF energy application can reflect inefficient energy delivery to the tissue due to poor catheter-tissue or catheter instability and should prompt catheter repositioning and verification of adequate tissue contact.

It is important to note that although coagulum formation is usually accompanied by an abrupt rise in impedance, the absence of impedance rise during ablation does not guarantee the absence of coagulum formation on the tissue contact site, which can unnoticeably be created on the tissue surface rather than on the ablation electrode. In addition, with large ablation electrodes, formation of blood clots may only minimally affect electrode impedance by covering a much smaller part of the electrode surface. Any increase in impedance during RF application can, however, indicate the beginning of coagulum formation or unintended catheter movement; in either case, RF application is discontinued.

The drop in generator impedance as a monitoring tool has several limitations. Generator impedance represents a global impedance value measured in all structures in series between the catheter tip and the indifferent body-surface electrode (skin patch). As such, generator impedance can be influenced by the electrical properties of not only the myocardium but also skin, subcutaneous tissue, lungs, and other thoracic structures. The large contribution of these variables dwarfs the relatively low magnitude of generator impedance change from ablation itself, thereby limiting its sensitivity to small drops in local impedance at the catheter tip resulting from RF effects and its clinical utility for real-time monitoring of tissue heating and lesion formation.

Local impedance monitoring

Novel ablation catheters (IntellaNav MIFI OI and IntellaNav StablePoint, Boston Scientific, Cambridge, MA, USA) capable of measuring local impedance at the distal electrode of the catheter (without a reference skin electrode) have recently been developed.

Similar to generator impedance, local impedance falls during RF-induced tissue heating. However, unlike generator impedance, local impedance measurement is based on the near electric field generated at the catheter tip and, therefore, it is more specific to the near-field myocardium beneath the catheter and more immune to respiratory/thoracic changes.

In vitro and in vivo animal models have shown that the magnitude of and rate of drop in local impedance during RF ablation dynamically followed the rate of intramural tissue (at 2-mm depth) temperature rise and strongly predicted lesion size ( Fig. 8.11 ). Additionally, local impedance drop indicated volumetric tissue heating with high fidelity to intralesion temperature across a range of delivered power. Importantly, models that incorporated local impedance drop during RF ablation showed a superior fit to lesion dimension data compared to models that incorporated other parameters such as force-time integral or generator impedance changes.

Recent data demonstrate that local impedance is potentially a valuable indicator of intralesion temperature and can provide valuable information on the catheter-tissue coupling, tissue characteristics, and lesion formation during RF catheter ablation. Both the absolute and percentage local impedance drops were significantly greater at the successful RF lesions than at unsuccessful lesions during ablation in the atrium and ventricle. Further, a change in local impedance during ongoing ablation was found to be more sensitive than that in generator impedance; drop in local impedance was more prominent than that observed in generator impedance at successful RF ablation points ( Video 8.1 ). Additionally, excessive drop in local impedance during RF ablation often portended excessive tissue heating and steam pops.

Recent evidence from experimental and initial clinical studies suggests that local impedance can potentially allow operators to monitor RF lesion formation. A local impedance drop of approximately 15 Ω during focal RF energy application can potentially be used as an indicator of adequate tissue heating and effective ablation lesion, while large drops (>35 Ω) can portend steam pops. However, the target local impedance drop values for optimal lesion formation for a variety of arrhythmias will need to be verified in future studies. ^,

Electrophysiologic effects of ablation

Effects of tissue heating on the recorded electrograms or the arrhythmia are important indicators for monitoring lesion formation. Interruption of tachycardia (VT, atrial flutter [AFL], supraventricular tachycardia) or block in conduction over a pathway (e.g., AVN or bypass tract [BT]) during the RF energy delivery provides immediate feedback about the disruption of tissue integrity. Additionally, an increase in local tissue pacing threshold and an attenuation in local electrogram amplitude can indicate tissue damage.

These factors, however, may not be easily monitored in real-time during the RF application, particularly the change in pacing threshold. Moreover, the decrease in electrogram amplitude is often not visible during RF application because of superimposed electrical artifact. Furthermore, reductions in amplitude and steepness of the local electrogram as indicators of tissue heating apply only to the unipolar distal electrogram. With a bipolar recording, the signal from the ring electrode may dominate the electrogram, and the bipolar amplitude can theoretically even rise during ablation because of a greater difference between the signals from both electrodes. On the other hand, the development of a monophasic R-wave (i.e., elimination of the negative deflection) in the unipolar electrogram can indicate transmurality in atrial ablations.

It is important to recognize that EP effects alone are not sufficient at ensuring an enduring RF lesion, since loss of EP function typically occurs well before cell necrosis and, hence, can be related to reversible cell injury. Experimental studies showed that reversible loss of cellular excitability occurs at tissue temperatures ranging between 43°C and 52°C (median, 48°C), whereas irreversible loss of cellular excitability and tissue injury develop at temperatures ≥50°C. Therefore, loss of EP function is inadequate to discriminate between reversible and irreversible myocardial tissue damage.

Input-based real-time ablation metrics

Effective delivery of RF energy to cardiac tissue is a balance of many factors including applied power, duration, tissue contact, flow rate, and tissue resistivity. In the absence of accurate and reliable real-time assessment of lesion development and transmurality, surrogate measures of lesion quality are commonly utilized.

The routine use of CF-sensing catheters now allows optimization of the catheter-tissue interface. However, the use of CF-guided ablation as a predictor of adequate tissue heating has not translated to the expected improved outcomes. Although CF is an important determinant of the lesion size during RF application, adequate CF is only one of several factors associated with lesion creation (in addition to ablation power, duration, contact surface area, and other factors), and when measured in isolation may not predict effective and durable lesions. Furthermore, measurement of average CF alone may actually be misleading when catheter stability or tip orientation is poor.

Subsequently, CF was used in conjunction with RF application time (force-time integral, FTI) as a better indicator of adequate lesion formation. As such, FTI considers not only the catheter-tissue contact but also the period of time for which this is applied, thereby providing a measure of the spatiotemporal stability of contact. However, as a marker of lesion quality, FTI has significant limitations. First, FTI does not account for power delivery, which is a modifiable ablation parameter and a critical determinant of lesion size. Second, FTI assumes a linear association between CF and time, but the actual interplay between CF and RF duration is dynamic and their relative contributions to lesion formation are more complex than the simple multiplication implied in the FTI formula. Third, the RF lesion in normal tissue is predominantly generated during the first 10 seconds of RF ablation, and it usually reaches its maximum size after 30 seconds once the thermodynamic equilibrium is achieved. Therefore, resistive heating does not lead to indefinite lesion growth with time and, contrary to implications of FTI, RF application with constant force and power for a longer period of time at the same location will not be more effective.

Several algorithms have been developed to overcome the limitations of CF and FTI by incorporating CF in conjunction with other metrics (such as catheter orientation with tissue and force directionality, RF power and time, impedance fall, lesion continuity and contiguity) to optimize lesion formation ( Fig. 8.12 ).

The Ablation Index (AI) is a marker of lesion quality that incorporates CF, RF power, and time in a weighted logarithmic formula. The AI has been shown to reliably predict lesion depth in animal models and has been proposed to allow intraprocedural guidance of RF application and improve outcomes in patients undergoing AF ablation. Nevertheless, the paucity of randomized controlled trials on AI-guided AF ablation makes it challenging to draw strong conclusions regarding its incremental benefit over conventional strategies. Moreover, the best AI targets and ablation power for an ideal ablation lesion remain to be identified. Low to moderate quality of evidence from nonrandomized data suggests that AI-guided PVI is associated with significantly lower procedure and ablation times, more frequent first-pass PVI and lower rate of acute PV reconnection, and lower incidence of AF recurrence, while it has safety profile comparable to non-AI catheter ablation. These studies used different ablation power and AI targets; the most commonly used energy was 35 W to the anterior and roof segments, and 30 W to the posterior and inferior segments. The most commonly used AI target was 500 to 550 for the anterior segment and 400 to 450 for the roof, posterior and inferior segments. High-power, short-duration lesions using up to 50 W with AI targets of 550 on the anterior wall and 400 on the posterior wall can be effective in achieving PVI safely and efficiently. ^{, ,}

Lesion size index (LSI) is a multiparametric index that incorporates CF, time, power, and impedance data recorded during RF ablation in a weighted formula. Ablations at fixed levels of CF, RF power, and duration but different impedances would all generate the same AI value, whereas LSI values would be different according to different levels of RF current. LSI was found to be highly predictive of RF lesion size in the in vitro model. Recent reports suggest that, during PVI, an LSI of 5.2 might be a suitable target for an effective lesion formation and an LSI of 4.0 might be acceptable on the posterior wall, especially in areas adjacent to the esophagus. However, confirmation by large clinical studies is warranted. ^,

It is important to understand that while FTI, AI, and LSI provide some means to standardize energy delivery tailored to tissue contact and can potentially improve point-by-point RF ablation, these lesion indexing algorithms assume that all tissues have a uniform response to RF dose (i.e., a prescribed power output, duration, CF, and contact stability); however, heterogeneity of tissue structure, thickness, disease state, and regional differences in blood flow challenge the uniformity assumption. Those differences in myocardial composition and conductance, in addition to effects of electrode contact angle, contact surface, or tissue-electrode interface geometry, are critically important to the extent of resistive heating and tissue response to RF ablation and are not accounted for in the input-based ablation lesion metric indices.

Recorded electrograms

Local electrogram characteristics (sharp, near-field electrogram versus far-field electrogram), beat-to-beat stability of recorded electrograms, and pacing capture threshold provide some information about tissue contact. However, electrograms only provide transient contact information and have been shown to be imprecise in judging contact. Also, assessing contact with abnormal myocardial tissue (e.g., infarct scar) can be difficult, since local electrograms are frequently low amplitude and local capture may not be achievable.

Fluoroscopy

Fluoroscopy can be useful to assess contact by monitoring catheter shaft and tip movement relative to the heart. However, utilization of fluoroscopy for this purpose significantly increases radiation exposure to both the patient and operator. Additionally, fluoroscopic monitoring is difficult to titrate and is not helpful for avoiding the application of excess catheter pressure against the cardiac wall.

Tactile feedback

Tactile feedback is often used as an indirect surrogate to catheter contact with the chamber wall. Tactile feedback, however, is subjective, a function of the operator’s experience, and is difficult to titrate. The use of steerable introducer sheaths can help improve catheter stability and contact, but such a technique often dampens the tactile feedback.

Data suggest that relying on only tactile feedback alone is not sufficient, and stronger effort should be made to gather as much information as possible about contact from as many different sensing modalities as possible. Furthermore, the use of three-dimensional (3-D) mapping systems alone, without the aid of fluoroscopy, to supplement tactile feedback may not be adequate.