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 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. 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). Their clinical application for catheter-based ablation has not yet been evaluated.^2,3

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 that electrical energy becomes dissipated as heat. Therefore, with excessive electrical resistance in the transmission line, the line actually warms up and power is lost. Current 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.^2,3

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 in contact with the electrode. The distribution between both depends on the impedance of both routes and also on how much electrode surface contacts blood versus endocardial wall. Whereas tissue heating is the target 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.^2,4

After leaving the electrode-blood-tissue interface, the current flows through the thorax to the indifferent electrode. Part of the RF power is lost in the patient’s body, including the area near the patch. Dissipation of energy can occur at the dispersive electrode site (at the contact point between that ground pad and the skin) to a degree that can limit lesion formation. In fact, if ablation is performed with a high-amplitude current (more than 50 W) and skin contact by the dispersive electrode is poor, it is possible to cause skin burns (Fig. 7-1).¹ Nevertheless, because the surface area of the ablation electrode (approximately 12 mm2) is much smaller than that of the dispersive electrode (approximately 100 to 250 cm2), the current density is higher at the ablation site, and heating occurs preferentially at that site, with no significant heating occurring at the dispersive electrode.⁵

FIGURE 7-1 Second-degree skin burns at the site of ablation ground pad contact to the patient’s thigh. Skin burns occurred secondary to poor contact between the skin and the pad during radiofrequency ablation of atrial fibrillation.

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.^2,3,5

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.^5,6

Tissue Heating

During alternating current flow, charged carriers in tissue (ions) attempt to follow the changes in the direction of the alternating current, 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. With a spherical electrode, the current flows outward radially, and current density therefore decreases with the square of distance from the center of the electrode. Consequently, 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.^2,3

Approximately 90% of all power that is delivered to the tissue is absorbed within the first 1 to 1.5 mm from the 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 as a result of heat conduction from this rim to the surrounding tissues. 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° and 90°C. However, whereas resistive heating starts immediately with the delivery of RF current, conduction of heat to deeper tissue sites is relatively slow and requires 1 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; however, deep tissue temperatures continue to rise for several seconds after interruption of RF delivery (the so-called thermal latency phenomenon).4 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 vessels determines the temperature profile within the tissue.^2,3

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. However, once the peak tissue temperature exceeds the threshold of 100°C, boiling of the plasma 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, which is accompanied by a sudden increase in electrical impedance that prevents further current flow into the tissue and further heating.^2,3

The range of tissue temperatures used for RF ablation is 50° to 90°C. Within this range, smooth desiccation of tissue can be expected. If the temperature is lower than 50°C, no or only minimal tissue necrosis results. 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.

Convective Cooling

The dominant factor opposing effective heating of myocardium is the convective heat loss into the circulating blood pool. 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 to 120 seconds). Therefore, the maximum lesion width is usually located intramurally, and the resultant lesion is usually teardrop shaped, with less necrosis of the superficial tissue.³

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 decreased 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, lesion size is reduced by such convective heat loss. On the other hand, when RF power delivery is not limited, convective cooling allows for more power to be delivered into the tissue, and higher tissue temperatures can be achieved (despite low temperatures measured by the catheter tip sensors), resulting in larger lesion size, without the risk of overheating and coagulum formation.^2,3

The concept of convective cooling can explain why there are few coronary complications with conventional RF ablation. Coronary arteries act as a heat sink; substantive heating of vascular endothelium is 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.⁵

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

Catheter Tip Temperature

Ablation catheter tip temperature depends on tissue temperature, convective cooling by the surrounding blood, tissue contact of the ablation electrode, electrode material and its heat capacity, and type and location of the temperature sensor.^2,7

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.^2,7

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; it can be significantly lower than tissue temperature.

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. 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. 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.^3,6–8

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 can be caused by changes in the cell membrane, protein inactivation, cytoskeletal disruption, nuclear degeneration, or other potential mechanisms.^2,3

Experimentally, the resting membrane depolarization is related to temperature. In the low hyperthermic range (37° to 45°C), little tissue injury occurs, and a minor change may 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° to 50°C), progressive depolarization of the resting membrane potential occurs, and action potential amplitude decreases. Additionally, abnormal automaticity is observed, reversible loss of excitability occurs, 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° to 51.5°C) and then permanent (at 51.7° to 54.4°C) conduction block develops, and fairly reliable irreversible myocardial injury occurs with a short hyperthermic exposure.³

In the clinical setting, the success of ablation is related to the mean temperature measured at the electrode-tissue interface. Block of conduction in an atrioventricular (AV) bypass tract (BT) usually occurs at 62° ± 15°C. During ablation of the AV junction, an accelerated junctional rhythm, which is probably caused by thermally or electrically induced cellular automaticity or triggered activity, is observed at temperatures of 51° ± 4°C, whereas reversible complete AV block occurs at 58° ± 6°C, and irreversible complete AV block occurs at 60° ± 7°C.

RF ablation typically results in high temperatures (70° to 90°C) for a short time (up to 60 seconds) at the electrode-tissue interface, but significantly lower temperatures at deeper tissue sites. This leads to 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 EP function can usually be demonstrated immediately after successful RF ablation, this finding can be delayed because tissue temperatures continue to rise somewhat after termination of RF energy delivery (thermal latency phenomenon). This effect can account for the observation that patients undergoing AV node (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 electrophysiological (EP) function after an initial successful ablation.⁵

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 on 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. 7-2). 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.²

FIGURE 7-2 Standard radiofrequency ablation catheter with char formed on the tip.

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 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. However, 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.³

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.²

Electrode Tip Temperature

It is important to understand that it is the amount of RF power delivered effectively into the tissue that determines tissue heating and thus lesion size, and the recorded catheter tip temperature is poorly correlated with lesion size. As noted, catheter tip temperature is always lower than the superficial tissue temperature. 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. In contrast, 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 Duration

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. Further extension of RF delivery during power-controlled RF delivery does not seem to increase lesion size further.

Electrode-Tissue Contact

The efficiency of energy transfer to the myocardium largely depends on electrode-tissue contact. An improvement in tissue contact leads to a higher amount of RF power that can be effectively delivered to the tissue. Consequently, the same tissue temperatures and lesion size can be reached at a much lower power level. However, at a certain moderate contact force, further increase in contact firmness results in progressively smaller lesions because a lesser amount of RF power is required to reach target temperature.⁴

Additionally, the surface area at the electrode-tissue interface influences the lesion size. When the catheter is wedged (i.e., between ventricular trabeculae or under a valvular leaflet), the electrode surface area exposed to tissue can be much higher. With half the electrode in contact with blood and the other half in contact with (twice) more resistive tissue, the amount of power delivered to blood is two times higher than the amount of power delivered to tissue (as compared with six times with 25% contact). The result is a more than twofold increase in tissue heating. In practice, however, power output rarely reaches 50 W in these situations; a temperature rise of the ablation electrode signals excessive tissue heating and limits power delivery.⁴

Close monitoring of changes in temperature and impedance provides information regarding the degree of electrode-tissue contact. Increased contact results in elevated initial impedance levels (by a mean of 22%) and an increase in the rate and level of impedance reduction, and the plateau is reached later. Additionally, increased electrode-tissue contact results in higher tissue temperatures, and the plateau is achieved later. Beat-to-beat stability of electrograms and pacing threshold also provide some information about tissue contact.

Electrode Orientation

When the catheter tip is perpendicular to the tissue surface, a much smaller surface area is in contact with the tissue (versus that exposed to the cooling effect of blood flow) than when the catheter electrode tip is lying on its side. Clinically, perpendicular electrode orientation yields larger lesion volumes and uses less power than parallel electrode orientation. However, if the RF power level is adjusted to maintain a constant current density, lesion size will increase proportionally to the electrode-tissue contact area, which is larger with parallel tip orientation. Lesion depth is only slightly affected by catheter tip orientation using 4-mm-long tip catheters, but lesions are slightly longer in the parallel orientation as compared with the perpendicular orientation. Moreover, 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 of 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.^2,3,10

Electrode Length

Ablation catheters have tip electrodes that are conventionally 4 mm long and are available in sizes up to 10 mm long (see Fig. 4-6). An increased electrode size reduces the interface impedance with blood and tissue, but the impedance through the rest of the patient remains the same. The ratio between interface impedance and the impedance through the rest of the patient is thus lower with an 8-mm electrode than with a 4-mm electrode, which reduces the efficiency of power transfer to the tissue. Therefore, with the same total power, lesions created with a larger electrode are always smaller than lesions created with a smaller electrode (Fig. 7-3). 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.⁴

FIGURE 7-3 Effect of electrode size and power delivery on lesion size. Representative blocks of myocardium are shown with lesions (dark shading) resulting from optimal radiofrequency energy delivery for each electrode type. The effect of 20 W delivered for 30 seconds is shown for a 4-mm electrode (left) and an 8-mm electrode (center). The larger electrode dissipates more of the energy into the blood pool and thus paradoxically creates a smaller lesion. Because of electrode cooling from its larger surface area, the same 8-mm electrode can deliver more power than the 4-mm electrode and could create a larger lesion (right).

When the power is not limited, however, catheters with large distal electrodes create larger lesions, both by increasing the ablation electrode surface area in contact with the bloodstream, thus 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 (see Fig. 7-3). 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.^2,3

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 may be 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 intramural gas explosions and blood clot formation. Another point of concern is that the formation of blood clots 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.⁴

The principal limitations of a large ablation electrode (8 to 10 mm in length) are 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, thus making it more difficult to identify the optimal ablation site. A larger electrode dampens the local electrogram, especially that of the distal electrode. With an 8- or 10-mm long distal and a 1-mm short proximal ring electrode, the latter can be the main source for the bipolar electrogram; this then confuses localization of the optimal ablation site. 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.⁴

Ablation Electrode Material

Although platinum-iridium electrodes have been the standard for most RF ablation catheters, gold exhibits excellent electrical conductive properties, as well as a 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 a higher mean rate of power 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.¹¹ Therefore, gold electrodes allow for greater power delivery to create deeper lesions at a given electrode temperature without impedance increases.¹² Enhanced electrode cooling allows for more RF power to be applied at constant temperature, before the temperature limit is reached or before the impedance of the electrode rises.⁴ However, the higher thermal conductivity of gold electrodes is no longer an advantage in areas of low blood flow (e.g., among 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.^3,11

Conflicting results were observed in clinical studies comparing 8-mm gold-tip with platinum-iridium–tip catheters for ablation of the cavotricuspid isthmus.¹³ 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. However, ablation with gold electrodes seemed to be safe and well tolerated and specifically did not increase the risk of AV block. 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.¹¹

Reference Patch Electrode Location And Size

The RF current path and skin reference electrode interface present significant impedance for the ablation current flow, thereby dissipating part of the power. Increasing patch size (or using two patches) provides for increased heating at the electrode-endocardium interface and thus increases ablation efficiency and increases lesion size. On the other hand, the position of the dispersive electrode (on the patient’s back or thigh) has little effect on the size of the resulting lesion.⁵

Blood Flow

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 lesion size is primarily dependent on the RF power delivered to the tissue, lesion size varies with the magnitude of local blood flow. 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 cavotricuspid isthmus, dilated and poorly contracting atria, and dilated and poorly contracting ventricles). 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° or 70°C only minimally increases RF power and increases the risk of thrombus formation and impedance rise. In contrast, increasing local blood flow is associated with increased convective cooling of the ablation electrode. Consequently, more power is delivered to the tissue to reach and maintain target temperature, thus resulting in larger lesion volumes.^2,3,7

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. The unipolar configuration creates a highly localized lesion, with the least amount of surface injury. Energy can also be applied in a bipolar mode between two endocardial electrodes. Bipolar energy delivery produces larger lesions than unipolar delivery. A novel ablation system (GENius, Ablation Frontiers, Carlsbad, Calif.) has been developed for ablation of atrial fibrillation (AF), in which duty-cycled and alternating unipolar and bipolar (between adjacent electrodes) energy RF ablation is used to create contiguous lesions in the pulmonary vein (PV) antrum using a ring catheter with 10 electrodes (Pulmonary Vein Ablation Catheter [PVAC], Ablation Frontiers, Medtronic Inc., Minneapolis, Minn.) (Figs. 7-4 and 7-5).¹⁴ Other catheter designs, which use the same multichannel duty-cycled RF generator, also have been developed for left atrial ablation of fractionated electrograms (Multi-Array Ablation Catheter [MAAC], Ablation Frontiers), septal ablation (Multi-Array Septal Catheter [MASC], Ablation Frontiers), and linear ablation (Tip-Versatile Ablation Catheter [TVAC], Ablation Frontiers) (see Fig. 7-4).¹⁵ One advantage of this approach is the simultaneous application of RF energy across an electrode array, intended to create contiguous lesions near an anatomical structure. Additionally, the multielectrode catheters allow selective mapping and ablation through any or all electrodes as required. This system is still undergoing investigation. Another novel ablation catheter system delivers RF energy through a flexible mesh electrode (C.R. Bard, Inc., Billerica, Mass.) that is deployed in the PV ostium. This system is undergoing investigation for PV isolation.¹⁶

FIGURE 7-4 Novel multipolar electrode catheters (Ablation Frontiers, Inc., Carlsbad, Calif.) using duty-cycled radiofrequency. A, The Pulmonary Vein Ablation Catheter (PVAC) is a multielectrode catheter used to map, ablate, and verify isolation of the pulmonary veins. B, The Multi-Array Ablation Catheter (MAAC) is a multielectrode catheter designed to map and ablate complex fractionated atrial electrograms in the left atrial body. C, The Multi-Array Septal Catheter (MASC) is a multielectrode catheter designed to map and ablate complex fractionated atrial electrograms along the left atrial septum. D, The Tip-Versatile Ablation Catheter (TVAC) is a multielectrode catheter for linear ablation.

(Courtesy of Medtronic, Inc., Minneapolis, Minn.)

FIGURE 7-5 Ablation Frontiers (Medtronic Inc., Minneapolis, Minn.) Pulmonary Vein Ablation Catheter (PVAC) in the left inferior pulmonary vein. The electrodes of the PVAC catheter, which is deployed over a guidewire into the antrum of a pulmonary vein, can be used for both recording and ablating around the perimeter of the vein. CS = coronary sinus; Eso = esophageal catheter; LAO = left anterior oblique; RAO = right anterior oblique.

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. 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. Larger dispersive electrodes and larger ablation electrodes result in lower impedance.^3,7

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.

The impedance drop during RF ablation occurs mainly because of a reversible phenomenon, such as tissue temperature rise, rather than from an irreversible change in tissues secondary to ablation of myocardial tissue. Therefore, impedance provides a useful qualitative assessment of tissue heating; however, it does not correlate well with lesion size.⁴ A 5- to 10-Ω reduction in impedance is usually observed in clinically successful RF applications, it correlates with a tissue temperature of 55° to 60°C, and it is rarely associated with coagulum formation. Larger decrements in impedance 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.^3,7

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, as needed, to maintain the impedance in the target range. A larger decrement of impedance should prompt reduction in power output.

The drop in impedance as a monitoring tool has several limitations. When blood flow rates are low, blood can also be heated, and electrode impedance drops accordingly. Moreover, a large rise in tissue temperature at a small contact area and a smaller rise with better tissue contact can result in a similar drop in impedance. Inversely, similar tissue heating with different tissue contact can result in a different change in impedance. In addition, resistive heating nearby is fast, whereas conductive heating to deeper layers is relatively slow. The former, at close distance, has a much greater effect on impedance than the latter, which occurs at greater distance. Like electrode temperature rise, the drop in impedance during RF application is not a reliable parameter for estimating tissue heating and lesion growth.^3,4,7

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 blood clot formation on the tissue contact site, which can unnoticeably be created on the lesion surface. In addition, as noted, 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 thrombus formation or unintended catheter movement; in either case, RF application is discontinued.

Temperature Monitoring

Monitoring of catheter tip temperature and closed-loop control of power output are useful to avoid excessive heating at the tissue surface, which can result in coagulum formation, and to accomplish effective heating at the target area. However, catheter tip temperature is affected by cooling effects and electrode-tissue contact and thus poorly correlates with lesion size. Tissue temperature can be markedly higher than catheter tip temperature; a higher target temperature can increase the incidence of tissue overheating associated with crater formation and coagulum formation. In high-flow areas, the tip is cooled and more RF power is delivered to the tissue to reach target temperature, thus resulting in relatively large lesions and vice versa.⁷

As mentioned earlier, temperature monitoring requires a dedicated sensor within the catheter. Two types of sensors are available: thermistor and thermocouple. No catheter or thermometry technology has been demonstrated to be superior in clinical use; however, closed-loop control of power output is easier to use than manual power titration. 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, and it is likely that the measured temperature underestimates the peak tissue temperature. Therefore, it is best if target temperatures no higher than 70° to 80°C are selected in the clinical setting.^3,7

Titration of RF energy using temperature monitoring is usually done automatically by a closed-loop temperature monitoring system. When manual power titration is directed by temperature monitoring, the power initially is set to 20 to 30 W and then is gradually increased until the target temperature is achieved. With both manual and automatic power titration, change in power output is frequently required throughout the RF application to maintain the target temperature. Application of RF energy is continued if the desired clinical effect is observed within 5 to 10 seconds after the target temperature or impedance is achieved. If the desired endpoint does not occur within this time, the failed application is probably because of inadequate mapping. If the target temperature or impedance is not achieved with maximum generator output within 20 seconds, the time it takes to achieve subendocardial steady-state temperature, the RF application can be terminated, and catheter adjustment should be considered to obtain better tissue contact.⁷

The target ablation electrode temperature varies according to the arrhythmia substrate. For AVNRT, target temperature is usually 50° to 55°C. For BT, AV junction, atrial tachycardia (AT), and ventricular tachycardia (VT), a higher temperature (55° to 60°C) is usually targeted.

When using 4-mm-tip catheters, the target temperature should be lower than 80°C. In high-flow areas in the heart, the disparity between tip temperature and tissue temperature is large and a lower target temperature should be considered (e.g., 60°C), whereas in low flow areas tissue temperature is much better reflected by tip temperature and a higher target temperature can be considered (e.g., 70° to 80°C). The duration of RF application can be limited to 30 seconds for nonirrigated 4-mm-tip electrodes. The lesion is predominantly formed within the first 30 seconds. A longer duration does not create larger lesions.

When using 8-mm-tip catheters, a larger portion of the ablation electrode is exposed to the blood and thus cooled by blood flow, and a relatively large difference between catheter tip temperature and tissue temperature can be expected. Consequently, a moderate target temperature (e.g., 60°C) should be chosen; the RF power may be limited to 50 to 60 W to avoid tissue overheating and coagulum formation.³

It is important to recognize that prevention of coagulum formation is difficult, even with temperature and impedance monitoring. The clot first adheres to the tissue because that is the site with the highest temperature and may only loosely attach to the cooler electrode. The denaturized proteins probably have higher electrical impedance than blood, but the contact area with the electrode can be small, and RF impedance may not rise noticeably. The absence of flow inside the clot and its presumed higher impedance accelerate local heating and, because of some contact with the electrode, also accelerate heating of the electrode. Desiccation and adherence to the electrode then lead to coagulum formation on the metal electrode and impedance rise. Automatic power reduction by temperature-controlled RF ablation compensates for the reduction of electrode cooling and may prevent desiccation and impedance rise. The clot, however, can still be formed as demonstrated by experimental in vivo studies, and this can remain unnoticeable until it detaches from the tissue. Therefore, the absence of thermal and electrical phenomena does not imply that the ablation has been performed safely.⁷

Ultrasound Imaging

Ultrasound imaging can allow assessment of tissue heating and pops. The presence of microbubbles on intracardiac or transesophageal echocardiography is typically associated with a tissue temperature higher than 60°C and increased lesion size, and continued RF application after the appearance of the bubbles is usually followed by an increase in impedance. Moreover, pops are not always audible but can be seen well on ultrasound, often with a sudden explosion of echocardiographic contrast.⁶ Microbubble formation, however, is not a straightforward surrogate for tissue heating. The absence of microbubble formation clearly does not indicate that tissue heating is inadequate or that the power level should be increased, nor does the presence of scattered microbubbles indicate safe tissue heating. This marker is fairly specific for tissue heating as judged by tissue temperatures but is not routinely sensitive. Specifically, scattered microbubbles can occur over the entire spectrum of tissue temperatures, whereas dense showers of microbubbles appear only at tissue temperatures higher than 60°C. Scattered microbubbles can represent an electrolytic phenomenon, whereas dense showers of microbubbles suggest steam formation, with associated tissue disruption and impedance rises.⁶

Mechanism

Excessive surface heating invites coagulum formation, carbonization, and steam popping. These adverse effects may limit the depth of RF lesions, thus making it difficult to produce lesions of sufficient depth in scar tissue or thickened ventricular walls. Such limitations of conventional RF systems have stimulated the evaluation of modified electrode systems. One important modification involves cooling of the ablation electrode, which was designed to prevent overheating of the endocardium while allowing sufficient energy delivery to achieve a larger lesion size and depth.¹⁷

There are two methods of active electrode cooling by irrigation: internal and external (Fig. 7-6). With the internal (closed-loop) system (Chilli, Boston Scientific, Natick, Mass.), cooling of the ablation electrode is performed by circulating fluid within the electrode.¹⁸ In contrast, with the external (open-loop) system (Celsius or Navistar ThermoCool, Biosense Webster, Inc., Diamond Bar, Calif.; and Therapy Cool Path, St. Jude Medical, Inc., St. Paul, Minn.), electrode cooling is performed by flushing saline through openings in the porous-tipped electrode (showerhead-type system).¹⁸ Another cooling system is sheath-based open irrigation, which uses a long sheath around the ablation catheter for open irrigation. The latter system was found to provide the best results, but this type of catheter tip cooling is not clinically available.

FIGURE 7-6 Schematic representation of irrigated electrode catheters. A, The closed-loop irrigation catheter has a 7 Fr, 4-mm-tip electrode with an internal thermocouple. B, The open-irrigated catheter has a 7.5 Fr, 3.5-mm-tip electrode with an internal thermocouple and six irrigation holes (0.4-mm diameter) located around the electrode, 1.0 mm from the tip. NaCl = sodium chloride.

(From Yokoyama K, Nakagawa H, Wittkampf FH, et al: Comparison of electrode cooling between internal and open irrigation in radiofrequency ablation lesion depth and incidence of thrombus and steam pop. Circulation 113:11, 2006.)

Active electrode cooling by irrigation can produce higher tissue temperatures and create larger lesions, compared with standard RF ablation catheters, because of a reduction in overheating at the tissue-electrode interface, even at sites with low blood flow. This allows the delivery of higher amounts of RF power for a longer duration to create relatively large lesions with greater depth but without the risk of coagulum and char formation. Unlike with standard RF ablation, the area of maximum temperature with cooled ablation is within the myocardium, rather than at the electrode-myocardial interface. Higher power results in greater depth of volume heating, but if the ablation is power limited, power dissipation into the circulating blood pool can actually result in decreased lesion depth (Fig. 7-7).¹⁸ Compared with large-tip catheters, active cooling has been shown to produce equivalent lesions with energy delivery via smaller electrodes, with less dependence on catheter tip orientation and extrinsic cooling, whereas larger electrodes have significant variability in their electrode-tissue interface, depending on catheter orientation (see Fig. 7-7). For the nonirrigated catheter, greater lesion volumes are observed with a horizontal orientation of the RF electrode compared with a vertical orientation. In contrast, for irrigated catheters, lesion volume increased with a perpendicular electrode orientation compared with the horizontal orientation.¹⁰

FIGURE 7-7 Effect of electrode type on lesion size. Representative blocks of myocardium are shown with lesions (dark shading) resulting from optimal radiofrequency energy delivery for each electrode type. A 4-mm electrode makes a small shallow lesion (left). An 8-mm electrode makes a larger, deeper lesion (center). The open-irrigated electrode makes a larger, deeper lesion, with a maximum width at some tissue depth (right).

Lesion depth seems to be similar between closed-loop and open-irrigation electrodes. However, open irrigation appears to be more effective in cooling the electrode-tissue interface, as reflected by lower interface temperature, lower incidence of thrombus, and smaller lesion diameter at the surface (with the maximum diameter produced deeper in the tissue). These differences between the two electrodes are greater in low blood flow, presumably because the flow of saline irrigation out of the electrode provides additional cooling of the electrode-tissue interface (external cooling). Ablation with the closed-loop electrode, with irrigation providing only internal cooling, in low blood flow frequently results in high electrode-tissue interface temperature (despite low electrode temperature) and thrombus formation.¹⁸

When maximum power and temperature parameters that did not result in increased rates of complications and any evidence of excessive heating (including popping, boiling, or impedance rise) for each catheter type were selected, closed-irrigation catheters resulted in slightly larger lesion volumes and greater lesion depths than the open-irrigation catheters. Both cooled catheters fared better than the standard 4-mm-tip and large 10-mm-tip catheters with larger lesions achieved within the range of safe energy delivery.¹⁹

Mechanism of Tissue Injury

The mechanisms underlying lesion formation by cryoenergy are twofold: direct cell injury and vascular-mediated tissue injury. The mechanisms of cellular death associated with tissue freezing involve immediate cellular effects, as well as late effects that determine the lesion produced.²²

Determinants of Lesion Size

During cryoablation, lesion size and tissue temperature are related to convective warming, electrode orientation, electrode contact pressure, electrode size, refrigerant flow rate, and electrode temperature.²⁸ Lesion sizes during catheter cryoablation can be maximized by use of larger ablation electrodes with higher refrigerant delivery rates. A horizontal electrode orientation to the tissue and firm contact pressure also enhance lesion size. For a given electrode size, electrode temperature may be a poor predictor of tissue cooling and lesion size. In contrast to RF ablation, cryoablation in areas of high blood flow can result in limited tissue cooling and smaller lesion sizes because of convective warming. Conversely, cryoablation lesion size can be maximized in areas of low blood flow.^22,28 The time to reach –80°C can predict cryoablation lesion size; cryoablation lesion volume increases as the time to reach –80°C decreases.²⁹ The 6- and 8-mm electrode-tip cryocatheters produce ablation lesions of similar depth that are more than two- and threefold larger than 4-mm catheters, respectively. Despite larger lesions, endothelial cell layers remain intact and devoid of thrombosis. Surface areas and volumes may be particularly sensitive to catheter tip-to-tissue contact angles with larger electrodes. As such, particular attention in catheter orientation is required with 8-mm electrode-tip cryocatheters to produce desired lesions.²⁷

Cryomapping

Cryomapping is designed to verify that ablation at the chosen site will have the desired effect and to ensure the absence of complications (i.e., to localize electric pathways to be destroyed or spared). This procedure is generally performed using various pacing protocols that can be performed during cryomapping (or ice mapping) at –32°C. At this temperature, the lesion is reversible (for up to 60 seconds), and the catheter is stuck to the adjacent frozen tissue because of the presence of an ice ball that includes the tip of the catheter (cryoadherence). This permits programmed electrical stimulation to test the functionality of a potential ablation target during ongoing ablation and prior to permanent destruction. It also allows ablation to be performed during tachycardia without the risk of catheter dislodgment on termination of the tachycardia.

In the cryomapping mode, the temperature is not allowed to drop to less than –30°C, and the time of application is limited to 60 seconds. Formation of an ice ball at the catheter tip and adherence to the underlying myocardium are signaled by the appearance of electrical noise recorded from the ablation catheter’s distal bipole. Once an ice ball is formed, programmed stimulation is repeated to verify achievement of the desired effect. If cryomapping does not yield the desired result within 20 to 30 seconds or results in unintended effect (e.g., AV conduction delay or block), cryomapping is interrupted, to allow the catheter to thaw and become dislodged from the tissue. After a few seconds, the catheter may be moved to a different site and cryomapping repeated.²²

Cryoablation

When sites of successful cryomapping are identified by demonstrating the desired effect with no adverse effects, the cryoablation mode is activated, in which a target temperature lower than –75°C is sought (a temperature of –75° to –80°C is generally achieved). The application is then continued for 4 to 8 minutes, thus creating an irreversible lesion. If the catheter tip is in close contact with the endocardium, a prompt drop in catheter tip temperature should be observed as soon as the cryoablation mode is activated. A slow decline in catheter tip temperature or a very high flow rate of refrigerant during ablation suggests poor electrode-tissue contact. In such cases, cryoablation is interrupted, and the catheter is repositioned.²²

Atrioventricular Nodal Reentrant Tachycardia

So far, slow pathway ablation for AVNRT by cryothermal energy represents the larger experience in the clinical application of this technology.^21,32 Previously published reports with cryoablation of AVNRT in pediatric patients demonstrated procedural success rates of 83% to 97% and recurrence rates ranging from 0% to 20%. Although the use of bonus cryoapplications to consolidate the acutely successful cryoablation and the availability of larger tip cryocatheters (6 to 8 mm and 6 mm versus 4 mm) to create larger lesions have been associated with fewer recurrences on long-term follow-up without compromising safety, the overall procedural success rates have remained consistently lower than those (95% to 99%) with RF ablation.^31,33–36

According to current data, cryothermal energy is a valuable and useful alternative to RF energy to treat patients with AVNRT. The absence of permanent inadvertent damage of AV conduction makes this new technology particularly useful for patients with difficult anatomy, after an unsuccessful prior standard ablation procedure, in pediatric patients, and in patients in whom even the small risk of AV block associated with RF ablation is considered unacceptable. Cryoablation can be of particular advantage in several situations, including posterior displacement of the fast pathway or AVN, a small space in the triangle of Koch between the HB and the CS ostium, and the need for ablation to be performed in the midseptum. However, given the high success rate and low risk of RF slow pathway ablation, it can be difficult to demonstrate a clinical advantage of cryoablation over RF ablation in unselected AVNRT cases.²²

Bypass Tracts

Cryothermal ablation of BTs in the superoparaseptal and midseptal areas, both at high risk of complete permanent AV block when standard RF energy is applied, is highly safe and successful (Fig. 7-10).³² Cryoablation can also be successfully and safely used to ablate selected cases of epicardial left-sided BTs within the CS, well beyond the middle cardiac vein, once attempts using the transseptal and transaortic approaches have failed.²¹ However, the experience with cryoablation in unselected BTs is more limited and less satisfactory; this can be related to multiple factors, including the learning curve and the smaller size of the lesion produced by cryoablation. In addition, all the peculiarities of cryothermal energy, which are optimal for septal ablation, are less important or even useless for ablation of BTs located elsewhere.²¹

FIGURE 7-10 Cryoablation of a para-Hisian bypass tract (BT). Right anterior oblique fluoroscopic view of cryoablation catheter (Cryo) position and ECG and intracardiac recordings during cryoablation of para-Hisian BT. The first two QRS complexes at left are preexcited, with the second QRS complex (blue arrow) less so than the first (red arrow), as cryoablation begins to affect pathway conduction. Preexcitation disappears afterward. The ablation recordings at bottom are severely disrupted by the delivery of cryoenergy. Abl_dist = distal ablation site; Abl_prox = proximal ablation site; Abl_uni = unipolar ablation site; CS = coronary sinus; CS_dist = distal coronary sinus; CS_prox = proximal coronary sinus; His_dist = distal His bundle; His_mid = middle His bundle; His_prox = proximal His bundle; HRA = high right atrium; RA = right atrium; RV = right ventricle; RVA = right ventricular apex.

In more recent series, the short-term success rate of cryoablation of BTs in the superoparaseptal and midseptal regions exceeded 90%; however, resumption of BT conduction with recurrence of symptoms can occur in up to 20% of patients, and overall success rates have been lower than those of AVNRT cryoablation. It is, however, important to note that alternatives for elimination of some of those BTs are limited by a prohibitive risk of AV block with RF ablation. Although transient modifications of the normal AVN conduction pathways can be observed during cooling and right bundle branch block has occurred on occasion, no permanent modifications have been observed, and inadvertent AV block has yet to be reported. In fact, immediate discontinuation of cryothermal energy application at any temperature on observation of modification of conduction over normal pathways results in return to baseline condition soon after discontinuation.³⁰

Focal Atrial Tachycardia

Occasionally, successful cryoablation of focal AT has been reported. Its safety has also been confirmed for ablation of atrial foci located close to the AVN. Cryoenergy can also be particularly valuable in ablating focal sources within venous structures.³⁰

Typical Atrial Flutter

Complete cavotricuspid isthmus block can be achieved by cryothermal ablation. In addition to the safety advantages of cryoenergy discussed earlier, cryolesions are limited by convective warming effects of blood flow, which reduces the risk of damage to nearby coronary arteries. Another advantage of using cryothermal energy for the ablation of typical AFL is the absence of pain perception related to energy application.^30,37 Previous studies reported cryoablation of the cavotricuspid isthmus for typical AFL with short-term and long-term success comparable to results with RF ablation. However, a more recent prospective randomized study comparing RF and cryothermal energy for ablation of typical AFL suggested that lesion durability from cryoablation was significantly inferior to that of RF ablation. Although acutely successful ablation rates in the cryoablation group were comparable to those for RF ablation (89% versus 91%), persistence of bidirectional isthmus block in patients treated with cryoablation reinvestigated 3 months following ablation was inferior to that in patients treated with RF ablation, as evidenced by the higher recurrence rate of symptomatic, ECG-documented AFL (10.9% versus 0%), and higher asymptomatic conduction recurrence rates (23.4% versus 15%). Additionally, compared with RF ablation, cryoablation is associated with significantly longer procedure times. This is driven mainly by differences in ablation duration, which can be attributed to the longer duration of each cryoablation (4 minutes) compared with RF ablation (up to 60 seconds).^22,38

Pulmonary Vein Isolation

For the characteristics mentioned earlier, cryothermal energy ablation can be considered an ideal and safer energy source for PV isolation, and the incidence of PV stenosis and thromboembolic events is expected to be dramatically reduced compared with RF ablation. Additionally, cryoablation has never been associated with atrioesophageal fistula.^39,40 On the other hand, cryothermal injury is sensitive to surrounding thermal conditions. The high flow of the PVs can present a considerable heat load to cryothermal technologies, which can limit the size and depth of the lesion produced by cryothermal energy at the ostium of the PV. Furthermore, cryoablation of PVs by standard steerable catheters is very time-consuming, considering that one single ablation point takes about 4 minutes.

A new catheter design for circumferential ostial ablation of the PVs has been developed, with the option of deploying an inflatable balloon in the PVs to reduce the heat load related to blood flow (Fig. 7-11).^21,39 The cryoballoon is deployed at the ostium of targeted PVs, and cryoenergy is delivered over the occluding balloon system to create circumferential lesions around the PV ostium. The clinical short-term and long-term success rates of this approach are acceptable; however, the risk of phrenic nerve injury is relatively high because of ablation in the region of the right superior PV. To date, thromboembolism or PV stenosis has yet to be reported.^{23–25,41–45}

FIGURE 7-11 Cryothermal balloon catheter.

(Courtesy of Medtronic, Inc., Minneapolis, Minn.)

Ventricular Tachycardia

Data regarding cryoablation for VT remain scant. Small reports described feasibility and success of cryoenergy for ablation of outflow tract VT. Additionally, cryothermal energy can be of potential advantage in percutaneous subxiphoid epicardial ablation of VT because of less potential damage to the epicardial coronary arteries. The reduced heat load in the pericardial space related to the absence of blood flow limits RF energy delivery but can be to the advantage of cryoablation, with the possibility of producing larger transmural lesions.²²