Radiofrequency (RF) ablation induces tissue necrosis by heating the targeted tissue. Most cardiac ablation systems deliver RF energy between a unipolar platinum-iridium electrode at the tip of the ablation catheter and a large dispersive grounding patch on the patient’s skin, typically on the abdomen or thigh ( Fig. 8.1 ). Ablation catheters also have electrodes to record electrical activity from within the heart ( Fig. 8.2 ). When the catheter is placed at the target site, alternating current (AC) in the RF range (500 to 1000 kHz) is applied to the tip electrode. As RF current passes from the tip of the ablation catheter to the grounding patch, resistive heating occurs in the intervening tissue at a rate proportional to the square of current density. Current density is highest in the immediate vicinity of the ablation catheter tip, which has a small surface area. As a result, significant resistive heating occurs only at the catheter tip–tissue interface and in a small volume of surrounding tissue. Deeper lesions result from conductive heating of myocardium adjacent to the site of resistive heating. Application of electric current of any frequency will result in tissue heating; however, higher frequencies (in the MHz range) are transmitted to tissues distant from the electrode, and heat dissipation can occur over a much larger tissue mass, whereas lower frequencies can lead to stimulation of cardiac muscle and nerves, resulting in arrhythmias and pain.
Heat generation in tissue
The electrode at the tip of the catheter is not insulated, and electric energy can travel into the targeted tissue surrounding the probe. Because the electrical energy is delivered in the form of alternating current, the electric field created within the tissue periodically reverses direction ( Fig. 8.3 ). This causes the ions in the tissue to try to follow this same alternating path, which results in ions drifting relative to one another and relative to neutral particles. These rapid movements (for a typical frequency of 500 kHz the direction of ion movement changes a million times per second) result in collisions between the particles that limit their drift velocities and convert some of their kinetic energy into thermal energy. This phenomenon is called frictional or resistive heating. Resistive heat production within the tissue is directly related to RF power density, which is proportional to the square of the current density (I ). When RF energy is delivered in a unipolar mode, the current distributes radially from the source. The current density decreases in proportion to the square of the distance (r ) from the RF electrode source. Thus direct resistive heating of the tissue decreases in relation with the distance from the electrode to the fourth power. , Hence, doubling the distance from the catheter tip reduces resistive heating by 94%. Because of the rapid reduction in heating with distance, lesions created by RF energy are typically small and well circumscribed. Deeper tissue heating occurs solely as a result of heat conduction from the narrow radius of volume heating around the electrode. Indeed, only a 1- to 2-mm rim of tissue directly adjacent to the catheter tip is heated resistively, whereas deeper tissues are heated by passive thermal conduction ( Fig. 8.4 ). This conductive heating is responsible for most of the lesion volume from RF ablation catheters.
The skin below the ground pad will also be heated, but because of the large pad surface area, temperature rises are low. A single application lasts typically 45 to 120 seconds and produces an ablation lesion 5 to 10 mm in diameter. The operator can control the size of the lesion by varying the time period and intensity of application, contact force, and the size of the ablation electrode. Generally there are three methods to control applied intensity : (1) power control, in which electrical voltage applied to the RF electrode is adjusted to keep applied RF power constant; (2) temperature control, in which one or more thermal sensors (either thermocouples or thermistors) are integrated in the RF electrode, typically near the tip, and the applied RF power is adjusted to keep the measured temperature at a defined target value; and (3) impedance control, in which RF power is adjusted depending on tissue impedance, which is measured between the RF electrode and the ground pad. Most ablation catheters employ temperature control, where temperature measured within the electrode tip is used to adjust applied RF power. The location of maximum tissue temperature is a few mm from the catheter tip and is affected by local intracardiac blood flow, which considerably affects tissue heating and resulting size of the ablation zone. In general, larger ablation zones are possible at locations with high flow rate because of increased convective cooling. A major consideration for RF ablation is that catheter tip temperature can differ substantially from underlying tissue temperature because of the complex relationship between RF heating of adjacent tissue and convective cooling from intracavitary blood flow. This can result in a measured catheter tip temperature that is significantly lower than true tissue temperature, especially with larger-tipped catheters, which are exposed to more convective cooling because of their larger surface area. In fact, tissue temperatures can exceed 100°C, even while a much lower temperature is recorded from the catheter tip. This can boil fluid within the tissue, resulting in an explosive steam pop, damage to adjacent structures, crater formation, cardiac perforation, and tamponade. Catheters of different lengths and diameters are commercially available depending on desired size of the ablation zone. Newer catheter designs employ internal cooling to increase ablation zone size.
When catheter ablation is performed under temperature control, RF power is regulated to maintain a constant electrode–tissue interface temperature (commonly 55°C or 60°C). The ablation electrode–tissue interface temperature is dependent on the opposing effects of heating from the tissue and cooling by the blood flowing around the electrode. At any given electrode–tissue interface temperature, the RF power delivered to the tissue is significantly reduced in areas of low blood flow. The reduced cooling associated with low blood flow causes the electrode to reach the target temperature at lower power levels. Because lesion size is primarily dependent on the RF power delivered to the tissue, lesion size will vary with the magnitude of local blood flow. Areas where lesion size is adversely affected because of low local blood flow (poor electrode cooling), increasing electrode–tissue interface temperature to 65°C or 70°C only minimally increases RF power and increases the risk of thrombus formation and impedance rise. , , To increase electrode cooling to allow RF power to be maintained in a desirable range in areas of low blood flow, fluid irrigation of the electrode either by circulating fluid within the electrode (closed loop system) , or flushing saline through openings in the electrode (open irrigation system) , is used ( Fig. 8.5 ). The “active electrode cooling” by irrigation allows sustained RF power, even at sites with low blood flow, to produce deeper lesions.
Electrode–tissue contact force
Lesion generation in cardiac RF catheter ablation is dependent on the interaction between electrode and contacting tissue. The contact force (CF) between the catheter tip and the target tissue is a key factor to safe and effective lesion formation. , Insufficient CF may result in an ineffective lesion, whereas excessive CF may result in complications such as heart wall perforation, steam pop, thrombus formation, or esophageal injury. Despite its importance, CF cannot be measured directly with available ablation catheters; thus other measures have been used as surrogates for CF, including the pattern of motion of the catheter tip under fluoroscopy, the amplitude of the unipolar and bipolar potentials, and impedance. To measure catheter–tissue CF in real time during catheter mapping and RF ablation, two designs of ablation catheters using different technologies have been developed. One type of catheter uses a small spring connecting the ablation tip electrode to the catheter shaft with a magnetic transmitter and sensors to measure microdeflection of the spring (SmartTouch). The other type of catheter uses three optical fibers to measure microdeformation of a deformable body in the catheter tip (TactiCath). Both systems have CF resolution less than 1 g in bench testing. ,
The 7.5F SmartTouch CF sensing catheter has a 3.5-mm tip electrode with six small holes (0.4 mm diameter) around the circumference for saline irrigation. A tiny spring is located just proximal to the ablation tip electrode. A magnetic signal emitter is attached to the tip electrode (distal to the spring), and three magnetic sensors are located proximal to the spring to measure microdeflection of the spring. The microdeflection is computed to the magnitude and angle of CF every 25.6 ms ( Fig. 8.6 ). CF is displayed both continuously and as the average value (over 1 second) on an electroanatomic mapping system. This catheter also has a magnetic location sensor for conventional electroanatomic mapping. CF catheters have resulted in improved long-term outcomes compared with traditional non-CF catheters. ,