8. QUIZ
Increasing the size of the tip of an RF catheter ablation while
maintaining a constant power level, duration of power
delivery and perpendicular catheter tip orientation to the
tissue surface will:
1. Decrease lesion size
2. Increase lesion size
↑ tip size →
↓power density
3. Result in no significant change in lesion size
13. RF LESIONS - WHAT WE’VE
DISCUSSED SO FAR:
Temperature at the catheter tissue interface is a determinate
of lesion size.
Ablation catheter diameter is a determinate of lesion size.
Tissue temperature diminishes rapidly as you move away
from tissue catheter contact.
Power is directly proportional to the degree of resistive tissue
heating
17. HOWEVER...
When the power not limited, catheters large distal electrodes create
LARGER lesions
↑ electrode surface area in contact with bloodstream, resulting in augmented
convective cooling effect, and
↑ the volume of tissue directly heated
18.
19.
20. QUIZ
Which catheter will create the larger lesion if all other
parameters are constant (power = 30 W, time = 60 sec,
contact force 20 gm-force)
1. Non perfused 7 Fr 4 mm tip catheter
2. Perfused 7 Fr 4 mm tip catheter
If power held constant →
↓ temp at tip of perfused
catheter
21. cooled, RF c power increase
traditional, non-cooled RF
cooled, RF s power increase
24. CONVECTIVE COOLING
Surface cooling
reduces risk of boiling
and coagulum formation
Higher power can be
used
Higher power results
in greater depths of
volume heating
Surface cooling prevents
monitoring of lesion formation
High power can cause
superheating in subendocardial
layers and pop lesions
If ablation is power limited,
power dissipation into the
circulating blood pool can yield
decreased lesion depth
25. QUIZ
At regions of high blood flow (as compared to low flow),
using temperature-controlled RF is likely to:
1. Overheat the tissue
↑ flow → ↑power to maintain temp →
↑ lesion size / depth
2. Less likely to overheat the tissue
3. No difference
32. ASSSESSMENT OF CATHETERTISSUE INTERFACE
Without good contact, only intracavitary blood will be heated, with insufficient myocardial
temperature to cause necrosis of targeted tissue.
Parameters used to assess contact:
beat-to-beat variability in electrograms
baseline electrode impedence
changes in impedance and temperature during ablation
fluoroscopy
visual assessment by echocardiography
pacing capture threshold
tactile feedback
35. TITRATION
Titration to efficacy: a specified power or loss of electrograms
Titration to safety: avoid complications of excessive energy
delivery.
coagulum formation
steam pops
cardiac perforation
collateral damage (coronary arteries)
42. CONCLUSIONS
Tissue injury occurs reproducibly at a temperature of about
50°C
Heat transfer in tissue is a predictable biophysical
phenomenon
Factors including temperature, power, time, impedance,
cooling, and contact force all contribute to lesion size and
depth
Notes de l'éditeur
Overall, the goal of catheter ablation is to destroy the arrhythmogenic tissue.
To accomplish this, to date, radiofrequency and cyroablation ablation have been used in the largest number of patients. We will focus on those modalities today.
RF energy is a form of alternating electrical current that generates a lesion in the heart by electrical heating of the myocardium. Specifically, the goal of RF is to transform electromagnetic energy into thermal energy in the tissue and destroy the arrhythmogenic tissues by heating them to a lethal temperature.
The mode of tissue heating by RF energy is resistive (electrical) heating. As electrical current passes through a resistive medium, the voltage drops, and heat is produced (similar to the heat that is created in an incandescent light bulb).
Typically, an oscillation frequency of 500 kHz is selected. Lower frequencies are more likely to stimulate cardiac muscle and nerves, resulting in arrhythmia generation and pain sensation. Higher frequencies will result in tissue heating, but in the megahertz range the mode of energy transfer changes from electrical (resistive) heating to dielectric heating (as observed with microwave energy).
Generally, RF energy delivery can be divided into 2 main mechanisms, unipolar and bipolar. As you probably know already, the majority of RF we use clinically is unipolar.
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 gets 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.
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. 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
When RF energy is delivered in a unipolar fashion, the current distributes radially from the source.
Resistive tissue heating is proportional to the RF power density, which in turn is proportional to the square of the current density (J). The current density decreases in proportion to the square of the distance from the electrode. The direct resistive heating of the tissue, thus, decreases proportionally with the distance from the electrode raised to the fourth power. As a result, only a narrow rim of tissue in close contact with the catheter electrode (i.e. 2-3mm) is heated directly. Most of the tissue heating, particularly deeper layers are heated by thermal conduction from direct resistive heating that occurs at the electrode-tissue interface.
Actual measurements of temperatures in issue and then in an in vivo setting were performed. At different source temperatures, a fairly predictable decrease in the tissue temperature was seen with increasing distance from the source.
In vitro and in vivo studies have shown that the radial tissue temperature profile follows an inverse proportion with increasing distance from the ablation electrode resulting in a steep tissue temperature gradient. As I mentioned, most of the tissue heating occurs as a result of thermal conduction from direct resistive heating source. The tissue temperature change with increasing distance from the heat source is called the radial temperature gradient. At onset of RF energy delivery, the temperature is very high at the source of heating and falls off rapidly over a short distance . As time progresses, more thermal energy is transferred to deeper tissue layers by means of thermal conduction. Eventually, the entire electrode-tissue system reaches steady state
As I mentioned before, there is a direct inter-relationship between power and temperature. The power density delivered is directly proportional to the degree of resistive tissue heating. Higher power delivery increases source current density – hence temperature and increases radius of the heat source thereby increasing lesion size in two ways.
In vivo studies have shown a dose-response relationship between delivered power and steady-state temperature measured at the catheter tip at individual ablaion sites, but a poor correlation of power and temperature between different ablation sites.
If RF power delivery is interrupted before steady state is achieved, tissue temperature will continue to rise in deeper tissue planes as a result of thermal conduction from more superficial layers heated to higher temperatures. In one study, the duration of continued tem- perature rise at the lesion border zone after a 10-second RF energy delivery was 6 seconds. The temperature rose an additional 3.4°C and remained above the temperature recorded at the termination of energy delivery for more than 18 seconds. This phenomenon, termed thermal latency, has important clinical implications because active ablation, with beneficial or adverse effects, will continue for a period of time despite cessation of RF current flow.
Temperature in the controlled in vitro setting was an excellent predictor of lesion depth. Figure 5 demonstrates a direct relationship between the depth of the lesion and the temperature. When temperature was compared to other electrical parameters including current, power, and energy, it seemed to be the best predicator (Fig. 6) – keep in mind this is in the absence of heat loss due to convective cooling by bloodflow. Clinically, the opposing effect of convective cooling by bloodflow diminshes the value of temperature monitoring to assess lesion size.
A study using an isolated perfused and superfused porcine right ventricular (RV) free wall preparation has shown that the steady- state temperature recorded at the ablation electrode-tissue interface was a more accurate predictor of lesion size than measured power, current, or energy.
This being said, there is a dose-response relationship between delivered power and steady-state temperature measured at the catheter tip at individual ablation site.
It is intuitive that if one starts out with a large heat source, a larger lesion should be formed. If a pinpoint source of heat is used, less energy transfers to the tissue and the lesion size might be anticipated to be smaller. Haines et al.5 made this observation in the controlled in vitro setting in 1990. Lesion size was found to be directly proportional to the diameter of the electrode used. Figure 7 demonstrates the relationship between electrode radius and depth and diameter of resultant lesion. 7-Fr OD catheter ~ 2.3mm
In vitro and in vivo studies have shown that significant tissue heating during RF delivery is associated with a slight decrease in measured impedance, typically in the range of 5 to 10 ohms.
Unfortunately, this process is limited in the biologic setting by the formation of coagulum and char at the electrode-tissue interface if temperatures exceed 100°C. At 100°C, blood literally begins to boil. This can be observed in the clinical setting with generation of showers of microbubbles if tissue heating is excessive.
As the blood and tissue in contact with the electrode catheter desiccate, the residue of denatured proteins adheres to the electrode surface. These substances are electrically insulating and result in a smaller electrode surface area available for electrical conduction. In turn, the same magnitude of power is concentrated over a smaller surface area, and the power density increases. With higher power density, the heat production increases, and more coagulum forms. Thus, in a positive-feedback fashion, the electrode becomes completely encased in coagulum within 1 to 2 seconds.
In a study testing ablation with a 2-mm-tip electrode in vitro and in vivo, a measured temperature of at least 100°C correlated closely with a sudden rise in electrical impedance. Almost all ablations without a sudden rise in electrical impedance had a peak temperature of 100°C or less, whereas all but one ablation manifesting a sudden rise in electrical impedance had peak temperatures of 100°C or more.
The major thermodynamic factor opposing the transfer of thermal energy to deeper tissue layers is convective cooling. With the case of RF catheter ablation, the heat is produced by resistive heating and transferred to deeper layers by thermal conduction. Simultaneously, the heat is conducted back into the circulating blood pool and metal electrode tip. Because the blood is moving rapidly past the electrode and over the endocardial surface, and because water (the main constituent of blood) has a high heat capacity, a large amount of the heat produced at the site of ablation can be carried away by blood.
During temperature-controlled RF ablation, the tip temperature, tissue temperature, and lesion size are affected by the electrode-tissue contact and by cooling effects resulting from blood flow. With good contact between catheter tip and tissue and low cooling of the catheter tip, the target temperature can be reached with little power, resulting in small lesions even though a high tip temperature is being measured. In contrast, a low tip temperature can be caused by a high level of convective cooling, which results in higher power delivery to reach the target temperature, yielding a larger lesion.
Greater power delivery can be achieved by cooling the tip of the catheter using either passive tip cooling (i.e. convective cooling by blood flow) or active tip cooling (i.e. an open or closed saline irrigated system). As the magnitude of convective cooling increases, there is decreased efficiency of heating due to more energy being carried away in the blood and less energy delivered to the tissue. Excess heating at the surface is decreased by cooling of the catheter tip. This allows the power to be turned up without the risk of coagulum and char formation. Increased power delivery means that heating occurs deeper in the tissue, making for greater depth of resistive heating into the myocardial target tissue, thus yielding larger lesions.
Alternatively ablation electrode can be actively irrigated with saline to reduce the electrode-tissue interface temperature and prevent an impedance rise. Compared with conventional RF application, cooled ablation allows passage of both higher powers and longer durations of RF current with less likelihood of impedance rises. In addition, because convective cooling from the bloodstream is not required, an irrigated electrode may be capable of delivering higher RF power at sites of low blood flow, such as within ventricular trabecular crevasse
An interesting and important side note... with traditional RF, the site of maximum temperature will be the electrode-tissue interface, whereas with an irrigated tip, the hottest point may actually be within the myocardium, not subject to convective cooling. That tissue may exceed 100 deg C, and is subject to steam popping and dissection, perforation, and thrombus formation.
Answer is 1. In the temperature-control mode, the generator alters power output to maintain a specific temperature at the tip of the electrode. This tip is simultanously being cooled by the high flow of blood. Therefore, the generator increases power output to increase the electrode temperature to the set value. This higher power output leads to an increase in current density, greater resistive tissue heating and an increase in lesion size/depth.
The predominant mechanism of tissue injury by RF energy delivery is presumed to be thermally mediated. The myocardial cellular electrophysiologic effects of hyperthermia have recently been described in an in vitro model of isolated guinea pig RV papillary muscle. In this study, excised RV papillary muscles were pinned in a high-flow tissue chamber and superfused with Tyrode's solution. The temperature of the superfusate was rapidly changed from 37°C to between 38°C and 56°C for 60 seconds and returned to 37°C. The muscles were impaled with a single microelectrode to measure the resting membrane potential, the maximum rate of increase of the action potential, action potential amplitude, and action potential duration. Hyperthermia was found to cause significant changes in myocardial cellular electrophysiologic properties that included (1) a marked depolarization of the resting membrane potential at temperatures > 45°C (Fig 9)
RV papillary muscle preparation during exposure to a temperature of 48.1°C. The muscle was paced at a frequency of 1 Hz. There was an initial depolarization of the membrane potential (Vm)after hyperthermic exposure that resulted in loss of action potentials {vertical spikes above baseline) in response to pacing stimuli (vertical bars below baseline). After returning to 37°C, the membrane potential repolarized to baseline values with immediate restoration of normal excitability.
mia on cellular excitability of isolated RV guinea pig papillary muscles. Vertical bars represent median values with the interquartile range. The median temperature associated with normal excitability (44.0°C) is significantly lower than the median temperatures associated with reversible loss of excit- ability (48.0°C) and irreversible loss of excitability and tissue injury (50.5°C). Subsequent histo-path studies have demonstrated that calcium overload occurs at temperatures 50°C resulting in irreversible tissue injury.
Key factors influencing the size and depth of an RF ablation lesion include current density at the electrode tip (in turn determined by delivered power and electrode surface area),6 electrode-myocardium contact, orientation of catheter tip,7 duration of energy delivery, achieved elec- trode tip temperature, and heat dissipation from intracavitary blood flow or nearby cardiac vessels. Because some of these factors are unknown during ablation, power is often increased to reach a prespecified goal (e.g., 40 to 50 W for ablation of the right atrial isthmus) or to a desired effect (e.g., loss of preexcitation or tachycardia termination). Power titration is also modulated by electrode impedance and temperature monitoring in the clinical setting.
Although efficacy is important, it is also critical to avoid complications of excessive energy delivery. Careful titration of RF power can minimize the probability of coagulum formation, steam pops, cardiac perforation, and collateral damage to intracardiac and extracardiac structures.
Worrisome? No.
Plot of impedance, power, and temperature during catheter ablation of a left posteroseptal accessory pathway using a conventional 0 4-mm-tip catheter. Blood flow was brisk, and convective cooling kept the catheter-tip temperature below 50°C despite high power (50 W). However, even without a high temperature at the catheter tip, evidence of tissue damage was seen. Accessory pathway conduction was blocked in less than 3 seconds, 36 and impedance fell by more than 15 ohms during energy application.
Data recorded during lesion application that resulted in steam pop and transmural left atrial tear from barotrauma. At the moment of microbubble release on intracardiac echocardiography, a small, non- sustained rise in impedance was observed (arrow). A few seconds later, electrode temperature rose abruptly, as bubbles engulfed the ablation electrode.
Excessive temperature rise during irrigated radiofrequency (RF) ablation. During ablation in the left atrium for atrial fibrillation, using an externally irrigated 4-mm-tip catheter with flow rate of 17 mL/min, the operator noticed a steadily rising temperature and discontinued RF when it reached 45°C. This probably resulted from intimate tissue contact that prevented adequate cooling of the electrode tip. Other possible responses would have been increasing the irrigation flow rate, reducing power, or repositioning the catheter.
