The Biophysics of Radiofrequency Ablation and Factors Affecting Lesion Size

Unipolar radiofrequency ablation (RFA) has been the mainstay of interventional electrophysiology since its conception.¹ The objective of ablation is to eliminate arrhythmia by either disrupting crucial circuitry, for example with re-entrant arrhythmia, removing triggers, for example with focal atrial/ventricular tachycardias, or by isolating arrhythmogenic tissue, such as the pulmonary vein.

To achieve this using RFA, viable myocardium is rendered electrically inert via irreversible damage from tissue heating. However, creating large ablation lesions is limited by the risk of complications such as tissue charring or steam pops.² Therefore, to obtain optimal outcomes, knowledge of the biophysics and procedural factors in creating a RFA lesion are fundamental.

Biophysics of Radiofrequency Ablation

The Ablation Circuit

To understand the development of an RFA lesion, it is important to be aware of the underlying electrical circuit. This can be simplified to a lumped element model with three resistors in series connected by perfectly conducting wires (Figure 1 ).³

These three resistors are:

• Z_A = the resistance at the interface of the electrode tip;
• Z_B = the resistance of body tissues, such as lungs, blood vessels and adipose tissue, en route to the dispersive electrode; and
• Z_DE = the resistance of the tissue around the dispersive electrode (mainly keratinised epidermis at the skin patch).

As the electrode tip is in simultaneous contact with both blood and myocardium which have different resistances, Z_A is divided into two resistors in parallel: Z_EBI (electrode blood interface) and Z_ETI (electrode tissue interface). Resistors convert electrical energy into heat which spreads over the surface area of the resistor.

During RFA, energy is emitted from an ablating catheter at Z_A as an alternating current of 350–700 kHz. This frequency is capable of heating tissue, but it is too rapid to be able to electrically capture the myocardium, induce arrhythmia or cause dielectric heating similar to microwave ablation. The electrical current then flows via the patient’s tissues (Z_B) to the dispersive electrode (Z_DE). The surface areas of Z_B and Z_DE are large and therefore their heating effect is negligible as the current spreads across them and is estimated to be 0.01°C for an individual weighing 75 kg.⁴ At Z_A, the tip of a 4 mm ablation catheter has a surface area of only 27 mm². This small surface area enables the tip to transfer focused current to the myocardium when the circuit is activated. At the ETI of the intended ablation site, the tissue experiences an area of high current density and, due to acceleration of cellular ions caused by the alternating electrical current, undergoes resistive heating. In contrast, the EBI experiences minimal heating due to the constant convective cooling effect of circulating blood.

It should be emphasised that the lumped circuit model above is a simplification of the true RFA circuitry. An interventional electrophysiologist should appreciate that due to heterogeneous cellular architecture, Z_ETI does not hold constant resistance throughout, even in healthy myocardial tissue.

Thermodynamics

Due to the complexity and varied resistance within biological tissue, resistive heating is distributed according to current density vectors during RFA. That is, the current per unit area based upon the amount of flow of electrical charge to that part of tissue. This is best defined thermodynamically by the continuum of Joule heating:

Heat generated per unit volume = Current density²/electrical conductivity

However, in practice, to simplify the understanding of the biophysical response of tissue to RFA, it can be assumed that the ETI has a constant resistance throughout, as per the lumped element model described above. One must also appreciate this model weakens as tissue heterogeneity increases, which happens when myocardial tissue becomes more diseased.

With this, we use an adapted Joule’s law to approximate the area of resistive heating which is proportional to the square of the current density in the tissue:

Heat generated = Current² × Impedance × Time

Furthermore, the current density itself is inversely proportional to the square of the distance from the ablating electrode to the tissue.

Current density ∝ 1/Distance from electrode tissue interface²

This is true in a homogeneous medium and therefore the approximation must be acknowledged, particularly for very diseased heterogeneous myocardium.⁵ This emphasises the importance of forming an effective ETI as the degree of resistive heating will be proportional to the distance of the electrode to the fourth power. Consequently, the current density rapidly falls within the myocardium in a radial manner when ablation is delivered. Therefore, the tissue volume experiencing exclusively resistive heating is relatively small at approximately 2 mm in all dimensions. From this, a steady state develops within a few seconds.

The site of resistive heating then radiates to adjacent tissues in a process called conductive heating, causing further damage and completing the ablation lesion. This process is slower, taking up to 2 minutes before reaching thermal equilibrium.^6,7 At this point, the energy delivered to generate heat is matched by its conductive loss to the tissues and convective loss to the blood pool and the lesion reaches its maximum size.

In this process, the conduction of heat through the tissue receiving RFA is affected by numerous dynamic factors, including the rate of temperature change, the tissue composition affecting conductivity, blood perfusion and tissue metabolism. This is defined by Penne’s bioheat equation:⁸

ρc. δT/δt=Δ. (kΔT)+ρ_bc_bω_b (T_a−T)+Q_m+Q_ext

ρ = density of tissue; c = specific tissue heat capacity; δT/δt = the change in tissue temperature over time; k = tissue thermal conductivity; ρ_bc_bω_b = the density, heat capacity and perfusion rate of blood; T_a = arterial blood temperature; Q_m = metabolic heat generation; Q_ext = external heat source (e.g. ablation catheter).

In 1989, a seminal study by Haines et al. was able to demonstrate this proof of concept. Using an in vitro canine model, their study showed that increasing the temperature at the ETI from 50°C to 85°C strongly correlated with larger ablation lesion depth and width. They also showed the tissue temperature fell with its distance from the ablating electrode and was predictable in a hyperbolic thermodynamic model.⁹

Cellular Mechanics of Thermal Damage

As RFA occurs and tissue temperature rises, cardiac myocytes go through a series of electrophysiological changes. At 45°C, marked depolarisation of the cell membrane occurs, followed by an increased rate of action potential rise (phase 0) before a fall in overall amplitude (phases 1 and 2). Finally, irreversible loss of excitability of the cell occurs at temperatures greater than 50°C after a duration of 60 seconds.¹⁰ Macroscopically, this results in a lesion with a pale necrotic core of coagulation necrosis surrounded by a haemorrhagic periphery. Microscopically, infiltration of inflammatory mononuclear cells and neutrophils are seen. Electron microscopy shows ultrastructural damage occurs up to 6 mm away from what is visible.¹¹ RFA lesions turn into a fibrous scar with granulation tissue, fat cells, cartilage and infiltrates that renders the tissue electrically inert.^12,13

The achievement of a thermal dose is necessary to achieve irreversible tissue damage. This is based upon the cumulative tissue temperature achieved across ablation duration and derived from the Arrhenius equation.^14,15 With this, small increases in temperature can increase tissue necrosis dramatically, and this forms the basis of high-power, short-duration ablation techniques.

Biophysics Summary

Fundamentally, ablation lesion size is dependent on the temperature that can be generated in the myocardium and its ability to conduct heat through it. Procedural factors determining this must affect one of the components of Joule’s law: the current supplied; the impedance at the ETI; the duration of delivery; or the thermoconductive tissue properties (Table 1 ).

Factors Affecting Lesion Size

Current

As discussed above, the amount of current delivered to the myocardium is the key electrical parameter determining heat generation and thus RFA lesion size.¹⁶ Interventional electrophysiologists will be more familiar with radiofrequency (RF) energy delivery as measured in Watts rather than current with ampères. However, the relationship between power and current is not linear, rather being proportional to the square of the current as (similar to the thermodynamic concepts above), Joule’s law states that:

Power = Current² × Tissue impedance

Therefore, an operator should be aware that for RFA, assuming all other biophysical and catheter factors are identical, selecting a higher power setting will not result in a directly proportional increase in current, but rather proportional to the square root of such an increase. For example, doubling power would result in an increased current proportion of √2, approximately 40% greater, rather than double.

By the same logic, increasing from seemingly low levels of power will have larger effects on current delivery than that at higher levels. For example, increasing from 15 W to 20 W is an equivalent rise in current as 30 W to 40 W as both have proportionally increased by one-third. In both scenarios 15% greater current will be delivered.

In multiple pre-clinical studies, with other factors controlled, increasing power results in larger RFA lesions.^16–19 However, with increasing power settings, larger numbers of complications were seen. This was initially seen with temperature-guided catheters in the form of plasma boiling at 100 ^oC. The resultant pops caused tissue shredding and formation of coagulum coating the electrode risking thromboembolic events.^19,20

Catheter Irrigation

Consequently, methods to cool the catheter tip while maintaining tissue temperature to maximise lesion size were developed.^2,21 Irrigated ablation catheters increase the size of ablation lesions by enhancing the convective cooling effect of the circulating blood pool. Cooling the electrode tissue interface decreases the risk of tissue charring and coagulum formation and thus higher power delivery and current density to the ETI can occur. Studies have demonstrated that ablation lesions with significantly greater depth and volume than non-irrigated catheters were created as a greater delivery of current could be provided at the expense of the knowledge of the catheter tip temperature. Perhaps unsurprisingly, the rate of cooling irrigation was then shown to also affect lesion dimensions, with a smaller surface diameter seen with increased flow.²² Unfortunately, irrigated catheters are also limited by complications at high temperatures, forming steam bubbles below the ablating electrode and resulting in crater formation and the risk of myocardial rupture.²

Irrigated ablation catheters come in closed- and open-loop varieties. Closed-loop catheters circulate fluid through the catheter to cool the electrode internally. By comparison, open-loop catheters provide cooling by allowing the fluid to exit at the site of the electrode, directly cooling the tip and surrounding tissue. The catheters produce comparable lesion sizes. However, it should be noted that in sites of low blood flow, open-loop catheters maintain a higher level of convective cooling and therefore are at lower risk of complications caused by excessive ETI temperatures. Due to ETI cooling, ablation lesions with irrigated catheters cause a peak temperature subendocardially, meaning its greatest diameter is below the surface and holds the shape of a spherical frustum rather than a hemisphere.

Recently, catheters have come to market that deliver very high-power, short-duration RFA lesions.^23,24 Fundamentally, in this method the biophysics of RFA lesion formation are the same, but there is a greater emphasis on current delivered over ablation duration. The premise behind very high-power, short-duration ablation is that with high-power delivery, resistive heating occurs over a larger tissue volume, creating the necessary lesion size and abolishing the requirement for conductive heating. The short duration of ablation (4 seconds) negates the chances of steam pop occurring and despite the short duration, an adequate thermal dose is achieved due to the high power. A recent meta-analysis has shown the advantages of using these catheters, with reduced procedural and ablation time resulting in comparable arrhythmia-free outcomes and without increases in the complication rate.²⁵

In summary:

• Increasing power delivers more current through the ablation circuit, increasing the current density and temperature at the ETI, resulting in a larger lesion.
• Increasing power results in raised current proportional to the square root of the relative increase.
• Augmentation of current can be facilitated by catheter irrigation but is limited by steam pops.
• Novel, very high-power, short-duration ablation catheters rely on resistive tissue heating over conductive.

Impedance at the Electrode Tissue Interface

Impedance at the ETI is a key component in RFA lesion size. Procedurally, impedance is produced by numerous interconnecting, often dynamic factors which need to be integrated by the operator to optimise each RFA lesion.

First, an operator must understand the origin of the impedance measurement displayed on their electrophysiology systems. Impedance (sometimes also known as generator impedance) is the measurement of the total resistance throughout the ablation circuit and Z_A+Z_B+Z_DE. Z_A – the specific impedance of the tissue intended to be ablated – is the value of most interest. However, impedance can be affected by Z_B, which varies with the patient’s body composition, BMI, breathing, movement and even different locations in the heart.^3,26,27 Therefore, one must appreciate the intra- and inter-patient variability in Z_B will affect impedance readings and consequently values from one patient cannot be translated directly to another.

Some ablation catheters offer a different, novel biophysical parameter known as local impedance (LI). LI overcomes the variation caused by Z_B by measuring Z_A only, giving the operator a focused assessment of the impedance specifically at the catheter tip. This is achieved by creating an electrical field between the catheter tip and a ring electrode and measuring any distortions caused within, for example, by tissue contact. The LI can then be measured by dividing the change in voltage by the supplied current.^28,29 Computer models, pre-clinical and clinical studies have all shown increases in LI are indicative of tissue contact and reflect underlying tissue architecture.^29–32

Second, the impedance of the catheter tip is the biophysical factor that determines the current delivered. From Joule’s law, as the power delivered is fixed by the operator, tissue with low baseline impedance must, by definition, receive greater current than tissue with a higher baseline impedance. However, the propensity of said tissue to undergo resistive heating and conduct this is a different matter.

Third, tissue impedance is dynamic during RFA, falling as myocardium heats. This fall in impedance is caused by an increase in conductance across the tissue as its intra- and extracellular fluid viscosities fall, allowing ions to travel more freely.³³ As the impedance falls, the current delivery to the tissue further increases as per Joule’s law as described above, allowing for greater resistive heating. Eventually, an impedance plateau will be reached, signalling a thermal equilibrium within the tissue. These changes in impedance follow a logarithmic model and help guide the operator to the evolution of the lesion (falling impedance) and its completion (plateau).^34,35 Further, the degree of impedance drop is a useful value, demonstrating strong correlation with ablation lesion size and alongside the rate of impedance drop, the chances of a steam pop.^19,20,36,37 Drops in LI have also shown similar outcomes, with some studies showing a stronger correlation with RFA lesion dimensions than generator impedance.^28,29

Fourth, impedance is indicative of tissue contact and the degree of electrode tissue coverage (ETC).³⁸ As the blood pool has a lower impedance than myocardium, any contact or proximity of contact will cause a rise in impedance.^28,31,32 It should be noted that the EBI is the preferred route for electrical current flow due to its lower impedance. Consequently, it is the role of the operator to optimise the ETC to minimise this electrical shunting. During the procedure, the ablating electrode will be in contact with the tissue (the ETI) and the blood pool (EBI). With greater ETC, the impedance will rise accordingly.³⁹ In this scenario, when providing RFA, greater current will be transferred to the ETI rather than dissipating via the EBI, creating a larger ablation lesion. ETC is not a defined parameter available on current electrophysiology systems and baseline impedance can be used to indicate this and the projected size of the ablation lesion.³⁸

Indeed, other factors known to affect ablation lesion size act as surrogates for ETC including contact force, cardiac rhythm and catheter orientation. Contact force is a well-researched parameter, with studies consistently showing that greater contact and greater contact with time (force-time integral or FTI) results in larger ablation lesions.^{6,19,29,39–41} The greatest effect of this is with initial contact in the 0–10 g range, with smaller increases seen from 10–40 g. From a biophysical standpoint, it is not the degree of force itself allowing for larger lesions, but the greater electrode tissue coverage and therefore current transfer it provides. Consistent with this, impedance also rises across these contact force levels, with plateaus seen from 20 g onwards.^30,32,42

Cardiac rhythm affects impedance by varying the tissue coverage during cardiac contraction. In sinus rhythm, the coordinated ‘kick’ of the atria causes greater variation in impedance and results in a lower baseline impedance level compared to AF.³² Similarly, studies have shown greater contact force variation in sinus rhythm compared with AF, and linked this to poorer impedance drops with RFA.⁴³ These findings are important, as they suggest that to optimise ETC, as demonstrated by a higher baseline impedance, ablating during an arrhythmia such as AF may be preferable to ablating during sinus rhythm.

Orientation has been shown in pre-clinical studies to affect impedance with parallel orientation showing higher values when controlled for contact force.^44,45 This is due to the greater ETC formed along the side of the catheter and consequently larger volume lesions are seen with non-irrigated catheters due to increased breadth in these simulated situations. In contrast, with irrigated catheters, smaller lesions have been seen due to the irrigation ports being in direct contact with tissue, causing a ‘cooling rail’ effect along the ablation electrode.⁴⁵ In clinical practice this is more complicated, as catheter orientation and the electrode tissue coverage it provides will be affected by anatomy, for example, being wedged in trabeculations or sliding along the surface of a smooth left atrium. Consequently, although parallel orientation has the potential to create a better biophysical condition for RFA, it is less reliable.

Fifth, the preference for electrical current to travel to the ETI can be affected by increasing the impedance of the EBI. This can be altered by varying the ionic composition of the fluid selected to flow through an irrigated catheter. Isotonic fluid, such as 0.9% sodium chloride, is commonly used for RFA, and as it has lower impedance than the myocardium, it shunts current away from the ETI. In contrast, using hypotonic solutions, such as 0.45% or ‘half’ sodium chloride, has a higher impedance than the myocardium, simply due to its lower concentration of ions. Consequently, current preferentially flows to the ETI, increasing heat, and using 5% dextrose increases this effect even further. ^46–48 Both produce significantly larger RFA lesions than isotonic fluids, but due to the increased heating, there is a larger risk of steam pops.

Finally, baseline tissue impedance is affected by the underlying tissue characteristics. Healthy, densely packed, uniform myocardium has been shown to have a higher impedance than diseased tissue composed of scattered, damaged cardiomyocytes, collagen and adipose.^49–51 This difference in impedance is progressive depending on the extent of disease. From a biophysical point of view, the lower impedance of diseased myocardium actually means it receives greater current and thus has more potential for resistive heating. However, as diseased tissue is fibrous and is not composed of free ions capable of generating heat, lower temperatures are achieved. Further, current is shunted from the area by adipose tissue, which is a poor thermal conductor. Consequently, in diseased myocardium, despite having a more electrically optimal circumstance for RFA, smaller ablation lesions are created.^37,52

In summary:

• Impedance shows inter- and intra-patient variability.
• Impedance at the catheter tip is a key factor in determining RFA lesion size by modulating the current delivered.
• The amount of ETC raises impedance by directing current to the ETI rather than shunting to the EBI. Therefore, increasing impedance indicates greater ETC and potentially a larger lesion size.
• Greater current can be diverted to the ETI by raising the impedance of the EBI with hypotonic irrigants.
• Contact force is a surrogate marker for ETC.
• Impedance changes during RFA and can be used to judge the evolution of and completion of a lesion.
• Baseline tissue impedance is progressively lowered by the extent of histological myocardial disease, and although greater current is delivered, this is offset by poor thermal conductance.

Electrode Size

Another factor that affects the size of the ablation lesion is electrode size. The most commonly used RFA catheters have a 4 mm diameter. With increased electrode surface area, both impedance (due to the relationship of ETI to EBI) and current density supplied to the ETI will be affected. Theoretically, with a larger electrode, greater current density could be supplied to the endocardium using a larger power input to create a larger lesion and improve procedural efficiency.⁵³ However, a larger tip will also increase the surface area in contact with the EBI, and subsequently, the current could shunt away from the target ETI. Increasing catheter size also confers poor endocardial contact due to stiff, difficult-to-manoeuvre catheters, resulting in smaller lesions being generated.⁵⁴

Duration

During RFA, the area of resistive heating rapidly achieves a steady-state temperature. However, the area of conductive heating takes more time to produce the full lesion size. Consequently, allowing for an adequate duration of RFA is necessary to achieve this. Haines et al. showed lesion width and depth increased with ablation duration monoexponentially, with a half-time of 7–10 seconds and eventually plateauing at 45–60 seconds.^6,9 Further pre-clinical studies have shown lesion size continues to grow even up to 90 seconds.^7,55 It is important to note that the mono-exponential model also demonstrates that ablation lesion size increases most rapidly in the first few seconds of RFA delivery. A recent study by Balhke et al. demonstrated lesion diameter and depth to increase by 2.71 and 1.08 mm/s respectively in this time before both slowing to 0.1 mm/s after 20 seconds.¹⁸ An operator should bear this in mind when delivering RFA as prolonged applications will have diminishing returns.

Pulsed Field Ablation

The advent of pulsed field ablation (PFA) has opened up a new method of inducing irreversible cellular damage via electroporation. The biophysics of PFA are beyond the scope of this review; however, the principles are comparable to RFA, with factors affecting lesion formation including field strength, pulse duration, pulse number, electrode size, configuration, contact force, tissue conductivity, anisotropy and impedance.⁵⁶

Conclusion

Knowledge of biophysical principles in creating an RFA lesion is fundamental to the interventional electrophysiologist. Lesion size is determined by the temperature achieved within the myocardium, in turn affected by thermodynamics affecting either the current supplied, the impedance at the electrode tissue interface, the duration of delivery and the thermoconductive properties of the tissue. By integrating these, lesion formation can be optimised according to the desired arrhythmic treatment.