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Zaltieri, M. Myocardial Radiofrequency Ablation Temperature Monitoring. Encyclopedia. Available online: https://encyclopedia.pub/entry/10686 (accessed on 16 November 2024).
Zaltieri M. Myocardial Radiofrequency Ablation Temperature Monitoring. Encyclopedia. Available at: https://encyclopedia.pub/entry/10686. Accessed November 16, 2024.
Zaltieri, Martina. "Myocardial Radiofrequency Ablation Temperature Monitoring" Encyclopedia, https://encyclopedia.pub/entry/10686 (accessed November 16, 2024).
Zaltieri, M. (2021, June 09). Myocardial Radiofrequency Ablation Temperature Monitoring. In Encyclopedia. https://encyclopedia.pub/entry/10686
Zaltieri, Martina. "Myocardial Radiofrequency Ablation Temperature Monitoring." Encyclopedia. Web. 09 June, 2021.
Myocardial Radiofrequency Ablation Temperature Monitoring
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This entry is adapted from the peer-reviewed paper 10.3390/s21041453

Myocardial catheter-mediated radiofrequency ablation (RFA) is a minimally invasive technique which exploits high-frequency alternating electrical current to induce irreversible damage in selected myocardial districts through hyperthermia. (RFA) has received substantial attention for the treatment of multiple arrhythmias. In this scenario, there is an ever-growing demand for monitoring the temperature trend inside the tissue as it may allow an accurate control of the treatment effects, with a consequent improvement of the clinical outcomes. There are many methods for monitoring temperature in tissues undergoing RFA, which can be divided as invasive and non-invasive. 

temperature measurements thermocouples thermistors fiber bragg grating sensors fluoroptic sensors infrared thermometry magnetic resonance thermometry ultrasound thermometry myocardial radiofrequency ablation cardiac radiofrequency ablation

1. Introduction

Catheter-mediated radiofrequency ablation (RFA) is the most widely used procedure in the field of cardiac electrophysiology. In fact, since its first application in cardiology in 1987 [1][2][3], RFA has emerged as the key procedure for the treatment of multiple arrhythmias, due to the low mortality and morbidity associated with this practice, together with its high success rate [4].

Myocardial RFA induces irreversible damage in selected myocardial districts through hyperthermia. During RFA, high temperatures of at least 50 °C are reached to cause irreversible damage on the target tissue with consequent cell death. Instead, temperatures equal or above 100 °C should not be attained since are often cause of dangerous complications [5]. In fact, steam popping, tissue perforations and hematic clots upon the catheter tip are the main possible operative drawbacks, since at 100 °C the water contained into the cells undergoes immediate boiling, the blood proteins denature and the tissues surrounding the catheter tip incur drying [6][7][8]. From a macroscopic point of view, RFA produces lesions that are constituted by a central portion of necrotic tissue bordered by a zone of inflamed tissue, in which cellular excitability is zeroed [9].

As shape and dimension of the produced lesions, and consequently the outcome of procedure, are strongly related to the temperature and its history, being able to monitor the temperature increase of the treated tissue during cardiac RFA may be of fundamental importance.

2. General Principles of Radiofrequency Ablation

During RFA procedures, an alternating electrical current at high-frequency is provided to the target tissue by means of a catheter. Specifically, a continuous unmodulated sinusoidal waveform current whose frequency ranges from 350 kHz to 750 kHz is produced by the RF generator. These frequency values are not high enough to induce ventricular fibrillation [5]. The current is delivered between the antenna tip electrode and a ground plate which is located in contact with the patient’s skin with the help of electrical conducting gel. On the tip electrode, called ablation electrode, the passage of the electrical current is focused, while on the ground plate (also called dispersive or indifferent electrode) minimum current density is ensured thank to its large area of contact (typically greater than 10 cm2) [4].

The formation of the thermal damage (or lesion) is the result of the heating mechanism. Such process can be considered the outcome of two main contributes: the resistive heating involving the tissue surrounding the tip, and the conductive heat transfer into the underneath layers [10]. Considering the capacitive effects negligible, the power delivered into the target tissue during RFA is reported in the equation below:

P = I2 R

(1)

where I is the current provided, and R the total resistance (i.e., the sum of the resistances relative to the catheter, blood, tissue, and ground plate). Hence, resistive heating process is strictly related to the local power density (p) which, in turn, depends on the current density (j) and R, according to the following equation:

p = j2 R

(2)

Given that j decreases as 1/r2, where r is the distance from the catheter application point, p drops within the tissue as 1/r4 [11]. Therefore, only a thin layer (that is about 1 mm [4]) of tissue surrounding the tip can be considered subjected to the resistive heating action [12]. Instead, the total damaged volume (whose dimensions depend on several factors such as delivered power, treatment time and pression exerted on the tissue by the tip) is governed by both conductive and convective heating exchanges. In fact, the layers below the antenna exchange conductive heat, while the interaction with flowing blood and tissues at lower temperatures provokes convective cooling.

The cells destruction and irreversible injury depend on temperature and exposure time [13]. In myocardial tissue, cells depolarization with consequent excitability loss starts at about 43 °C and remains reversible until reaching 48 °C for any treatment time. Several in vitro experiments have identified in the temperature of about 50 °C the value at which permanent injury occurs [5]. Moreover, as the temperature increases, the time for reaching cytotoxicity shortens, as described by the Arrhenius’ equation [14][15]:

Ω = ∫A exp(-E/RT(t))dt 

(3)

where Ω is the natural log of the ratio of the concentration of the altered tissue state to the original state, A the collision frequency, E the activate energy, T(t) the absolute temperature, and R is the universal gas constant.

3. Techniques for Temperature Monitoring of Myocardial Tissue undergoing Radiofrequency Ablation

Monitoring the temperature increase of the treated tissue during cardiac RFA and, more generally, during all kind of RF thermal treatments which are typically exploited for the solid tumors’ removal (e.g., liver, lung, pancreatic, and kidney cancers), may be of fundamental importance, not only to ensure the success and the safety of the procedures, but also to adjust in progress the parameters set (e.g., power delivery and treatment time). Given this need, in the last decades the effort made by the researchers to develop performant methods for temperature monitoring during RFA procedures has led to the achievement of several solutions exploiting various technologies. These solutions can be divided into invasive methods (i.e., requiring direct contact or insertion into the tissues to be monitored) such as thermocouples, thermistors, fluoroptic sensors and fiber Bragg grating sensors, and non-invasive methods (i.e., contactless solutions) such as Computed Tomography (CT), Magnetic Resonance Imaging (MRI), ultrasound tomography and infrared (IR) thermometry.

Concerning temperature monitoring during RFA cancer treatments, MRI thermometry [16][17][18] has played a key role, while more rarely IR thermometry [19][20] has been used. Moreover, some attempts have been made also through the use of ultrasound-based thermometry [21][22][23] and invasive methods, such as thermocouples [24][25] thermistors [26][27][28][29], and fiber optics [30][31][32][33], but limited to laboratory settings.

Focusing on myocardial RFA, thermocouples and thermistors embedded in RF emitting antennas have been widely adopted both in the clinical practice [34][35][36][37][38][39][40][41] and in the experimental field [42][43][44][45]. Nevertheless, such configuration records a single-point measurement which does not provide direct information regarding the temperature of the tissue undergoing RFA. To avoid this issue, one of the first approach used, but still widely employed for its ease of use and low costs, is the insertion of thermal probes holding either thermocouples [43][46][47][48][49] or thermistors [50][51][52] directly into the treated tissue. Thermocouples are typically encapsulated into needles or shields to promote chemical isolation from the surrounding biological environment. These sensors are well known, present small size, low price, and adequate performances relative to the application context. However, the conductive metallic components which constitute thermocouples may interact with the incident RF field, so inducing measurement artifacts [53][54]. These effects can be partially overcome by preferring small diameters probes inserted into the tissues with the long axis orthogonal to the incident electromagnetic field [55]. In addition, the presence of one or more thermocouples probes within the operating field might obstruct the already crowded area, so limiting their usage in the clinical scenario. The application of thermistors in the field of cardiac RFA presents an evolution comparable to that of thermocouples for similar characteristics and performances [56]. Low cost, small size, and fast time response, as well as better accuracy are the main advantages related to this sensor. On the contrary, thermistors’ measurements can be affected by noise as a result of the coupling between the connecting wires holding metallic components and the RF radiation [55].

In addition, more than a decade later, fluoroptic probes [57][58][59] were employed to obtain more reliable and high-resolved temperature measurements. Among all the FOSs, fluoroptic sensors are the prevailing technology in the literature for temperature monitoring during myocardial RFA. These sensors present two possible configurations: embedded into probes or, more rarely, without any shield. The presence of the probe causes a significant increment of the response time. The quality of the measurements, the immunity to electromagnetic fields, small size, flexibility, and the multiplexing capability have led over the years to an ever-greater use of this type of sensor. On the contrary, the use of FBGs in the field of cardiac RFA is limited to a very recent study [60]. Good thermal sensitivity and accuracy, short response time, multiplexing capability, immunity to electromagnetic fields are the main advantages brought by this technology. On the contrary, the fragility and the cross-sensitivity to strain (caused for example by cardiac and respiratory activity) may limit the use of this methodology in in vivo trials on myocardium. However, compared to thermocouples and thermistors, both fluoroptic fibers and FBGs present significantly higher costs.

In the 2000s, the interest in contactless thermometric methodologies led to the exploitation of techniques based on ultrasound tomography [61], IR [62][63][64][65], and MRI [66][67][68][69][70].

non-invasive techniques allow the thermal pattern reconstruction of the myocardial tissue subjected to RFA avoiding direct contact with the organ. MRI, ultrasound imaging and IR imaging ensure no insertion of additional tools inside the treated area, so reducing overcrowding in already flocked operating fields with experimental or surgical equipment. Also, all the presented imaging methodologies did not use ionizing radiations. Moreover, the lack of contrast fluids intake makes them suitable for repeated applications.

Among the others, MRI is the most exploited. The use of MRI offers high precision, good temporal resolution, linear relationship with temperature variation from -15 °C to 100 °C and no outcome dependency from the specific tissue. On the contrary, such technique is deeply affected by the motion artifacts caused by the cardiac and respiratory activities. Algorithms devoted to artifacts elimination could be exploited, but at the expense of the computational costs. Moreover, the need to operate within the MRI room, thus to use MRI-compatible surgical tools limited its use in the clinical practice. Also, the costs (of both MR scanners and specific sequences for thermometry) are higher compared to IR- and ultrasound-based imaging.

Regarding ultrasound thermometry, ultrasound probes are significantly less expensive than MRI scanners and also readily available in hospital environments. Good accuracy and spatial resolution can be achieved by carefully choosing both a performing motion algorithm and a proper ultrasound pulse. However, once again, this entails an incrementation of the computational costs. The main drawbacks related to this technology are: the large occurrence of motion artifacts caused by the organs’ physiological activity (i.e., cardiac and breathing activities) and the measurement errors due to the change in the speed of sound in tissues exposed to temperatures greater than 50 ° C.

Finally, IR imaging provides real-time bidimensional color-coded map of easy interpretation. To date, this technique has been mostly exploited for preventing unwanted injuries of anatomical structure surrounding the myocardium (e.g., esophageal lumen), instead of reconstructing the cardiac temperature. This is because the IR system (which is a catheter) needs to be placed in correspondence of the measurement site. Nevertheless, this method is highly affected by the surrounding environment as instruments and operators could invade the scanning field thus distorting the measurements. Moreover, no information regarding the inner layers of the treated tissues is provided by the usage of this technique.

Despite the obvious advantages brought by these approaches (e.g., uncluttered surgical field and lack of additional devices to be managed), to date they are not in widespread usage due to many limitations that technology has not yet managed to overcome [71][72].

A summary of the main features that characterize each invasive and non-invasive technique is reported in Table 1.

Table 1. Summary of the features of the presented invasive and non-invasive solutions devoted to temperature monitoring of myocardial tissue undergoing RFA.

Technology

Features

Thermocouples

Invasive; single point measurement; accuracy of approximately 1 °C; robust; well-known technology; almost constant sensitivity in a wide range of T; adequate dynamic response considering the application of interest; wide measuring interval; potential presence of measurement artifacts

Thermistors

Invasive; single point measurement; accuracy better than 0.3 °C; robust; well-known technology; high sensitivity but can significantly decrease for high T; adequate dynamic response considering the application of interest; wide measuring interval; potential presence of measurement artifacts

Fluoroptic Sensors

Invasive; single point measurement; accuracy up to 0.2 °C; fragile; adequate dynamic response considering the application of interest; wide measuring interval; immunity to electromagnetic interference

FBGs

Invasive; multi-point measurement with resolution even better than 1 mm; accuracy even higher than 0.1 °C; fragile; adequate dynamic response considering the application of interest; wide measuring interval; constant sensitivity in a wide range of T; immunity to electromagnetic interferences; potential presence of motion artifacts

MRI

Non-invasive; 3D distribution of T; sensitivity up to -0.01 ppm·°C-1; constant sensitivity for T from -15 °C to 100 °C; no tissue dependency in case of PRFST; temporal resolution better than 2 s; motion artifacts  

Ultrasound Imaging

Non-invasive; 3D distribution of T; easy supply in clinical settings; high T resolution at high computational costs; motion artifacts; decrease in sensitivity for T up to 50 ° C

IR Imaging

Non-invasive; 2D color-coded T map; no T information regarding the inner layers; measurement affected by the surrounding environment

In conclusion, several methods for temperature monitoring during cardiac RFA have been presented and their main benefits and burdens have been discussed. Currently, no approach can be considered free from drawbacks, so suggesting that the choice of a particular solution depends on the needs and the circumstances related to the individual case.

In clinical practice, the currently used techniques do not allow to perform real temperature estimation inside the myocardial tissue. In fact, thermocouples and thermistors held into the emitting antennas provide temperature at the antenna tip which only gives indirect information for the adjustment of the treatment parameters. Also, infrared probes placed in the esophagus lumen are used solely to avoid unwanted injuries.

In the clinical scenario, despite its limits, MRI-based thermometry might be the most promising technique for myocardial thermal monitoring, given the already positive results present in the literature.

In the research context, invasive systems are exploited for in vivo and ex vivo experiments to characterize the effects of cardiac RFA by means of internal temperature analysis. Moreover, such competences may help in optimizing myocardial RFA procedure on patients, also making the process safer. Among other techniques, FBGs present enormous potential and could have a major impact in analyzing the influence of different parameters (i.e., temperature, power delivered, treatment time and force exerted by the catheter) on the mechanism of lesion formation. By means of this new awareness, a relevant contribution would be given on the performance evaluation of new RF delivery devices, as well as on the design guidance of these tools.

Finally, it is possible to conclude that at present finding the optimal measurement system for myocardial temperature monitoring during RFA is still an open challenge.

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Martina Zaltieri
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