1. Introduction

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]–[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].

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]–[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].

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 [13], being able to monitor the temperature increase of the treated tissue during cardiac RFA may be of fundamental importance.

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 cm

²

) [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:

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:

Given that j decreases as 1/r

²

, where r is the distance from the catheter application point, p drops within the tissue as 1/r

⁴

[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]:

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]:

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

2. 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.

Concerning temperature monitoring during RFA cancer treatments, MRI thermometry [16]–[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]–[23] and invasive methods, such as thermocouples [24], [25] thermistors [26]–[29], and fiber optics [30]–[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].

Focusing on myocardial RFA, thermocouples and thermistors embedded in RF emitting antennas have been widely adopted both in the clinical practice [34]–[41] and in the experimental field [42]–[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]–[49] or thermistors [50]–[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 addition, more than a decade later, fluoroptic probes [57]–[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]}.

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

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.

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

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.

References:

[1] S. K. Huang, S. Bharati, A. R. Graham, M. Lev, F. I. Marcus, and R. C. Odell, “Closed chest catheter desiccation of the atrioventricular junction using radiofrequency energy—a new method of catheter ablation,” J. Am. Coll. Cardiol., vol. 9, no. 2, pp. 349–358, 1987.

[2] T. Budde, G. Breithardt, M. Borggrefe, A. Podczeck, and J. Langwasser, “Initial experiences with high-frequency electric ablation of the AV conduction system in the human,” Z. Kardiol., vol. 76, no. 4, pp. 204–210, 1987.

[3] M. Borggrefe, T. Budde, A. Podczeck, and G. Breithardt, “High frequency alternating current ablation of an accessory pathway in humans,” J. Am. Coll. Cardiol., vol. 10, no. 3, pp. 576–582, 1987.

[4] S. Nath, J. P. DiMARCO, and D. E. Haines, “Basic aspects of radiofrequency catheter ablation,” J. Cardiovasc. Electrophysiol., vol. 5, no. 10, pp. 863–876, 1994.

[5] J. L. Dinerman, R. D. Berger, and H. Calkins, “Temperature monitoring during radiofrequency ablation,” J. Cardiovasc. Electrophysiol., vol. 7, no. 2, pp. 163–173, 1996.

[6] J. F. Viles-Gonzalez, E. Berjano, and A. d’Avila, “Complications of radiofrequency catheter ablation: can we prevent steam pops?” JACC: Clinical Electrophysiology, 2018.

[7] M. Tokuda et al., “Outcomes of cardiac perforation complicating catheter ablation of ventricular arrhythmias,” Circ. Arrhythmia Electrophysiol., vol. 4, no. 5, pp. 660–666, 2011.

[8] D. E. Haines and A. F. Verow, “Observations on electrode-tissue interface temperature and effect on electrical impedance during radiofrequency ablation of ventricular myocardium.,” Circulation, vol. 82, no. 3, pp. 1034–1038, 1990.

[9] F. Morady, “Radio-frequency ablation as treatment for cardiac arrhythmias,” N. Engl. J. Med., vol. 340, no. 7, pp. 534–544, 1999.

[10] C. V Burton, “RF lesion generation,” Stereotact. Funct. Neurosurg., vol. 39, no. 2, pp. 77–79, 1976.

[11] L. W. Organ, “Electrophysiologic principles of radiofrequency lesion making,” Stereotact. Funct. Neurosurg., vol. 39, no. 2, pp. 69–76, 1976.

[12] D. E. Haines, D. D. Watson, and A. F. Verow, “Electrode radius predicts lesion radius during radiofrequency energy heating. Validation of a proposed thermodynamic model.,” Circ. Res., vol. 67, no. 1, pp. 124–129, 1990.

[13] R. J. Stafford, D. Fuentes, A. A. Elliott, J. S. Weinberg, and K. Ahrar, “Laser-induced thermal therapy for tumor ablation,” Crit. Rev. Biomed. Eng., vol. 38, no. 1, 2010.

[14] M. E. Ulucakli, “Simulation of Radiofrequency Ablation and Thermal Damage to Tissue,” in Proceedings of the IEEE 32nd Annual Northeast Bioengineering Conference, 2006, pp. 93–94.

[15] S. Labonte, “Numerical model for radio-frequency ablation of the endocardium and its experimental validation,” IEEE Trans. Biomed. Eng., vol. 41, no. 2, pp. 108–115, 1994.

[16] M. Lepetit‐Coiffé et al., “Real‐time monitoring of radiofrequency ablation of rabbit liver by respiratory‐gated quantitative temperature MRI,” J. Magn. Reson. Imaging An Off. J. Int. Soc. Magn. Reson. Med., vol. 24, no. 1, pp. 152–159, 2006.

[17] S. Terraz et al., “Radiofrequency ablation of small liver malignancies under magnetic resonance guidance: progress in targeting and preliminary observations with temperature monitoring,” Eur. Radiol., vol. 20, no. 4, pp. 886–897, 2010.

[18] J. Gellermann et al., “Methods and potentials of magnetic resonance imaging for monitoring radiofrequency hyperthermia in a hybrid system,” Int. J. Hyperth., vol. 21, no. 6, pp. 497–513, 2005.

[19] F. Primavesi et al., “Thermographic real-time-monitoring of surgical radiofrequency and microwave ablation in a perfused porcine liver model,” Oncol. Lett., vol. 15, no. 3, pp. 2913–2920, 2018.

[20] X. Geng et al., “Comparison of ultrasound temperature imaging with infrared thermometry during radio frequency ablation,” Jpn. J. Appl. Phys., vol. 53, no. 4, p. 47001, 2014.

[21] Y.-D. Liu et al., “Adaptive ultrasound temperature imaging for monitoring radiofrequency ablation,” PLoS One, vol. 12, no. 8, p. e0182457, 2017.

[22] M. J. Daniels and T. Varghese, “Dynamic frame selection for in vivo ultrasound temperature estimation during radiofrequency ablation,” Phys. Med. Biol., vol. 55, no. 16, p. 4735, 2010.

[23] T. Varghese et al., “Ultrasound monitoring of temperature change during radiofrequency ablation: preliminary in-vivo results,” Ultrasound Med. Biol., vol. 28, no. 3, pp. 321–329, 2002.

[24] J. D. Crowley, J. Shelton, A. J. Iverson, M. P. Burton, N. C. Dalrymple, and J. T. Bishoff, “Laparoscopic and computed tomography-guided percutaneous radiofrequency ablation of renal tissue: acute and chronic effects in an animal model,” Urology, vol. 57, no. 5, pp. 976–980, 2001.

[25] X. Buy, C.-H. Tok, D. Szwarc, G. Bierry, and A. Gangi, “Thermal protection during percutaneous thermal ablation procedures: interest of carbon dioxide dissection and temperature monitoring,” Cardiovasc. Intervent. Radiol., vol. 32, no. 3, pp. 529–534, 2009.

[26] F. E. Diehn, Z. Neeman, J. L. Hvizda, and B. J. Wood, “Remote thermometry to avoid complications in radiofrequency ablation,” J. Vasc. Interv. Radiol., vol. 14, no. 12, pp. 1569–1576, 2003.

[27] E. J. Patterson, C. H. Scudamore, D. A. Owen, A. G. Nagy, and A. K. Buczkowski, “Radiofrequency ablation of porcine liver in vivo: effects of blood flow and treatment time on lesion size.,” Ann. Surg., vol. 227, no. 4, p. 559, 1998.

[28] D. E. Dupuy, R. Hong, B. Oliver, and S. N. Goldberg, “Radiofrequency ablation of spinal tumors: temperature distribution in the spinal canal,” Am. J. Roentgenol., vol. 175, no. 5, pp. 1263–1266, 2000.

[29] W. Fan, X. Li, L. Zhang, H. Jiang, and J. Zhang, “Comparison of microwave ablation and multipolar radiofrequency ablation in vivo using two internally cooled probes,” Am. J. Roentgenol., vol. 198, no. 1, pp. W46–W50, 2012.

[30] D. Tosi et al., “Fiber-optic chirped FBG for distributed thermal monitoring of ex-vivo radiofrequency ablation of liver,” Biomed. Opt. Express, vol. 5, no. 6, pp. 1799–1811, 2014.

[31] P. Saccomandi, E. Schena, M. Diana, F. M. Di Matteo, G. Costamagna, and J. Marescaux, “Multipoint temperature monitoring in liver undergoing computed tomography-guided radiofrequency ablation with fiber Bragg grating probes,” in 2016 38th Annual International Conference of the IEEE Engineering in Medicine and Biology Society (EMBC), 2016, pp. 5174–5179.

[32] D. Tosi, E. G. Macchi, and A. Cigada, “Fiber-optic temperature and pressure sensors applied to radiofrequency thermal ablation in liver phantom: methodology and experimental measurements,” J. Sensors, vol. 2015, 2015.

[33] E. G. Macchi, D. Tosi, G. Braschi, M. Gallati, A. Cigada, and E. Lewis, “Optical fiber sensors-based temperature distribution measurement in ex vivo radiofrequency ablation with submillimeter resolution,” J. Biomed. Opt., vol. 19, no. 11, p. 117004, 2014.

[34] H. Calkins, E. Prystowsky, M. Carlson, L. S. Klein, J. P. Saul, and P. Gillette, “Temperature monitoring during radiofrequency catheter ablation procedures using closed loop control. Atakr Multicenter Investigators Group.,” Circulation, vol. 90, no. 3, pp. 1279–1286, 1994.

[35] E. Anter et al., “A lattice-tip temperature-controlled radiofrequency ablation catheter for wide thermal lesions: first-in-human experience with atrial fibrillation,” JACC Clin. Electrophysiol., vol. 6, no. 5, pp. 507–519, 2020.

[36] L. T. Blouin, F. I. Marcus, and L. Lampe, “Assessment of effects of a radiofrequency energy field and thermistor location in an electrode catheter on the accuracy of temperature measurement,” Pacing Clin. Electrophysiol., vol. 14, no. 5, pp. 807–813, 1991.

[37] L. A. Pires, S. K. S. Huang, A. B. Wagshal, R. S. Mittleman, and W. J. Rittman, “Temperature-guided radiofrequency catheter ablation of closed-chest ventricular myocardium with a novel thermistor-tipped catheter,” Am. Heart J., vol. 127, no. 6, pp. 1614–1618, 1994.

[38] I. A. N. D. McRURY, J. G. Whayne, and D. E. Haines, “Temperature measurement as a determinant of tissue heating during radiofrequency catheter ablation: an examination of electrode thermistor positioning for measurement accuracy,” J. Cardiovasc. Electrophysiol., vol. 6, no. 4, pp. 268–278, 1995.

[39] J. J. Langberg et al., “Temperature monitoring during radiofrequency catheter ablation of accessory pathways.,” Circulation, vol. 86, no. 5, pp. 1469–1474, 1992.

[40] S. Willems et al., “Temperature-controlled radiofrequency catheter ablation of manifest accessory pathways,” Eur. Heart J., vol. 17, no. 3, pp. 445–452, 1996.

[41] A. Thiagalingam, A. d’Avila, C. McPHERSON, Z. Malchano, J. Ruskin, and V. Y. Reddy, “Impedance and temperature monitoring improve the safety of closed‐loop irrigated‐tip radiofrequency ablation,” J. Cardiovasc. Electrophysiol., vol. 18, no. 3, pp. 318–325, 2007.

[42] G. Hindricks et al., “Radiofrequency coagulation of ventricular myocardium: improved prediction of lesion size by monitoring catheter tip temperature,” Eur. Heart J., vol. 10, no. 11, pp. 972–984, 1989.

[43] H. H. Petersen, X. Chen, A. Pietersen, J. H. Svendsen, and S. Haunsø, “Temperature-controlled radiofrequency ablation of cardiac tissue: an in vitro study of the impact of electrode orientation, electrode tissue contact pressure and external convective cooling,” J. Interv. Card. Electrophysiol., vol. 3, no. 3, pp. 257–262, 1999.

[44] H. H. Petersen, X. Chen, A. Pietersen, J. H. Svendsen, and S. Haunsø, “Lesion dimensions during temperature-controlled radiofrequency catheter ablation of left ventricular porcine myocardium: impact of ablation site, electrode size, and convective cooling,” Circulation, vol. 99, no. 2, pp. 319–325, 1999.

[45] M. K. Jain and P. D. Wolf, “In Vitro Temperature Map of Cardiac Ablation Demonstrates the Effect of Flow on Lesion Development,” Ann. Biomed. Eng., vol. 28, pp. 1066–1074, 2000, doi: https://doi.org/10.1114/1.1310218.

[46] H. Cao, V. R. Vorperian, J.-Z. Tsai, S. Tungjitkusolmun, E. J. Woo, and J. G. Webster, “Temperature measurement within myocardium during in vitro RF catheter ablation,” IEEE Trans. Biomed. Eng., vol. 47, no. 11, pp. 1518–1524, 2000.

[47] O. J. Eick and D. Bierbaum, “Tissue Temperature‐Controlled Radiofrequency Ablation,” Pacing Clin. Electrophysiol., vol. 26, no. 3, pp. 725–730, 2003.

[48] U. Halm et al., “Thermal esophageal lesions after radiofrequency catheter ablation of left atrial arrhythmias,” Am. J. Gastroenterol., vol. 105, no. 3, pp. 551–556, 2010.

[49] P. Halbfass et al., “Incidence of asymptomatic oesophageal lesions after atrial fibrillation ablation using an oesophageal temperature probe with insulated thermocouples: a comparative controlled study,” Ep Eur., vol. 19, no. 3, pp. 385–391, 2017.

[50] P. Kovoor et al., “Comparison of radiofrequency ablation in normal versus scarred myocardium,” J. Cardiovasc. Electrophysiol., vol. 17, no. 1, pp. 80–86, 2006.

[51] D. P. Redfearn et al., “Esophageal temperature monitoring during radiofrequency ablation of atrial fibrillation,” J. Cardiovasc. Electrophysiol., vol. 16, no. 6, pp. 589–593, 2005.

[52] I. Rodríguez, J. L. Lequerica, E. J. Berjano, M. Herrero, and F. Hornero, “Esophageal temperature monitoring during radiofrequency catheter ablation: Experimental study based on an agar phantom model,” Physiol. Meas., 2007, doi: 10.1088/0967-3334/28/5/001.

[53] D. Chakraborty and I. Brezovich, “Error sources affecting thermocouple thermometry in RF electromagnetic fields,” J. Microw. Power, vol. 17, no. 1, pp. 17–28, 1982.

[54] D. P. Chakraborty and I. A. Brezovich, “A source of thermocouple error in radiofrequency electric fields,” Electron. Lett., vol. 16, no. 22, pp. 853–854, 1980.

[55] D. A. Christensen, “Thermal dosimetry and temperature measurements,” Cancer Res., vol. 39, no. 6 Part 2, pp. 2325–2327, 1979.

[56] P. R. N. Childs, J. R. Greenwood, and C. A. Long, “Review of temperature measurement,” Rev. Sci. Instrum., vol. 71, no. 8, pp. 2959–2978, 2000.

[57] M. A. Wood, K. M. Shaffer, A. L. Ellenbogen, and E. D. Ownby, “Microbubbles during radiofrequency catheter ablation: composition and formation,” Hear. Rhythm, vol. 2, no. 4, pp. 397–403, 2005.

[58] I. A. Thyer, P. Kovoor, M. A. Barry, J. I. M. Pouliopoulos, D. L. Ross, and A. Thiagalingam, “Protection of the coronary arteries during epicardial radiofrequency ablation with intracoronary chilled saline irrigation: assessment in an in vitro model,” J. Cardiovasc. Electrophysiol., vol. 17, no. 5, pp. 544–549, 2006.

[59] I. Watanabe et al., “Temperature-controlled cooled-tip radiofrequency ablation in left ventricular myocardium,” Int. Heart J., vol. 51, no. 3, pp. 193–198, 2010.

[60] M. Zaltieri, G. Allegretti, C. Massaroni, E. Schena, and F. M. Cauti, “Fiber Bragg Grating Sensors for Millimetric-Scale Temperature Monitoring of Cardiac Tissue Undergoing Radiofrequency Ablation: A Feasibility Assessment,” Sensors, vol. 20, no. 22, p. 6490, 2020.

[61] C. H. Seo et al., “The feasibility of using thermal strain imaging to regulate energy delivery during intracardiac radio-frequency ablation,” IEEE Trans. Ultrason. Ferroelectr. Freq. Control, vol. 58, no. 7, pp. 1406–1417, 2011.

[62] M. Wood et al., “Direct measurement of the lethal isotherm for radiofrequency ablation of myocardial tissue,” Circ. Arrhythmia Electrophysiol., vol. 4, no. 3, pp. 373–378, 2011.

[63] M. G. Daly, I. Melton, G. Roper, G. Lim, and I. G. Crozier, “High-resolution infrared thermography of esophageal temperature during radiofrequency ablation of atrial fibrillation,” Circ. Arrhythmia Electrophysiol., vol. 11, no. 2, p. e005667, 2018.

[64] J. P. Hummel et al., “Atrial Fibrillation Thermographic and Endoscopic Monitoring of Patients: Safety Algorithm for the Esophagus: AF TEMP-SAFE Study,” Circ. Arrhythmia Electrophysiol., vol. 11, no. 12, p. e006814, 2018.

[65] R. T. Borne and D. T. Nguyen, “Red Alert: Infrared Thermography for Esophageal Monitoring.” Am Heart Assoc, 2018.

[66] A. Kolandaivelu, M. M. Zviman, V. Castro, A. C. Lardo, R. D. Berger, and H. R. Halperin, “Noninvasive assessment of tissue heating during cardiac radiofrequency ablation using MRI thermography,” Circ. Arrhythmia Electrophysiol., vol. 3, no. 5, pp. 521–529, 2010.

[67] B. D. De Senneville, S. Roujol, P. Jaïs, C. T. W. Moonen, G. Herigault, and B. Quesson, “Feasibility of fast MR‐thermometry during cardiac radiofrequency ablation,” NMR Biomed., vol. 25, no. 4, pp. 556–562, 2012.

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.

[68] S. Toupin et al., “Feasibility of real-time MR thermal dose mapping for predicting radiofrequency ablation outcome in the myocardium in vivo,” J. Cardiovasc. Magn. Reson., vol. 19, no. 1, pp. 1–12, 2017.

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.

[69] V. Ozenne et al., “Improved cardiac magnetic resonance thermometry and dosimetry for monitoring lesion formation during catheter ablation,” Magn. Reson. Med., vol. 77, no. 2, pp. 673–683, 2017.

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.

[70] S. Toupin et al., “Combination of principal component analysis and optical-flow motion compensation for improved cardiac MR thermometry,” Phys. Med. Biol., vol. 62, no. 4, p. 1208, 2017.

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.

[71] V. Rieke and K. Butts Pauly, “MR thermometry,” J. Magn. Reson. Imaging An Off. J. Int. Soc. Magn. Reson. Med., vol. 27, no. 2, pp. 376–390, 2008.

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.

[72] V. Rieke, K. K. Vigen, G. Sommer, B. L. Daniel, J. M. Pauly, and K. Butts, “Referenceless PRF shift thermometry,” Magn. Reson. Med. An Off. J. Int. Soc. Magn. Reson. Med., vol. 51, no. 6, pp. 1223–1231, 2004.