A thesis submitted in partial fulfilment of the requirement of Nazarbayev University for the degree of

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FIBER OPTIC SENSORS FOR IN SITU MONITORING DURING THERMAL ABLATION OF TUMORS

Madina Jelbuldina

A thesis submitted in partial fulfilment of the requirement of Nazarbayev University for the degree of

Doctor of Philosophy in Science, Engineering and Technology

July 2021

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2 ABSTRACT

High-temperature tumor ablation, or hyperthermia, is a minimally invasive therapy applied to treat benign and malignant tumors in different organs, most commonly for small and mid-size hepatic tumors.

The modern clinical standards determine 60 °C as a threshold for the almost immediate thermal coagulation of biological tissues. In order to perform the procedure effectively and avoid burning of healthy tissue, the limitations of ablation procedure have to be overcome. The limitations are related to incomplete ablation, the lack of the real time control of the procedure and inability to treat large size tumors. Real-time monitoring of the temperature dose within the target zone provides information about the amount of damaged tissue, hence allowing a clinician to regulate TA settings and to control procedural outcomes.

In this work two technological solutions is proposed to overcome the limitations of TA: (1) real-time detection of the temperature profiles in situ by means of fiber optic sensors that offer significant advantages over conventional thermometry techniques – thermocouples and imaging; (2) nanoparticles introduced in situ, which can mediate the thermal treatment by stabilizing and modifying the thermal, electrical, and optical properties of the tissue and increase the size of treatable tumors.

The overall goal of the thesis was to develop fiber optic based sensing for the multi-point real time temperature monitoring during thermal ablation.

This thesis presents a comparative study of sensing performance of 3 fiber optic sensing technologies, for the application of RF ablation. Temperature distribution during HIFU thermal treatment is performed ex vivo on a sample of breast fibroadenoma with the help of FBG sensors.

Next goal is to investigate the effect of nanoparticles on the outcomes of ex vivo TA with in-situ thermal profiling by means of fiber optic sensors. FBG based sensory system was implemented to investigate the effects of magnetite (Fe3O4) nanoparticles with 2 mg/mL and 5 mg/mL concentrations on RFA and MWA procedures. Results are presented in the forms of thermal

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maps, reporting the extension of lethal 60 °C isotherm by 20% for MWA, and by 60% for RFA.

The work presents results on thermal profiling by means of non-standard optical fibers, such as polymer fibers and fibers with enhanced scattering profile. Sensing performance of the chirped fiber Bragg grating fabricated on microstructured polymer optical fiber is investigated and validated during ex vivo RF ablation.

Finally, the thesis reports a novel setup based on MgO doped optical fibers that provides high resolution two dimensional temperature monitoring of thermal ablation. The setup utilizes specialty optical fibers with enhanced scattering profile (around 40 dB with respect to SMF) and allows for the OBR/OFDR interrogation of multiple fibers with a single scan. The proposed multiplexing configuration is validated in ex vivo laser ablation of liver phantom. Results of temperature measurements are two-dimensional thermal maps exhibiting high spatial resolution of 2.5 mm.

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Figure 3.6 Photo of RF hyperthermia and fiber optic sensors interrogation setup, arranged on the bench. On the left side, the OBR (Luna OBR4600) with its control PC. In the center, the RF generator with the active electrode probe, and the metallic plate acting as passive electrode. On the right, a custom- made interrogator made by SLED and spectrometer with its control PC, for dual detection of FBGs and CFBG.

Figure 3.7 Schematic of the experimental setup based on RFA hyperthermia and fiber optics sensing systems.

Figure 3.8 (a) Photographic image and (b) schematics of positioning of RF probe and three fiber optic sensors in the phantom tissue prior to RFA heating.

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Figure 3.9 Reflected spectra of the CFBG and of the 5-element FBG array, detected on the spectrometer.

Figure 3.10 CFBG demodulation technique: first part algorithm - preliminary simulations of reflection spectra.

Figure 3.11 CFBG demodulation technique: second part algorithm – identification of gradient components during thermal ablation.

Figure 3.12 Temperature data recorded by 5 FBGs

Figure 3.13 Peak temperatures recorded in real time by FBG array, CFBG and distributed sensing methodologies during RFA.

Figure 3.14 Thermal map obtained with OBR distributed sensor during RFA. X axis shows duration of ablation. Y axis shows the sensor length with Y = 0 coordinate corresponding to the RFA tip.

Figure 3.15 Thermal map obtained with 5 FBGs during RFA. X axis shows duration of ablation. Y axis shows the sensor length in cm, corresponding to the length of each FBG 5 mm, and 5 mm spacing between the gratings.

Figure 3.16 Thermal map obtained with CFBG during RFA. X axis shows duration of ablation. Y axis shows the sensor length in cm.

Figure 4.1 X-Ray powder diffractogram of synthesized magnetite nanoparticles.

Figure 4.2 AFM topographical images in air of the synthesized magnetite nanoparticles (MNP); the data report the AFM images (2.5 m  2.5 m) with 2D view. A) Mica substrate in absence of MNPs; B) images of MNPs diluted in solution 1:10 concentration water buffer; C) images of MNPs diluted in solution 1:100 concentration water buffer. The synthesized MNP are 8 nm - 40 nm size.

Figure 4.3 Photographic image of the (a) RF ablation and FBG measurement setup; (b) Custom-made interrogation system.

Figure 4.4 a) Schematics of the experimental setup for MNP-enhanced RFA;

b) Positioning of the 15 FBGs arranged in 3 arrays on the xy plane.

Figure 4.5 Spectra of the 15 FBGs arranged in 3 arrays, acquired every 20 s during a RFA experiment.

Figure 4.6 Calibration functions for the FBG sensors; the chart reports the wavelength shift as a function of the applied temperature for the first array of the FBGs during a temperature cycle in a water bath.

Figure 4.7 Photographs of the RF-ablated lesions in the liver phantom, in absence of NP and after injection of 5 mg/mL concentration. The ruler shows cm units.

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Figure 4.8 Tissue impedance recorded with the impedance meter throughout the RFA.

Figure 4.9 Temperatures recorded by FBG sensors during RFA (no nanoparticles). The legend shows each of 15 FBGs label according to its position inside the 3 arrays.

Figure 4.10 Thermal maps for RFA experiments (no nanoparticles injected).

Temperature levels are presented in °C as a function of x and time for the first (left, y = 0 mm), second (center, y = 5 mm) and third (right, y = 10 mm) FBG arrays.

Figure 4.11 Temperature recorded by FBG sensors during MNP-enhanced RF ablation (5 mg/mL density of MNPs). The legend shows each of 15 FBGs label according to its position inside the 3 arrays.

Figure 4.12 Thermal maps for MNP-mediated RFA experiments (5 mg/mL concentration). Temperature levels are presented in °C as a function of x and time for the first (left, y = 0 mm), second (center, y = 5 mm) and third (right, y = 10 mm) FBG arrays. Double arrows indicate the diameter of 60 °C isotherm.

Figure 4.13 Photographic images of the tissue undergoing MWA: a) no NPs injected, b) with 2 mg/mL NPs concentration, and c) with 5 mg/mL NPs concentration. The ruler shows cm units.

Figure 4.14 Temperature recorded by FBG sensors during MW ablation (no nanoparticles injected). The legend shows each FBG label according to its position inside the 3 arrays.

Figure 4.15 Thermal maps for MWA experiments (no nanoparticles injected).

Temperature levels are presented in °C as a function of x and time for the first (left, y = 0 mm), second (center, y = 5 mm) and third (right, y = 10 mm) FBG arrays.

Figure 4.16 Temperature recorded by FBG sensors during NP-enhanced MW ablation (2 mg/mL MNPs concentration). The legend shows each FBG label according to its position inside the array.

Figure 4.17 Thermal maps for NP-enhanced MWA experiments (2 mg/mL MNPs concentration). Temperature levels are presented in °C as a function of x and time for the first (left, y = 0 mm), second (center, y = 5 mm) and third (right, y = 10 mm) FBG arrays.

Figure 4.18 Temperatures recorded by 15 FBG sensors during MNP-mediated MW ablation (5 mg/mL MNPs concentration). The legend shows each FBG label according to its position inside the array.

Figure 5.1 Experimental setup for the chirped Bragg grating inscription by means of KrF excimer laser and a uniform phase mask [3].

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Figure 5.2 Reflection spectrum (left) and group delay (right) of mPOF CFBG in reference condition, before exposure to thermal gradients [3].

Figure 5.3 Reflection spectrum (left) and group delay (right) of mPOF CFBG during heating cycles in water bath.

Figure 5.4 mPOF CFBG central wavelength shift as a function of temperature;

the chart shows the experimental data and a linear fit.

Figure 5.5 Setup for the temperature detection experiment: a case of the linear gradient.

Figure 5.6 Results of the temperature profiling with the mPOF CFBG: upper chart: thermal map, lower chart: isotherms. The colorbar shows temperature in

°C degrees.

Figure 5.7 Schematic of thermal ablation experiment: the LUNA OBR measures spectra from mPOF CFBG, which is placed in proximity of RF applicator during the ablation.

Figure 5.8 Measurement of Gaussian temperature gradient with a mPOF CFBG: thermal profile reconstructed with the CFBG as a function of distance along grating and time during ablation. The colorbar shows temperature in ◦C degrees.

Figure 5.9 Temperature graphs for Gaussian-shaped RFA temperature profile;

the chart reports the temperature as a function of time, for different values of position along the grating length d.

Figure 6.1 Rayleigh scattering mechanism in the core of an optical fiber.

Figure 6.2 SEM cross section of the MgO-doped fibers: (a) M01 fiber, (c) G22 fiber. Simulation of LP01 mode shape of the (b) M01 and (d) G22 fibers. [4]

Figure 6.3 Backreflected power of the EBFs, as detected by OBR.

Figure 6.4 Temperature sensitivity of the MgO-doped fiber

Figure 6.5 Photograph of the laser ablation and EBF based distributed sensing setup. (a) View of the whole setup, including laser diode, and OBR- based fiber multiplexing setup. (b) Positioning of the 4 MgO-doped fibers on the liver phantom; and the laser output fiber fixed perpendicular to the liver surface.

Figure 6.6 Schematics of the laser ablation and EBF based distributed sensing setup. The fiber multiplexing setup consists of 3 splitters, 4 extenders and 4 MgO-doped fibers, positioned in parallel to each other at distance d = 5 mm.

Figure 6.7 P-I characteristics of the 980 nm diode laser.

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Figure 6.8 View of the laser beam focused on the liver phantom and the positioning of 4 MgO-doped fibers during LA (left); photo of the ablated tissue with 4 fibers in parallel, the distance between fibers is 5 mm.

Figure 6.9 Relative lengths of SMF extenders spliced to MgO-doped fibers Figure 6.10 Two-dimensional thermal maps reporting temperature on the XY plane for different elapsed time (20 s, 30 s, 50 s, 70 s, 100 s, and 120 s).

Figure 6.11 Photographic image of the phantom after performing two laser ablations: with and with no MNPs.

Figure 6.12 Two-dimensional thermal maps, reporting temperature distribution on the XY plane for different elapsed time (30 s, 50 s, 70 s, and 100 s). Top row – pristine ablation (no NPs), bottom row - with injected MNPs solution.

The plain considered is 15×40 mm, with 4 sensing elements spaced 5 mm on Y axis.

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13 LIST OF ABBREVIATIONS

AFM Atomic force microscopy

BDK Benzyl dimethyl ketal

CFBG Chirped FBG

CT Computer tomography

DTG Draw tower grating

EBF Enhanced backscattering fibers

EM Electromagnetic

et al. lat. et alii; meaning “and others”

FBG Fiber Bragg grating

FDA United States Food and Drug

Administration

HCC Hepatocellular carcinoma

HIFU High intensity focused ultrasound

LA Laser ablation

MNPs Magnetite nanoparticles

MRI Magnetic resonance imaging

MWA Microwave ablation

NA Numerical aperture

Nd:YAG Neodymium-doped yttrium aluminum

garnet

NPs Nanoparticles

OBR Optical backscatter reflectometer

OFDR Optical frequency-domain reflectometry

OTDR Optical time domain reflectometer

PMMA Polymethyl methacrylate

POF Polymer optical fibers

RFA Radiofrequency ablation

SEM Scanning electron microscope

SMF Single mode fiber

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TA Thermal ablation

TC Thermocouple

US Ultrasound

WDM Wavelength division multiplexing

XRD X-ray diffraction analysis

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15 AUTHOR’S DECLARATION

I hereby declare that the research presented in this thesis is the original work of the author, and that all sources used in researching it are fully acknowledged and all quotations properly identified. The thesis has not been previously submitted to this or any other university for a degree.

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16 ACKNOWLEDGEMENT

I would like to express my sincere gratitude to my supervisor, Prof Daniele Tosi for supporting me with invaluable advice and positive attitude, and always encouraging me to do research. I’m thankful to School of Engineering and Digital science of Nazarbayev University, for their financial support and research facilities.

During my PhD study it was a great pleasure to work with great researches and colleagues, who became my friends: Takhmina Ayupova, Marzhan Sypabekova, Madina Shaimerdenova, Zhannat Ashikbayeva. I would especially like to thank Sanzhar Korganbayev, who shared his experience with me and was my lab buddy during three years, and Alina Korobeynik, for the synthesis of nanoparticles. I’m very thankful to Aliya Bekmurzayeva, Marzhan Sypabekova, for always being helpful.

I extend my gratitude to all former and current members of Prof Daniele’s group, namely Prof Carlo Molardi, Sultan Sovetov, Aizhan Issatayeva, Aidana Beisenova, Kanat Dukenbayev (for helping with AFM measurements). I’m very thankful to our collaborators, namely Rui Min, Carlos Marques (fabrication of polymer CFBG), Zhazira Seidagaliyeva (clinical HIFU), Wilfried Blanc (fabrication of MgO-doped optical fibers), Alina Korobeynik (synthesis of nanoparticles) and co-supervisors (Zhandos Utegulov and Guido Perrone).

Thank you to Dr. Luis Rojas for improving our PhD program, it would not be possible to succeed in the program without your constant support and advice.

On a more personal level, I’m grateful to my friends for making these five years an interesting and exciting journey: Kamilya, Gulden, Bagda, Asem, Damira, Aigerim, Ainura.

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CHAPTER 1

Introduction

1.1 Motivation and background

Among different cancer cure procedures, surgical resection remains the only option to the patients diagnosed with primary and secondary hepatic tumours, however very limited patients would be recommended to such a procedure [5], [6]. The obstructed access to the tumor, and associated with it incomplete resection, as well as invasiveness of the procedure, and frequent lethal cases motivate clinicians to utilize alternative tumor treatments.

Treatment of tumors with high temperatures has been introduced into medical practice more than a century ago, but only starting from 1980s ablation techniques consolidated into an independent technology to cure benign and malignant tumors [1]. Since then, energy-based ablation has been improved and extended its applications thanks to development of laparoscopic medical devices and imaging techniques [7], [8].

Interstitial thermal ablative techniques are currently performed in medical practice as a minimally-invasive alternative to traditional surgery in the treatment of benign and malignant tumors. Nowadays, percutaneous thermal ablation, or thermotherapy, is primarily used for the treatment of small, unresectable tumours, including liver [9], [10], kidney, lung and bone cancers [11], as well as soft-tissue tumours of breast [12], [13], adrenal glands, and head.

Being a minimally invasive technology, thermal ablation has demonstrated more favorable outcomes compared to surgical treatment [14]. The main advantages of thermal coagulation therapies are possibility to treat patients who cannot have surgery: elderly people, patients with health issues, or patients who have multiple tumors of small size. Other important advantages must be mentioned, such as less injury of the surrounding healthy tissue, lower morbidity rates, lower costs and shorter recovery period [15].

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The term thermal ablation applies to the variety of treatments utilizing cooling or heating tumors within certain temperature ranges. Different types of cells are not equally susceptible to extreme temperatures, however, temperatures below −40°C and above 60°C are considered cytotoxic for most types of tissues [16]. Several types of energy sources are applied, such as radiofrequency (RF) current, microwave, laser, ultrasound and thermal conduction based devices. The most widely used thermal techniques in modern clinical practice are radiofrequency ablation (RFA) and microwave ablation (MWA), which employ high temperatures to induce cell necrosis, as well as cryoablation, that affects tumor through cooling down the tissue to low temperatures.

RFA technique has gained popularity due to its relative simplicity and the possibility to achieve large ablative lesions (up to 4-6 cm). This approach utilizes energy produced by RF generator and is mostly used in the treatment of hepatic tumors [17]. The simple configuration utilizes a percutaneous RF needle positioned into the target tissue, and a passive electrode, which is placed on the patient’s skin [18]. Similar to RFA, MWA relies on the use of electromagnetic waves to produce heat up to 100 °C and higher. Heat is transferred into the target region through a microwave antenna. The main advantage of MWA over RFA is the possibility of treatment tissues with higher impedance like lung and bone [7]. The most common medical application of MWA is ablation of hepatocellular carcinoma [19], [20].

The principle of recently introduced therapies, like laser ablation and high- intensity focused ultrasound (HIFU) is overall similar to other hyperthermia methods, but are less common and not so well studied, as RFA or MWA.

Among all high-temperature modalities the only non-invasive technique is HIFU, since it utilizes ultrasound waves, so no device has to be inserted into patient’s body [21]. These acoustic waves are focused then in the specific target region to elevate local temperatures up to 60 °C, which causes coagulative necrosis [22]. Laser ablation (LA) generates electromagnetic heating, as do RFA and MWA, with the advantage of laser precision and efficiency during laser ablation. LA devices consist of a solid-state laser or fiber laser, coupled into optical fiber, and positioned in contact with the tissue [23], [24]. Limitations of this modality are related to light scattering and

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absorption, so the ablation lesions are typically small and don’t exceed 2 cm². LA is used often for thyroid and brain tumor treatments, and its performance is substantially dependent upon the absorption coefficient of the tissue [25], [26].

According to the semi-empirical equations derived by Sapareto et al. cell damage during hyperthermia is strongly related not only to temperature increment, but to the exposure time [27]. The proposed concept has been mathematically formulated in the thermal dose (TD) relationships via the Arrhenius rate analysis [28]. TD relationships were applied to different cells and tissues and it was shown that the rate of heat-induced cell death is almost linearly related to exposition time and exponentially dependent on the temperature increment (within specific temperature ranges) [29]. In order to quantify the amount of thermal damage, in 1984 Sapareto and Dewey [22]

introduced the unit of thermal dose in the form of cumulative equivalent minutes at 43 °C (CEM43°C or t43). Although the overall damage to tissue depends on tissue sensitivity, which has variations across species and different tissues and organs, as well as temperature and exposure time, multiple experiments revealed the breakpoint in the rate of cell death to be around 43 °C [30]. The application of thermal dose and thermal damage values have been validated for different temperature ranges: thermal doses of 120–240 min at 43 °C generate considerable tissue necrosis, but the sensitivity between tissue types is variable. A lethal temperature threshold was estimated around 50-55 °C for short treatment times (less than 5 min) based on different methodological approaches, which was consistent with the TD concept [31]. Irreversible damage to vital compartments of cells occur at temperatures as high as 60 °C and is accompanied with almost immediate tissue coagulation. By reaching 100 °C and above, the tissue due to its mostly watery content, undergoes vaporization and carbonization. A useful diagram (Figure 1.1) demonstrates stages of thermal injury of the tissue induced by temperatures higher than body temperature [32]. For successful ablation, the tissue temperature should be maintained in the ideal range to ablate tumor tissue adequately and avoid carbonization around the tip of the electrode due to excessive heating. To achieve this goal, an accurate temperature monitoring is necessary in situ throughout the course of ablation.

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Figure 1.1 Biological response of the tissue to thermal effects. Adopted from [32].

In order to perform the TA procedure in a most efficient way, a combination of several factors have to be achieved. First is related to accurate positioning of the applicator inside the tumor, that is achieved through localization of the tumor and identification of its contours [33]. Next, the clinical settings of the device, such as power, duration, the amount of energy, should be adjusted for each patient not only prior to the ablation but throughout the course as well.

Such a control is impossible without real time monitoring of the temperature dose within the target region. However, the current ablation technologies lack accurate real time monitoring, and this issue impedes a full clinical integration of thermal therapies [17]. This limitation is caused by the fact that precise temperature data is missing during ablation therapy. Temperature increase is a function of the tissue properties, such as thermal and electrical conductivity, absorption coefficients, and blood flow, which vary from patient to patient and even within a single organ [34]. The effects of blood flow are more pronounced over the longer duration of ablation because of the heat sink effect and the local variations in perfusion [35].

Thus, there is a high demand in accurate temperature profiling and real time thermal mapping as this will serve the goal of overcoming technological limitations of hyperthermic procedures through performing several tasks:

i) Estimate the amount of energy delivered to the treated area, ii) Determine the volume of tissue that was coagulated;

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iii) Based on the thermal maps provide the clinician data to adjust the settings of the ablation and/or terminate the procedure [36]

iv) Avoid damage to structures adjacent to the target area.

The appropriate temperature monitoring technique should meet the following requirements:

 Biocompatibility with human body;

 Minimal invasiveness;

 Compatibility with ablation tools and other medical devices (generators, applicators etc.);

 Appropriate spatial and temporal resolution, and

 Sufficient temperature accuracy.

Several temperature measurement methodologies are applied and studied recently to guide energy-based treatments in research, and more recently in clinical practice. These approaches are divided into thermocouples and imaging-based modalities (magnetic resonance imaging (MRI)-based techniques, ultrasound (US) imaging and X-ray computer tomography (CT) imaging modalities).

Contact-type temperature sensors have to be placed into the target tissue, usually in the way when sensors are attached to or embedded in the percutaneous ablation RF or MWA applicators, and are inexpensive modality providing relatively good accuracy, about 1 °C. Thermocouples consist of two metallic wires, and commercially available systems are presented in the form of single sensor or multisensory thermistor-based or thermocouple-based systems. Despite their low cost, significant drawbacks of thermocouples are related to high thermal conductivity of metallic wires which causes undesirable heating of the sensor itself. Extra heating leads to incorrectness of temperature readings. The measurements are also limited in space (within the closest proximity to the RF or MWA needle), as well as in number of measuring points. It is not possible to obtain two- or three-dimensional temperature distribution based on the thermocouples systems.

MRI is considered a standard thermal monitoring technology used during TA treatments and is the only FDA approved modality for non-invasive thermal ablation monitoring [36], [37]. Despite high resolution of MRI thermometry and the possibility of obtaining a 3D temperature map, significant drawbacks

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related to MRI-based techniques are related to motion sensitivity and associated with it image distortions; non-compatibility with EM fields, meaning that most TA devices cannot be placed in the MR scanner. Requirements of high strength magnetic field and high costs of the equipment are also considered as disadvantages.

Another imaging modality commonly used for image guided ablation is ultrasound. The benefits of ultrasonic systems are their high temporal resolution, their availability, and their low cost. Still, ultrasound suffers from significant disadvantages, which include inferior image quality compared to MRI and CT, and limitations which arise from inability to image and treat with HIFU targets located behind overlying bones or structures containing gas, which obstruct the ultrasonic waves [38], [39].

X-ray CT, on the other hand, is a widely available imaging modality that offers high temporal resolution combined with high spatial resolution. It is substantially less expensive than MRI and provides images which are superior to ultrasound. Moreover, it can also image the brain and lungs, which ultrasound cannot, and does not impose any restrictions on the ablation equipment located within its vicinity as MRI does. Hence, it can potentially be utilized as a quantitative method for non-invasive monitoring of thermal ablation.

In relation to thermal ablation procedures, CT technique has been applied as a temperature monitoring tool during ex vivo RFA [40], [41], [40] and laser ablation [42], [43]. Variations of CT Hounsfield Unit with temperature as well as the effect of the tissue shrinking during ex vivo MWA was analyzed via CT imaging on the porcine phantoms [8], [44], [45]. All the experiments required a phantom to be placed in PMMA box [45] to reduce CT imaging artefacts, which is significant limitation for the real practical applications. Several investigations were focused on ex vivo HIFU ablation [46]. Brace et al.

performed a comparative study of RFA and MWA on in vivo swine lungs, but the problem of high radiation from X-ray presents another limitation of CT thermometry [47].

Fiber optic sensors (FOSs) thanks to their exceptional characteristics are getting interest for the sensing applications in biomedical procedures. Made up of silica, FOSs are biocompatible and almost inert to most of the chemicals

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[48], [49]. Due to the small diameter, around 100-200 microns, FOSs can be encapsulated into needles, flexible or semiflexible catheters and other surgical instruments [50]. Other important property is immunity to electromagnetic fields resulting in compatibility with MRI environment [51], and possibility to fabricate high-temperature resistant sensors. Fiber Bragg gratings (FBGs) have been reported as for multi-point temperature sensors during LA [52] as well as RFA, and MWA [53]. In addition, FBG arrays have been demonstrated to detect two dimensional thermal patterns of ablation procedures through wavelength-division multiplexing [54].

Moreover, chirped FBG (CFBG), particularly a linearly chirped FBG, has been recently used for biomedical applications in thermal ablation [3]. Its main advantage over a standard FBG array is the possibility for not only temporally but also spatially resolved measurements with a resolution as high as one millimeter over the grating length [16]. On the other hand, data analysis is associated with a more complex and time-consuming demodulation procedure [15, 16]. Recent works have been focused on developing quasi-distributed sensor systems [55] to achieve spatially resolved temperature sensing in radiofrequency ablation (RFA).

1.2 Objectives of the thesis

The overall goal of the thesis was to develop fiber optic based sensing for the multi-point real time temperature monitoring during thermal ablation. To achieve this goal, it has been divided into the following objectives:

1. Develop the FOS system to investigate thermal distribution during ex vivo radiofrequency ablation, compare the performance of different types of optical fiber sensors.

2. Study the temperature distribution during HIFU thermal treatment performed ex vivo on a sample of breast fibroadenoma with the help of FBG sensors.

3. Implement FBG based sensory system to investigate the effects of magnetite (Fe3O4) nanoparticles on RFA and MWA procedures.

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Compare the results of experiments: in absence and after injection of nanoparticles solution.

4. Study the sensing performance of the chirped fiber Bragg grating fabricated on microstructured polymer optical fiber.

5. Develop the fiber optic multiplexing setup for the distributed measurements by utilizing fibers with enhanced scattering properties 6. Perform two-dimensional temperature mapping during laser ablation on

ex vivo porcine liver by using non-standard fibers with enhanced scattering properties.

1.3 Structure of Thesis

Together with Introduction, this Thesis is divided into seven chapters.

Chapter 2 presents the overview, biophysical and technological aspects of minimally invasive thermal ablation procedures, and the challenges associated with them. Importance of real-time accurate monitoring during thermal ablation procedures is discussed. Finally, the Chapter covers common thermometry techniques used in TA, their limitations.

Chapter 3 is dedicated to role and application of optical fibers for the real time temperature monitoring during thermal therapies. Overview of important fiber optic sensors technologies is presented. Results on temperature profiling throughout the energy-based ablation procedures, like RFA, MWA, HIFU, are presented. Chapter 3 reports a comparative study of three fiber optic sensing methods, such as FBGs, CFBG and distributed sensor.

In Chapter 4 the effect of nanoparticles to improve the outcomes of TA is studied. Heat distribution during RFA and MWA is analyzed with the fiber Bragg grating sensory system and presented in the form of thermal maps.

Results of thermal mapping are reported for pristine ablation, and after injection of nanoparticles.

Chapter 5 reports a CFBG inscribed into a microstructured polymer optical fiber for the temperature sensing application.

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Chapter 6 presents results of the work involving specialty fiber optic sensors for the in situ monitoring during laser ablation. The Chapter reports a novel multiplexing setup based on enhanced backscattering fibers for the two- dimensional thermal profiling during laser ablation.

Thesis is then commenced with a Chapter 7, Conclusion and Future perspectives, which gives an overall summary of the work described in the above chapters as well as proposes possible future directions.

1.4 Role of collaborators

In an article on thermal profiling during HIFU ablation [56], methodology was developed by Zhazira Seidagaliyeva and Daniele Tosi; a fibroadenoma sample was prepared by Zhazira Seidagaliyeva; HIFU setup and ablation was performed by Zhazira Seidagaliyeva and Turlybek Tuganbekov; data analysis was performed by Sanzhar Korganbayev, Madina Jelbuldina and Sultan Sovetov.

For the Chapter 4 related to nanoparticle-mediated thermal ablation [57], [58]

conceptualization belongs to Daniele Tosi, Madina Jelbuldina and Alina Korobeynik. Design of methodology was done by Madina Jelbuldina, Sanzhar Korganbayev and Daniele Tosi. Magnetite nanoparticles synthesis was done by Alina Korobeynik. Data analysis was performed by Sanzhar Korganbayev, Madina Jelbuldina and Daniele Tosi.

For a Chapter 5, and a paper on temperature measurements with mPOF CFBG fabrication of mPOF CFBG was done by Rui Min and Carlos Marques;

design of experiment and setup was developed by Sanzhar Korganbayev, Madina Jelbuldina, Daniele Tosi; data analysis was performed by Sultan Sovetov, Sanzhar Korganbayev and Daniele Tosi.

The work on two-dimensional thermal mapping by means of enhanced backscattering optical fibers reported in Chapter 6 and published in [59], [60]

conceptualization belongs to Daniele Tosi and Carlo Molardi; fabrication of specialty fibers with enhanced scattering profile was performed by Wilfried Blanc; multiplexing setup was prepared by Madina Jelbuldina, Aidana Beisenova, Aizhan Issatayeva and Carlo Molardi; Laser ablation setup was

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prepared by Madina Jelbuldina and Azat Abdullayev; Data analysis was performed by Sanzhar Korganbayev, Arman Aitkulov and Daniele Tosi.

Formal analysis was performed by Madina Jelbuldina, Carlo Molardi and Daniele Tosi.

1.5 Research articles

1. M. Jelbuldina, A. Korobeinyk, S. Korganbayev, V.J. Inglezakis and D.

Tosi, “Fiber Bragg grating based temperature profiling in ferromagnetic nanoparticles-enhanced radiofrequency ablation”. Optical Fiber Technology, 43, pp.145-152 (2018).

2. M. Jelbuldina, A. Korobeinyk, S. Korganbayev, D. Tosi, K. Dukenbayev and V.J. Inglezakis, “Real-Time Temperature Monitoring in Liver During Magnetite Nanoparticle-Enhanced Microwave Ablation With Fiber Bragg Grating Sensors: Ex Vivo Analysis”, IEEE Sensors Journal, 18(19), pp.8005-8011 (2018).

3. M. Jelbuldina, S. Korganbayev, Z. Seidagaliyeva, S. Sovetov, T.

Tuganbekov and D. Tosi, “Fiber Bragg Grating Sensor for Temperature Monitoring During HIFU Ablation of Ex Vivo Breast Fibroadenoma”, IEEE Sensors Letters, 3(8), pp.1-4 (2019).

4. Z. Ashikbayeva, A. Aitkulov, M. Jelbuldina, A. Issatayeva, A.

Beisenova, C. Molardi, P. Saccomandi, W. Blanc, V.J. Inglezakis and D. Tosi, “Distributed 2D temperature sensing during nanoparticles assisted laser ablation by means of high-scattering fiber sensors”, Scientific Reports, 10(1), pp.1-12 (2020).

5. S. Korganbayev, R. Min, M. Jelbuldina, X. Hu, C. Caucheteur, O. Bang, B. Ortega, C. Marques and D. Tosi “Thermal profile detection through high-sensitivity fiber optic chirped Bragg grating on microstructured PMMA fiber,” Journal of Lightwave Technology, 36(20), pp.4723-4729 (2018).

6. A. Beisenova, A. Issatayeva, S. Sovetov, S. Korganbayev, M.

Jelbuldina, Z. Ashikbayeva, W. Blanc, E. Schena, S. Sales, C. Molardi and D. Tosi, “Multi-fiber distributed thermal profiling of minimally

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invasive thermal ablation with scattering-level multiplexing in MgO- doped fibers”, Biomedical optics express, 10(3), pp.1282-1296 (2019).

Conference Publications

1. M. Jelbuldina, A. Issatayeva, A. Aitkulov, S. Korganbayev, A.

Beisenova, N. Kulmukhanova, D. Koshen, C. Molardi and D. Tosi,

“Multi-fiber distributed temperature profiling in ex vivo magnetite nanoparticle-mediated laser tissue ablation”, In Optical Interactions with Tissue and Cells XXXI (Vol. 11238, p. 112380H). International Society for Optics and Photonics (2020 February).

2. M. Jelbuldina, S. Korganbayev, A. Korobeinyk, V.J. Inglezakis, and D. Tosi “Temperature Profiling of ex-vivo Organs during Ferromagnetic Nanoparticles-Enhanced Radiofrequency Ablation by Fiber Bragg Grating Arrays” In 2018 40th Annual International Conference of the IEEE Engineering in Medicine and Biology Society (EMBC) (pp. 1-4), IEEE (2018 July).

3. M. Jelbuldina, S. Korganbayev, A.V. Korobeinyk, V.J. Inglezakis and D. Tosi “Fiber Bragg grating sensors for temperature monitoring during nanoparticle-assisted microwave ablation” In Optical Fiber Sensors (p. WF3). Optical Society of America (2018 September).

4. A. Issatayeva, A. Beisenova, S. Sovetov, S. Korganbayev, M.

Jelbuldina, Z. Ashikbayeva, W. Blanc, C. Molardi and D. Tosi, “2D temperature sensing obtained by multiplexing of optical backscattering reflectometry,” In Optical Fibers and Sensors for Medical Diagnostics and Treatment Applications XX (Vol. 11233, p.

112330T), International Society for Optics and Photonics (February 2020).

5. A. Issatayeva, A. Beisenova, S. Sovetov, S. Korganbayev, M.

Jelbuldina, Z. Ashikbayeva, W. Blanc, E. Schena, S. Sales, C.

Molardi and D. Tosi, 2019, November, “Multiplexing of distributed temperature sensing achieved by nanoparticle-doped fibers,”

In Optics in Health Care and Biomedical Optics IX (Vol. 11190, p.

111900H), International Society for Optics and Photonics (November 2019).

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CHAPTER 2

Temperature monitoring during thermal therapies

Chapter 2 overviews modern thermal ablation modalities, their working principle, as well as technological limitations (Section 2.1). Next, motivation for the real-time temperature monitoring during thermal therapies is discussed (Section 2.2). Section 2.2 also overviews thermometry techniques which are currently applied in clinical practice, as well as the issues related to the conventional thermometry techniques.

2.1 Minimally invasive thermal therapies

2.1.1 General ablation principles

Tumor treatment with high temperatures has been introduced into medical practice more than a century ago, but it’s only in the 1980s when tumor ablation started to evolve as an independent technology to treat benign and malignant tumors. This became possible thanks to development of cross- sectional imaging techniques and laparoscopic medical devices [1]. Since then, percutaneous energy-based ablation has significantly evolved and extended its applications to treatment of various types benign and malignant tumor, including liver, kidney, lung and bone cancers, as well as soft-tissue tumours of breast, adrenal glands, and head [7], [10], [12], [61]. Nowadays, percutaneous thermal ablation is primarily used for the treatment of small, unresectable tumours or for patients who are poor surgical candidates, for instance, their overall health is weak or they have multiple tumors of small sizes.

The term thermal ablation, or hyperthermia applies to procedures which utilize cooling or heating tumors within certain temperature ranges, thus inducing lethal damage to the cancer cells. Different types of cells are not equally

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susceptible to extreme temperatures, however, temperatures below −40°C and above 60°C are considered cytotoxic for most types of tissues [16].

Being a minimally invasive technology, thermal ablation has demonstrated more favorable outcomes compared to surgical treatment, and at the same time it is therapeutically effective, for example, radiofrequency ablation for the treatment of hepatic tumor [14]. The overall advantages of ablation therapies are less injury of the surrounding healthy tissue, lower morbidity rates, lower costs and shorter recovery period [15]. Besides this the real time monitoring of the procedure through imaging, and the possibility to treat patients who cannot have surgery, such as elderly people, or patients who have multiple tumors of small size. The reasons for the growing interest in energy-based modalities and their increasing use in the clinical practice are related to several factors:

a. Different types of tumours are detected at an earlier stage;

b. The number of elderly patients is increasing;

c. Energy-based ablation technology is evolving and devices are improving, for example HIFU technology, cryoablation;

d. Imaging technologies are rapidly developed and are integrated into the minimally invasive hyperthermia systems;

e. Hyperthermia can improve the outcomes of radiotherapy and chemotherapy [62], [63].

These factors promote researchers of biomedical, engineering fields to further study thermal ablation techniques and to search the methods to improve the outcomes of the procedures. The efficacy and outcomes of thermal ablation have indeed to be improved due to drawbacks associated with it, such as limits of the ablative lesions [64], making the ablation suitable only for the small and mid-size tumors, as well as imperfection of the intra-procedural temperature monitoring techniques, that will be discussed thoroughly later in this chapter.

The most widely used thermal techniques in modern clinical practice are radiofrequency ablation (RFA) and microwave ablation (MWA), which employ high temperatures to induce cell necrosis, as well as cryoablation, that affects tumor through cooling down the tissue to low temperatures. Figure 2.1 shows the timeline of development of different image-guided TA techniques.

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Figure 2.1 Development and evolution of image-guided thermal ablation techniques [1].

While multiple techniques provide hyperthermia therapy, they all rely on several physical principles on delivering heat energy to the body:

(i) Thermal conduction of heat, which flows from higher to lower temperature at a rate dependent on the thermal gradient and thermal properties of all contacting material.

(ii) Resistive or dielectric losses from an applied electromagnetic (EM) field. At radiofrequencies below 20 MHz, an electric potential difference between inserted and surface contacting electrodes creates an electric current. The amount of induced heat is defined by Joule law and is proportional to the tissue impedance and the electric current: P = I²R.

(iii) Mechanical losses due to molecular oscillations induced by an applied ultrasound (US) acoustic wave.

Depending on the temperature range, tumor cell damage occurs through a certain mechanism. For example, increase in temperature up to 40°C will not induce a significant damage to the targeted tissue, due to the heat shock response. This is a response of the body characterized by increased blood flow that helps to prevent excess heating. Exposing the tissue to higher temperatures leads to random protein denaturation in cell nuclei and the subsequent cell death [28], [65]. Protein denaturation inside the nuclei is the predominant mechanism of the cell necrosis [66]. Applying thermal dosage of 42 °C – 46 °C for 10 minutes is sufficient to cause irreversible damage.

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Temperatures as high as 60°C generally mark the lethal isotherm since above that level plasma membrane is melting which leads to instantaneous cell necrosis.

Cryotherapy, or cryoablation, on the contrary, relies on killing pathological tissues through inducing osmotic shock [67]. Cooling down the tissue to

−40°C will destroy cellular metabolism.

Next sections of this Chapter are dedicated to mechanisms of each particular hyperthermia modality – RFA, MWA, HIFU and LA, as well as cryoablation.

The advantages and disadvantages of the procedures as well as the areas of clinical use are also the topic of further discussion. Limitations of thermal ablation procedures are also presented. Section 2.2 covers common thermometry techniques utilized during TA, their drawbacks are discussed as well.

2.1.2 Radiofrequency ablation

Radiofrequency (RF) thermal ablation is currently being implemented as a standard treatment replacing surgical resection for the treatment of malignant hepatic lesions and hepatocellular carcinoma (HCC) [68], [69], since only 30%

of patients with HCC and 20% patients with hepatic colorectal metastases (CRM) are qualified for surgical resection [70].

Percutaneous RFA system comprises a generator and one or multiple electrodes: a percutaneous metal rod which is directly placed into the target tumor tissue under the ultrasound (US) or other imaging guidance, and the grounding pad that is positioned on the patient’s skin [71]. RF generator, operating typically below 1 MHz, generates electric field between two electrodes. The mechanism of tissue heating with RFA is based on resistive energy loss due to impedance of the biological tissue [33]. The RF generator acts as a voltage source, so the average power delivered can be calculated from Ohm’s law: P = V²/Z, where Z is the impedance of the tissue volume, adjacent to the electrode [18]. For example, liver is relatively conductive because of its high water and ion content, so it creates a low-impedance electrical current path. Conversely, aerated lung and fat have lower water and ion contents, so are associated with much higher electrical impedance. This

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makes RF ablation challenging in lung since even electrically conductive tumors are surrounded by lung parenchyma [71].

The resistive heating, described by Ohm’s law, leads to temperature increase and the following coagulative necrosis in the area surrounding the RF applicator tip. The heat then spreads outwards into adjacent tissue by conduction. In order to prevent the risk of tumor recurrence, it is necessary to produce the ablation lesion that extends 0.5 to 1.0 cm outside the tumor margin. Temperatures above 60 °C induce lethal cell injury, however, temperatures higher than 100 °C are less effective, due to tissue burns and vaporization of watery components, which leads to increase in the tissue impedance and therefore limits further electrical conduction through the remaining tissue [72]. Additionally, cytotoxic temperatures are difficult to maintain if the ablated tumor is near large blood vessels. This heat-sink effect is a commonly described limitation of RFA and occurs when heat that is absorbed by flowing blood or air is carried away from the area of ablation, thereby dissipating the hyperthermia and decreasing RFA efficacy; because of this, tumor tissue that is adjacent to vasculature is less susceptible to thermal damage.

Overall, RF ablation has found the greatest utility in treating small tumors (up to 3 cm diameter) in the liver and kidney. The advantage of RFA technique is the possibility to use deployable devices (umbrella or star-shaped array of deployed tines) or multiple-electrode systems to improve the efficacy of RF ablation and increase the ablated volume for medium tumors (up to 5 cm diameter). Such commercially available devices are depicted in Figure 2.2 [73].

Figure 2.2 (A) 14-Gauge multitined StarBurst Talon RFA device (photo courtesy of Angio Dynamics). (B) 14-Gauge LeVeen needle showing 12 retractable electrodes

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(photo courtesy of Boston Scientific). (C) 17-Gauge Cool-tip RF electrode (photo courtesy of Covidien).

2.1.3 Microwave Ablation

Microwave ablation (MWA) is a relatively newer modality that was developed in Japan [74], [75], and a common MWA system consists of a microwave generator, flexible coaxial cable, and microwave antenna. Similar to RFA, MWA relies on the use of electromagnetic energy to generate heat and to induce cell death through direct hyperthermic injury. Electromagnetic waves, usually in the range of 900–2500 MHz are delivered to the target tumor region through a percutaneous antenna. Electromagnetic (EM) field causes polar water molecules, which have weak unequal dipoles, to orient in alignment with the field [76]. High frequency oscillations of EM field induces rotative motions of water dipoles, which results in increase of the kinetic energy. Higher kinetic energy of water molecules implies temperature rise, and when sufficient temperature values are achieved, - the target tissue undergoes coagulative necrosis.

Unlike RFA, heat propagation mechanism during MWA does not rely on conduction or impedance of the tissue, which makes this technique suitable for treatment of tissues with higher impedance, like lungs and bone, and generally, more attractive than LA and RFA. Other advantages over RFA, include the ability to achieve better heating of larger tumor volumes and a lower susceptibility to heat-sink effects because higher temperatures are achieved within a shorter duration [77]. During RFA, the zone of active heating is limited to a few millimeters around the active electrode, and the remainder of the treated tissue is heated by thermal conduction. By contrast, MWA at certain frequencies can heat tissue up to 2 cm away from the antenna [18].

Multiple antennas can be connected simultaneously making possible the treatment of bigger tumors or several small tumors within the same course of ablation [77]. Among disadvantages are the system itself which is more technically complicated, and larger size of MWA cables. Moreover, the antenna is prone to overheating, so the system will require a cooling mechanism.

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Modern technology allows integration of mechanisms of control into the ablation system: thermal control, field control, and wavelength control, that serve the goal of achieving a more reproducible and controlled ablation [78], [79]. Reports show utilizing MWA for the treatment of liver and renal tumor [80], and some data are also available about lung and pancreatic cancer [81], [82].

2.1.4 Laser ablation

Lasers play an important role in various medical applications related to ophthalmology, dentistry and skin treatment. The devices and systems used for laser tumor ablation are similar to those used for other clinical applications.

The earliest work related to laser thermal ablation has been performed back in 1980s and was dedicated to the treatment of hepatic tumor [83]. Since then, treatment of cancer with laser has become a competitive modality applied to many types of tumors, however laser ablation is not so well studied and introduced into clinical practice like RFA or MWA [84].

During laser ablation temperature elevation and the subsequent thermal injury of the tissue is described by the mechanisms of laser-tissue interactions.

Laser light interacts with various tissue components depending on the light wavelength, but most laser ablation devices operate at 800–1100 nm range to ensure better energy penetration. Such a wavelength range is explained by the presence of tissue specific chromophores inside the living cells, these chromophores absorb laser light predominantly at characteristic wavelengths 800 to 1100 nm.

A typical laser ablation system comprises a power source, laser medium, and a laser delivery fiber. In modern techniques laser delivery is performed trough a thin optical fiber and can be realized percutaneously or laparoscopically [23].

An optical fiber is introduced into a target tumor region through a needle or catheter, most commonly 13 to 15 gauge for modern quartz fibers. Diode laser (laser wavelength 980 nm) and Nd:YAG laser (working at 1064 nm) are commonly used for the hyperthermia application [85]. Laser light is energetic and is absorbed rapidly by the tissue near the applicator. This causes a rapid temperature elevation, and further heat propagates via scattering of the light and thermal conduction into adjacent tissues. The sufficient amount of energy

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will induce lethal thermal injury. The amount of thermal damage during such therapy is defined by the optical properties of the target tissue [86], as well as the laser parameters, such as type of fiber applicator, laser wavelength and power.

As in case of RFA, carbonization limits the maximal zone of ablation, since the ablation lesion produced by a single applicator are not larger than 2 cm in diameter. The system utilizing multiple applicators or applicator with diffusing tip may help improve the outcome of LA [83]. External cooling may be used if the applied power is sufficient to cause applicator heating.

A significant advantage of applying laser fiber applicators is MRI compatibility, allowing for preprocedural planning and intraprocedural treatment monitoring using a variety of temperature-sensitive techniques. Clinical application of laser ablation includes treatment of hepatic tumor [23], breast carcinoma [87], benign prostatic hyperplasia, and cancer of the pancreas. In addition, modern LA systems can be incorporated with MRI and US tools to improve the safety and efficacy of the procedure through tumor localization, targeting and monitoring of the ablation [88].

2.1.5 High intensity focused ultrasound

During HIFU, through tissue as a high-energy acoustic wave [12]. HIFU transducers operate typically in the range 0.2 MHz to 3.5 MHz and deliver ultrasound energy with power densities in the range of 100–10000 W·cm^-2 to the focal region. Such short but high intensity pulses induce temperature increase up to 60–95 °C within a few seconds without damaging adjacent tissues [12], [89]. The predominant mechanism of coagulative necrosis is thermal, i.e. conversion of acoustic energy into heat. In addition to thermal mechanism, cell death can be achieved through non-thermal effects, known as (boiling) histotripsy. For histotripsy, very short (micro- or millisecond long) acoustic pulses of high intensity (>5 times as high compared to thermal ablation) force bubbles in the tissue to oscillate and, subsequently, burst, causing mechanical damage to tissues at a subcellular level [89]. Currently, HIFU is used for the ablation of tumors in the liver, prostate, breast, and kidney, and benign thyroid nodules [2]. Figure 2.3 illustrates a working principle of this technique.

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Figure 2.3 HIFU treatment of intraabdominal tumor. Adopted from [2].

The possibility to selectively damage tumor without the need for skin incision, as well as no bleeding and no radiation make HIFU an attractive option for patients with breast carcinomas [90]–[92]. Multiple clinical studies were successfully performed on HIFU treatment of breast cancer. Wu et al. showed complete necrosis of breast tumors of 2-4.8 cm in diameter by HIFU technique [91], [93]. Authors demonstrated a 95% five-year disease-free survival, and the cosmetic results were estimated as excellent by most of the patients. Li et al. reviewed 11 arms of breast cancer treatment using HIFU therapy during the period 2002-2010. The complete necrosis rates achieved with MRI-guided HIFU were 59%, while US-guided HIFU showed 96% of complete necrosis [92].

2.1.6 Cryoablation

In contrast to high temperature thermal ablation techniques, cryoablation, or cryotherapy, provides the destruction of a pathological tissues by freezing.

Cooling down the tissue to −40°C will induce osmotic shock to the cells, and destroy cellular metabolism [67].

It has been used for cancers of the retina, skin, prostate, kidney, liver, breast, lung and bone [94]. This technique utilizes liquefied cryogenic gases (e.g., argon) injected inside the cryoprobe. There is a small chamber at the end of

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the cryoprobe; under low pressure inside the chamber gas undergoes adiabatic expansion, resulting in a temperature drop to as low as –160°C and the formation of an iceball around the probe tip. Cell death is associated with cellular dehydration, as freezing of extracellular medium causes an osmotic gradient, cell shrinkage and distortion of the plasma membrane [95]. During the thaw, the intracellular compartment becomes hypertonic, and fluid shift causes the cell to burst [95]. Cell mortality occurs –40°C and –20°C, and studies show that this temperature needs to persist to 10 mm beyond the tumor boundary [67].

2.2 Thermometry for monitoring thermal ablation procedures

2.2.1 Introduction

In previous chapter I presented the overview of thermal ablation (TA) procedures, their working principles and limitations. In order to perform the TA procedure in a most efficient way, a combination of several factors have to be achieved. First is related to accurate positioning of the applicator inside the tumor, that is achieved through localization of the tumor and identification of its contours [33]. Next, the clinical settings of the device, such as power, duration, the amount of energy, should be adjusted for each patient not only prior to the ablation but throughout the course as well. Such a control is impossible without real time monitoring of the temperature dose within the target region. However, the current ablation technologies lack accurate real time monitoring, and this issue impedes a full clinical integration of thermal therapies [17]. This limitation is caused by the fact that precise temperature data is missing during ablation therapy. Temperature increase is a function of the tissue properties, such as thermal and electrical conductivity, absorption coefficients, and blood flow, which vary from patient to patient and even within a single organ [34]. The effects of blood flow are more pronounced over the longer duration of ablation because of the heat sink effect and the local variations in perfusion [35]. Thus, there is a high demand in accurate temperature profiling and real time thermal mapping as this will serve the goal of overcoming technological limitations of hyperthermal procedures through performing several tasks:

A thesis submitted in partial fulfilment of the requirement of Nazarbayev University for the degree of

FIBER OPTIC SENSORS FOR IN SITU MONITORING DURING THERMAL ABLATION OF TUMORS

Madina Jelbuldina