Abstract and keywords
Abstract (English):
Hyperthermia, i.e. tissue heating to a temperature of 39-45°C, is considered to be a very promising technique to increase the sensitivity of tumor cells to ionizing radiation and chemical preparations. At the present time, there are numerous methods for producing hyperthermia with the optimum method dependent on the required volume, depth, and site of heating. This paper presents the results of preliminary theoretical and in vivo confirmation studies of the feasibility of intraoperative local hyperthermia via induction heating of ferromagnetic material within a tumor bed implant that fills a resected tumor cavity. The implant is made during the surgical removal of tumor by mechanically filling the tumor bed with a self-polymerizing silicone paste in which very fine electroconductive ferromagnetic particles are uniformly distributed. Therefore, the implant can accommodate unique characteristics of each patient’s tumor bed. For the laboratory experiments, a prototype induction heating system was used to produce an alternating magnetic field with a frequency of about 100 kHz and a maximum intensity up to 3 kA/m inside an induction coil of inner diameter 35 cm. Experiments were conducted to heat a 2.5 cm diameter spherical implant both in open air and inside the thigh of a living rabbit. The results in both cases are in good agreement with our theoretical estimations. It was established that the temperature gradient near the implant surface decreases with increasing implant size, and for typical size tumor bed implants produces effective hyperthermia to a distance of more than 5 mm from the implant surface. This result provides hope for a decrease in relapse after treatment of malignant tumors using our combination heat plus intraoperative high dose rate local radiotherapy approach. Moreover, the externally coupled implant heating can be combined with local chemotherapy by applying a self-resorbable polymer film containing antineoplastic agents to the surface of the implant.

Local intraoperative hyperthermia, personalized tumor bed implant, induction heating, hot source heating
INTRODUCTION Oncological diseases are one of the most common pathologies in modern society. Their treatment, especially in the case of advanced stages of malignant growth, often leads to loss of function of the affected organ as a result of its surgical removal. Sometimes the loss of physiological functions of the body may be associated with the loss of other important body functions. A vivid example of social disability is loss of speech following treatment of laryngeal cancer. For this reason, interest in the development of new organ-preserving methods of combined and complex treatment in oncology is increasing. As a rule, the organ- preserving effect is achieved by combining various types of radiotherapy and chemotherapy [1-4]. Sometimes these treatments are combined with surgery or microsurgery [5]. One of the promising methods of increasing the sensitivity of tumor cells to ionizing radiation and chemical preparations is hyperthermia, i.e. tissue heating to a temperature of 39-45°C [6]. The main mechanisms of the sensitizing effect of hyperthermia on tumor cells are the following: Slowing down the process of repair of sublethal damage caused by exposure to drugs and ionizing radiation [7-9]; An increase in the oxygen content in tumor cells in a state of hypoxia, and as a result, their radio/chemosensitization [10, 11]; Increased bioavailability of drug agents due to increased tumor permeability and transmembrane transport as a result of local hyperthermia, as well as the use of heat-sensitive polypeptides [12, 13]. Please cite this article in press as: Vasilchenko I.L., Osintsev A.M., Stauffer P.R., Loyko I.E., Pastushenko D.A., Zharkova O.V., Braginsky V.I., Rynk V.V., and Vasilchenko N.V. Preliminary tests of local hyperthermia based on inductively heated tumor bed implant. Science Evolution, 2017, vol. 2, no. 2, pp. 15-21. DOI: 10.21603/2500-1418-2017-2-2-15-21. Copyright © 2017, Vasilchenko et al. This is an open access article distributed under the terms of the Creative Commons Attribution 4.0 International License (http://, allowing third parties to copy and redistribute the material in any medium or format and to remix, transform, and build upon the material for any purpose, even commercially, provided the original work is properly cited and states its license. This article is published with open access at http:// 15 Hyperthermia can also be used as an independent method in the treatment of cancer, because as a result of its application, the following changes in tumor cells occur: Direct thermal destruction or thermoablation of cells as a result of their heating to a temperature above 50°C, for example, when a magnetic nanoparticle is introduced into a tumor and coupled to external magnetic field [14, 15]; Apoptosis, as a result of impaired cell differentiation due to increased tumor temperature, [16, 17]; Change in cellular metabolism as a result of tumor heating [18]; Stimulation of the antitumor immune response of the body [19, 20]. Three types of hyperthermia are distinguished by the volume of exposure: general, regional and local. In case of general hyperthermia, the whole body of the patient undergoes heating, while in the case of the regional hyperthermia - a separate region of the body is heated. Obviously during regional hyperthermia, not only tumor tissues are heated, but also healthy tissues and organs. Local hyperthermia is now considered the most promising way of hyperthermia and involves heating of only the tumor itself and immediately adjacent tissues. There are several basic methods for implementing local hyperthermia [21]. Ultrasonic hyperthermia is based on focusing ultrasound waves with a frequency of about 1 MHz at the location of the tumor and converting their energy into heat [22]. Since the length of such waves is about 1 mm, and the diffraction limit of focusing is limited to approximately half the wavelength of the radiation divided by the refractive index, an almost pinpoint positioning of the heating zone is possible. However, the depth of penetration of ultrasound into the body tissue is limited to a maximum of 7-8 cm at this frequency. In addition, the propagation of ultrasound in the body tissues strongly depends on their heterogeneity. For example, the presence of air or bones creates reflections and absorption of ultrasound energy and sharply reduces the accuracy of focusing. Electromagnetic hyperthermia is accomplished by directing electromagnetic waves in the radiofrequency range (about 10-200 MHz) or in the microwave range (usually 430 MHz, 915 MHz and 2.45 GHz) into the tumor region using a single waveguide antenna or a set of phase focused antennas [23, 27]. The penetrating power of electromagnetic waves in the tissues of the human body decreases noticeably with increasing frequency and for microwave radiation at a frequency of 2.45 GHz is no more than 1-2 cm, which guarantees only surface and near-surface local hyperthermia. Reducing the frequency increases the depth of penetration, but greatly reduces the ability to focus precisely. As a result, for a frequency of 100 MHz, the size of the heated zone is generally at least 12-20 cm across. Another method of local heating is to use interstitial implants heated by various methods [28]. Among them, in particular, include the use of arrays of ferromagnetic needles injected directly through the skin into the tumor and heated by induction in an alternating external magnetic field. The main problem of the latter method is that a sharp temperature gradient exists in tissue near the surface of small heat sources, resulting in significant overheating of tissue immediately adjacent to the surface of the heat source and poor heating of more distant tissue. It is possible that a more uniform temperature distribution can be achieved by immersing the tissue region in a variable external magnetic field after introduction of magnetic nanoparticles into the circulatory system of the patient and accumulation of nanoparticles in the tumor due to peculiarities of blood flow and extravasation in tumor tissue [29, 30]. Another way to heat nanoparticles accumulated in a tumor is to transform the energy of near-infrared optical radiation into heat [31]. However, there is no reliable data on the clinical use of nanoparticle heating methods yet. Perhaps this is due to the problems of creating a high concentration of particles in tumor, as well as their subsequent removal from the body. In any case, at concentrations of nanoparticles achievable with systemic delivery, the amplitude of magnetic field intensity at 100 kHz can be estimated to be no less than 10 kA/m. In our opinion, a method for intraoperative local hyperthermia based on induction heating of an intracavitary or interstitial implant has good prospects for clinical use [32]. The personalized implant, which allows for individual characteristics of a patient’s tumor bed, is made during the removal of the tumor by mechanically filling its bed with a self-polymerizing mass in which the fine electroconductive ferromagnetic particles are uniformly distributed. Heating occurs mainly due to eddy currents in the small ferromagnetic particles which are uniformly distributed throughout the volume of the implant. At a magnetic field frequency of about 100 kHz, only the implant itself and tissue immediately adjacent to the tumor bed are heated. This method is ideally combined with the method of high- dose intraoperative local radiotherapy [33]. In addition, it can be easily combined with chemotherapy by applying a film of a self-absorbable polymer material with the addition of antitumor drugs to the surface of the implant [34]. The purpose of this work is to test the possibility of conducting hyperthermia based on induction heating of the implant under conditions close to clinical, including heating of in vivo tissues. MATERIALS AND METHODS A laboratory prototype of a clinical installation for induction heating (Fig. 1) has been developed together with LLC Research and Production Complex “Magnit M” (Tomsk, Russia). The device is an induction heating system with a coil consisting of 5 turns of internal diameter 30 cm, operating at a frequency of about 100 kHz. The coil has water cooling and an electrostatic shield. Thermocouples made of thin (0.1 mm) wires are introduced into the coil perpendicular to the lines of magnetic induction and serve to measure temperatures in the heated region and provide feedback to control the system power level. Accuracy of the temperature measurements with each probe was ± 0.7°С. R2 w T (R )  T  0 , (2) 0 0 3 and the temperature gradient near the surface of the implant, as follows from (1) and (2): T(r) 0 0 0 R w T  T (R )   . (3) r r  R0 3 R0 Fig. 1. Equipment setup for induction heating: (1) control unit; (2) transformer and matching circuit unit; (3) inductor coil. Implants were made from self-polymerizing silicone paste as used in dentistry for making casts. Each implant consisted of 11 g of a polymer base into which 11 g of steel balls, each with 1 mm diameter, were homogenously integrated. The diameter of the overall spherical implants was 2.5 cm. One catheter was inserted into the center of the polymer sphere for subsequent insertion of thermocouple sensors to measure internal implant temperature. To conduct in vivo studies, identical implants were introduced into both thighs of an anesthetized live rabbit, under the supervision of a veterinarian. To clarify the position of the implants, a 3D reconstruction of the object was performed on the basis of computer X-ray tomography (Fig. 2). Work with laboratory animals was carried out in accordance with the protocol of research according to the Geneva Convention of 1985 "The International Guiding Principles for Biomedical Research Involving Animals" and the Helsinki Declaration of 2000 "On the We have carried out elementary estimates. To achieve a hyperthermic effect, the surface temperature of the applicator should be at least 43°C. Then at T0 = 37°С, taking the thermal conductivity of body tissues equal to the thermal conductivity of water  = 0.6 W/(mK) and the dimensions of the applicator R0 = 0.01 m, we obtain the specific thermal power released in the applicator for the estimation, w  105 W/m3 = 0.1 W/cm3. The temperature gradient, according to (3), is approximately 0.5°C/mm. For example, at an applicator temperature of 43°C, the boundary of hyperthermal heating (39°C) will pass at a distance of 8 mm from the surface of the applicator. This means that the tissues lying in the so-called risk zone, that is at a distance of ≤ 5 mm from the tumor, are guaranteed to be heated. The decrease in the temperature gradient near the surface of the implant with an increase in its size, determined by expression(3), is an important fact that allows one to get rid of the main drawback inherent in needle-like ferromagnetic implants, as noted above. At a mass fraction of 50%, the concentration of steel balls with a radius RB = 0.5 mm is about 400 cm-3. The heat energy released in one ball per unit time is [33]: 2 2 Humane Treatment of Animals". Ps  3 RB Hm 20  /  , (4) RESULTS AND DISCUSSION According to the estimates given in [35], the temperature distribution T(r) in tissues surrounding a spherically symmetrical implant with a radius of R0, in the absence of convective transfer or perfusion of blood in the stationary limit is determined by the following expression: 0 R3 w 1 where  - electrical conductivity of steel,  - magnetic permeability of steel, 0 - permeability of free space,  = 2f, f - frequency of magnetic field, Hm - Amplitude of magnetic field strength. Expression (4) makes it possible to estimate the intensity of the alternating magnetic field necessary to produce effective hyperthermia from the implant. For the above parameters, Hm  200 A/m. Stauffer et al. [33] present the evaluation of magnetic field for the case of basal T (r)  T0  3 , (1) r (0.7 kg / m3 / s) and active (6.2 kg / m3 / s) blood perfusion in tissue surrounding the tumor [37]. They were 224 A/m where  - thermal conductivity of tissues; w - volume power density of heat sources, а T0 - ambient temperature. The temperature near the surface of the applicator is, respectively: Fig. 2. Results of 3D-modeling of the object. and 414 A/m, respectively, which in any case is much less than the intensity of the alternating magnetic field required to heat a tumor containing magnetic nanoparticles accumulated from systemic delivery. Edge Center 25 mm 55.0 Probe temperature, °C 50.0 45.0 40.0 35.0 30.0 25.0 0 2 4 6 8 10 Heating time, min Center Edge (b) Fig. 3. Results of heating the implant in the open air: (a) location of thermocouples; (b) heating kinetics. Fig. 3 shows the results of heating the implant in the open air at a temperature of 25°C. Heating was carried out while maintaining a temperature of 50 ± 1°С at the center of the implant with a magnetic field strength in the coil of 1000 ± 50 A/m. Fig. 3b presents data averaged over three measurements. As can be seen, a rapid heat-up to the set temperature occurred within about 3 minutes, which agrees well with estimates made by us earlier [35]. It is obvious that low thermal conductivity and heat capacity of air, even with significant convection due to a large temperature difference, does not create a significant heat dissipation from the implant. This results in a sufficiently rapid heating and a steady-state output. Fig. 4 presents the laboratory results of an in vivo study of the temperature distribution around the 2.5 cm diameter implant embedded in the thigh of a rabbit. We should note that the junctions of thermocouples that measured the temperature of tissues in the zone of conditional risk, that is at a distance of ≤ 5 mm from the implant surface, were installed with an error of ± 1 mm. The figure shows the data averaged over the two implants. As can be seen, with the same inductor parameters, the time to reach the steady-state mode of implant heating increased approximately twofold. Obviously, this is due to the significantly higher heat losses into perfused rabbit thigh tissue. The surface temperature of the applicator was approximately 44.2 ± 0.7°C, and the temperature at the boundary of the conditional risk zone was approximately 41.7 ± 0.7°C. Thus, the average gradient near the surface of the implant was found to be approximately 0.5°C/mm, which agrees very well with the estimates given above. It is important to note that the junctions of thermocouples located at a large distance from the surface of the implant (17 ± 1 mm) showed, within the accuracy of the measurements, the same temperature equal to the body temperature of the rabbit. This fact confirms the absence of heating of the body tissues in an alternating magnetic field with a frequency of 100 kHz. CONCLUSIONS AND FUTURE DIRECTIONS Preliminary in vitro and in vivo studies confirm the feasibility of producing effective hyperthermia in tumor bed tissue around a resection cavity filled with an intraoperatively manufactured implant containing conductive ferromagnetic particles by applying an external sub-megahertz range magnetic field. The method involves removing tumor tissue and filling the tumor bed with a self-polymerizing paste containing conductive ferromagnetic particles. As a result of selective heating of the tumor bed to a temperature of 41-45°C, a significant reduction in possible relapses to radiotherapy treatment is expected. This method of local hyperthermia has low peripheral toxicity since both the radiation and thermal doses fall off rapidly with increasing distance from the implant. The local hyperthermia method is readily combined with radiotherapy by embedding brachytherapy catheters at appropriate positions within the implant at the time of formation within the tumor bed. Alternatively or in addition, radioactive seeds with a short half-life period could be embedded within the implant at the time of manufacture. Hyperthermia based on the intraoperatively created implant can be combined with chemotherapy by applying a self-absorbing polymer film containing antineoplastic or other drugs to the surface of the implant. In this case, the sessions of hyperthermia can enhance the bioavailability of the added drugs. It should be possible to improve control of heating by using a ferromagnetic substance with a Curie point near the desired implant surface temperature of 45°C. By relying on thermoregulation near the Curie point, there is no need for including a thermal sensor within the implant to provide feedback for varying the magnetic field strength to maintain constant temperature during the hyperthermia procedure. ACKNOWLEDGMENTS The authors are sincerely grateful to the rector of the Kemerovo State University, A. Yu. Prosekov, for assistance in organizing and conducting the research. (b) At risk Center Edge Body 5 mm 25 mm > 15 mm 52.0 50.0 Probe temperature, °C 48.0 46.0 44.0 42.0 40.0 38.0 36.0 34.0 0 2 4 6 8 10 Heating time, min. Center Edge At_risk Body (c) (d) Fig. 4. Results of the in vivo experiment: (a) implantation of an implant with catheters for temperature monitoring into a rabbit thigh; (b) placement of the rabbit in an inductor coil; (c) location of thermocouples; (d) heating kinetics.

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