Temperature controlled laser hyperthermia system using a newly developed laparoscopic system equipped with an ultra-compact thermal imager.

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Laser hyperthermia is one of the methods of treatment of malignant tumors. We have developed a thermal imaging endoscope with an ultra-compact thermal sensor and installed a new laparoscopic laser hyperthermia system to heat cancerous tissue to the appropriate temperature, focusing on the ability of a thermal imaging camera to perform 2D mapping of temperature facts. Hepatocellular carcinoma cells (N1S1) were implanted into the liver of Sprague-Dawley rats (n = 13) to obtain hepatocellular carcinoma in situ. Six rats underwent laparoscopic laser hyperthermia (70°C, 5 min) using a newly developed system, the rest of the rats received only laparoscopic injection. Lesion volume measurements and histological evaluations were performed in all rats. Laparoscopic laser hyperthermia systems provide stable temperature control. At a given temperature of 70°C, the temperature of the cancer target was maintained in the range of 68-72°C for 93.2% of the irradiation time (5 min). The median volume of heat-treated tumors was significantly smaller than that of untreated tumors. A newly developed laparoscopic laser hyperthermia system is capable of maintaining tumor surface temperature at any desired level and has been shown to be effective in the treatment of a rat model of hepatocellular carcinoma.
Hyperthermia is a very effective cancer treatment because cancer cells are easily affected by heat, and hyperthermia has been studied for a long time because of its minimal side effects1,2,3.
In recent years, attention has been drawn to laser hyperthermia (LHT), a method of heating tumor tissue with laser radiation. Laser heating occurs when light energy is absorbed by tissue and then converted into heat. The absorption of laser radiation by tissues depends on the composition of the tissue (proportion of extracellular matrix, collagen, water, etc.), and each organ has its own characteristics5. However, when translated into localized tissue heating, the thermodynamic effect on the tissue remains the same. The therapeutic effect of laser hyperthermia is due to tissue destruction due to water evaporation in tissues and apoptosis or necrosis of tumor cells6. Since LTT can be applied to organs in vivo using optical fibers, LTT can be used not only for cancers of luminal organs such as the esophagus 7 , but also for solid organ cancers such as liver cancer 8 , brain tumor 9 and renal cell anemia. carcinoma 10.
To achieve safe and effective LTT, it is necessary to control the temperature of the cancer tissue during heating and maintain it at an appropriate level. Magnetic resonance imaging (MRI) based temperature monitoring has been used for interstitial LTT of brain tumors, and temperature control has been shown to be effective in therapy9,11,12. On the other hand, temperature monitoring based on radiant energy (infrared) detection has the following advantages: (1) non-invasiveness and (2) real-time acquisition of the object’s surface temperature. In addition, (3) one can obtain a two-dimensional thermal distribution. Based on these advantages, we have developed a thermal imager temperature control method and demonstrated its usefulness in laser hyperthermia. In particular, we have successfully developed a feedback system that uses the temperature information received from the thermal imager as an input to automatically control the laser power when the target tissue is heated. Using this system, we demonstrate that the temperature of the target tumor can be maintained at a stable level in animal models13, which we report results in beneficial therapeutic effects14.
On the other hand, in recent years, laparoscopic surgery for intraperitoneal malignancies has been widely used as a minimally invasive cancer treatment. In the laparoscopic technique, the surgeon uses carbon dioxide injection to inflate the abdominal cavity, so that cancers deep in the pelvic cavity and subphrenic cavity can be well observed and treated. Thus, in addition to cancers of the stomach, colon, and rectum, laparoscopic surgery is currently being used to treat a variety of cancers such as gallbladder and hepatocellular carcinoma15,16. Therefore, we proposed the idea of ​​introducing hyperthermia as an additional treatment in laparoscopic surgery. Hyperthermia can be applied to tumors that are difficult to remove surgically (eg, tumors with indistinct borders or tumors involving large blood vessels), and thus can complement surgical approaches.
For the application of LTT in laparoscopic surgery, we have developed a laparoscopic system equipped with a compact thermopile transducer17. In addition to the thermopile sensor, the laparoscopic system has a laser clamp hole and a rigid endoscope. The system can simultaneously acquire images of the observation site and a two-dimensional map of the surface temperature, and can also maintain constant heating of the target tissue at a given temperature. In this study, in order to demonstrate the usefulness of this laparoscopic treatment system, we performed non-contact LTT under laparoscopic conditions in a rat orthotopic hepatocellular carcinoma model and confirmed its therapeutic effect.
A rat model of hepatocellular carcinoma was laparoscopically treated with TC-LTT for 300 s at 70°C. The reason for setting the heating time to 300 seconds is based on the results of a previous study in which radiofrequency ablation was used to treat hepatocellular carcinoma18. The heating temperature setting (70°C) was determined based on the results of a preliminary study of the relationship between heating temperature and processing depth (Supplementary Figure S1). Since the tumor thickness in model animals is approximately 6 mm, 70°C was chosen as the lowest temperature required to treat this thickness. This temperature setting is also designed to minimize damage to normal tissues.
Figure 1 shows intraperitoneal images of a model of rat orthotopic hepatocellular carcinoma observed by laparoscopy (AIM1588, Stryker, San Jose, CA, USA) before and after treatment. Hepatocellular carcinoma before hyperthermia (Figure 1A) was identified as a white nodular lesion in the left lateral lobe that regressed after hyperthermia (Figure 1B). As shown in Figure 1, the surface of the tumor sometimes turned black after laser irradiation. Since the temperature was maintained at 70 °C, this is unlikely to be carbonization, and the coloration may be due to the formation of methemoglobin by heating hemoglobin 19 (Section 4 of the Appendix).
Additional video showing the actual treatment in progress. Thermal imaging (left part of the video) shows that after the start of laser irradiation, the irradiation area heats up and the temperature rises. The brightfield image (on the right of the video) shows that the laser target continues to overlap the tumor site as the surgeon fine-tunes the target. At the very end of the laser exposure, the reflected laser radiation at the site of the tumor disappeared in the brightfield imaging, and the color of the dot representing the highest temperature in the thermal image changed from red to green.
On fig. 2 shows changes in tumor temperature and laser power values ​​during heat treatment. The laser power reached its maximum within 1 s after the start of laser exposure, and the tumor temperature reached the set temperature (70 °C) after irradiation with maximum power for approximately 30 s. The maximum laser power of the device is 3 W/cm2. It was confirmed that the laser power was automatically controlled to maintain the tumor surface temperature at 70°C for the next 300 seconds. After the tumor temperature reached 70 °C, the median tumor temperature during temperature control was 69.8 °C (min of 67.8, max of 77.4 °C), with a distribution of temperature variation of < 68 °C: 0.2%, 68–72 °C: 93.2%, and > 72 °C: 6.6%. After the tumor temperature reached 70 °C, the median tumor temperature during temperature control was 69.8 °C (min of 67.8, max of 77.4 °C), with a distribution of temperature variation of < 68 °C: 0.2%, 68–72 °C: 93.2%, and > 72 °C: 6.6%. После того, как температура опухоли достигла 70 °C, средняя температура опухоли при контроле температуры составила 69,8 °C (минимум 67,8, максимум 77,4 °C), с распределением колебаний температуры < 68 °C: 0,2%, 68–72 °C: 93,2% и > 72 °C: 6,6%. After the tumor temperature reached 70°C, the mean tumor temperature at temperature control was 69.8°C (minimum 67.8, maximum 77.4°C), После того, как температура опухоли достигла 70°С, медиана температуры опухоли при контроле температуры составила 69,8°С (минимум 67,8, максимум 77,4°С), распределение изменения температуры <68°С: 0,2%, 68-72°С: 93,2%, > 72 °С: 6,6%. After tumor temperature reached 70°C, median tumor temperature on temperature control was 69.8°C (minimum 67.8, maximum 77.4°C), distribution of temperature change <68°C: 0.2%, 68-72°C: 93.2%, > 72°C: 6.6%. Samples stained with hematoxylin and eosin (HE) and samples stained with terminal deoxynucleotidyltransferase (TdT) dUTP (TUNEL) are shown in Figure 3. In the treatment group, necrotic degeneration was observed throughout the tumor area, and normal liver bordering the tumor margin was also thermally degenerated to a thickness of approximately 1.5 mm (average thickness 1.4 mm (minimum 0.6, maximum 2.6 mm)).
Tumor surface temperature (blue dots) and laser power (orange dots) versus time during laser hyperthermia. The temperature of the tumor surface increased with the start of laser irradiation, and once the temperature reached the set temperature (70°C), the laser power was automatically adjusted to maintain the tumor surface temperature at 70°C.
(A), (B) and (C), macroscopic images (red triangles) of thermally treated tumors. Each image refers to one of three different animals. The direction of the laser beam is indicated by the orange arrow. (D) and (E), macroscopic images of untreated tumors (red triangles). Each image refers to one of two different animals. Photo (F) is an enlarged view of the red box in photo (C), which is located at the border between normal liver tissue and tumor tissue. Photograph (H) shows necrotic changes in the tumor (TUNEL positive). Photo (G) is an enlarged view of the red rectangle in photo (D), which is located at the border between normal liver tissue and tumor tissue. Each tissue sample was cut to maximize the area in the sagittal plane of the tumor. HE, scale bar = 5 mm (A, B, C, D), 0.25 mm (F, G), TUNEL, scale bar = 0.25 mm (H, I).
Tumor volumes at the time of sacrifice in the experimental and control groups are shown in Fig.4. The median tumor volume was significantly smaller in the treatment group (treatment group: 1.0 × 102 mm3, control group: 9.4 × 102 mm3, P = 0.0043). Histopathological results showed that in the treated group, necrosis of the entire tumor area occurred, and tumor growth was almost completely suppressed.
Scatterplot of individual tumor volumes in heat treated and control groups. A significant decrease in tumor volume was observed in the heat treatment group (P = 0.0043).
No sick or dead mice were found throughout the study, and there were no treatment-related deaths in the treatment group. In addition, localized abscesses and hematomas found in previous studies have not been observed20,21.
In this study, we successfully destroyed cancerous tissue with non-contact LTT in an animal model of orthotopic tumor using a newly developed temperature controlled laparoscopic laser hyperthermia (TC-LTT) system. Continuous monitoring using temperature sensors without time delay allows real-time display of two-dimensional temperature distribution in the illuminated area. Surgeons can know in real time whether the tumor is heating up too much or not enough. In addition, the laser power feedback mechanism through temperature control allows precise control of the target lesion temperature during the heating process.
To achieve a good therapeutic effect in the hyperthermia of malignant tumors, it is important to heat the tissue and maintain it at an appropriate temperature. Our preliminary experiments show that both too low and too high processing temperatures lead to insufficient results (Supplementary Fig. S1). Overheating or overheating has been reported to cause unwanted thermal effects including evaporation, charring (Supplementary Fig. S3), and applicator damage or malfunction.
Temperature monitoring during laser interstitial hyperthermia (LITT) for malignant tumors has been reported using thermocouples, MRI, computed tomography (CT) and other temperature measurement methods 22 . The use of thermocouples allows temperature measurement without time delay, but requires the insertion of a thermocouple into the tissue, which carries the risk of bleeding and tumor seeding. On the other hand, temperature measurement with CT or MRI is attractive in that it is non-invasive and can measure the temperature distribution in three dimensions (temperature resolution: ±0.2°C). However, the MRI system has a time lag of 4-5 seconds before the measurement and cannot track the change in temperature in seconds. In addition, MRI is difficult to adapt to non-fixed organs due to the noise caused by body movement. CT is a problem of the impact of ionizing radiation on biological tissues. On the other hand, the biggest advantage of the TC-LTT system is that the temperature distribution can be obtained almost in real time (with a time delay of only 0.12 s) and in a two-dimensional non-invasive way without the use of ionizing radiation. radiation. In addition to managing heat with thermal sensors, the system also allows surgeons to see processes associated with thermal changes in treated tissue in brightfield images, allowing surgeons to treat with confidence.
Because the system is a non-contact (non-puncture) form of temperature measurement and laser irradiation of the target tissue, there is no mechanical penetration into the tumor. For LTT of solid organs, interstitial irradiation is usually used, in which laser fibers are injected into the tumor and heated. However, tumor puncture is associated with the risk of bleeding and tumor seeding associated with puncture. In addition, puncture-type light-emitting devices such as the NeuroBlate optical laser probe (Monteris Medical, Minnesota, USA) often require a cooling system to prevent overheating of the probe tip, which not only complicates operation but also risks injury due to breakage. On the other hand, our TC-LTT system uses bare fiber without punctures, which eliminates the above risks.
LTT in patients with early stage hepatocellular carcinoma has fewer complications and is not inferior to surgery in the short term.
When the tumor tissue is heated to temperatures from 50°C to 100°C, coagulation necrosis occurs. However, heating above 100°C is fraught with rupture of the tumor, carbonization and incomplete coagulative necrosis due to evaporation of water in the tissue. Therefore, in order to completely cure the tumor, it is necessary to provide thermal energy from 50 to 100 °C to the entire area of ​​the tumor. Temperature control is essential to avoid side effects. Even at the laser power (3 W/cm2) set in this study, the tumor surface temperature exceeded 100°C without temperature control, resulting in carbonization (Supplementary Fig. S3).
If a high power laser is used, the temperature of the irradiated area may be higher than the set temperature before the temperature control mechanism is activated. However, in the setup used in this study (maximum laser power: 3 W/cm2), no such event was observed, and the temperature rise during the sampling time (0.12 s) was only about 0.1 °C. If a high power laser is used in the future, the temperature rises above the set temperature during sampling, which can be avoided by (1) lowering the maximum laser power setting and (2) reducing the sampling time.
In this study, the treatment time of 300 s was based on the results of a previous study in which radiofrequency ablation was used to treat hepatocellular carcinoma, but another experiment showed that tumor growth inhibition could be achieved even in a shorter time (Suppl). , Figure S2). In an additional experiment (Suppl. Sec. 2), a heating time of 150 seconds resulted in a depth of tumor necrosis comparable to a heating time of 300 seconds. However, at shorter heating times (less than 75 seconds), the depth of tumor necrosis fluctuated (became unstable). Surprisingly, however, 37 seconds of heating produced a depth of tumor necrosis equivalent to 300 seconds of heating in some cases. Therefore, if there is a device capable of uniformly heating the entire tumor (for example, a heating device in which the distance between the fiber optic probe and the surface of the tumor does not change much when heated), this can be done in less than 150 seconds. process. heating).
The flow rate is estimated as follows. First, the diameter of the beam spot on the irradiated surface was measured relative to the distance from the position of the fiber tip to the irradiated surface (Supplementary Fig. S5). Because the distance between the tip of the endoscope and the surface of the tumor was estimated to be about 10 mm during intra-abdominal surgery in rats, the diameter of the beam spot on the tumor at that time was estimated to be about 10 mm. (area 0.79 cm2) as shown in the appendix. Figure S5. Thus, at a laser power of 3 W/cm2, the energy flux density during intraperitoneal manipulation can be estimated at 3.8 W/cm2.
The optical thickness of near infrared light in living tissue is about 5 mm26. However, a maximum treatment depth of up to 9.3 mm was obtained in this study (Supplementary Figure S1). This may be due to heat transfer from the heated tissue rather than direct heating by absorbing near infrared light. Thus, it has been found that if heating is maintained at a certain temperature, a processing depth greater than the optical thickness can be obtained.
Further miniaturization is actually possible by reducing the size of the thermopile array. In this study, a thermopile with a spatial resolution of 32 × 32 (Φ = 9 mm) was used, the outer diameter of the endoscope was 14 mm. A 5.3 mm thermocouple array (HTPA8 × 8d (spatial resolution 8 × 8 pixels), Heimansensor, Germany) is currently available so that the outer diameter of the endoscope tip can be reduced to approximately 9 mm. However, reducing the outer diameter of the endoscope tip to 1–5 mm remains difficult.
In this study, the depth distribution of temperature from the irradiated area to the antipodal area is unknown. However, observations of histopathological preparations showed that the average thermal depth from the point of irradiation to the antipodal point was 4.3 (minimum: 3.2, maximum: 4.7) mm, and the thermal energy reached the entire area of ​​the tumor in this tumor model. In addition, there was no inadvertent thermal injury to other organs due to less damage to normal liver tissue and no treatment-related deaths.
The maximum treatment depth obtained in this study was approximately 9 mm. However, given the depth of penetration of light with a wavelength of 808 nm into tissues, the therapeutic effect of tumors thicker than this depth will be insufficient. However, many light-absorbing nanoagents with high thermal conversion efficiency have been reported, and it is possible to enhance the therapeutic effect by combining these agents.
The distance between the tumor and the fiber apex varies, and the size of the spot is inconsistent due to the movement of the liver associated with the movement of diaphragmatic breathing and the movements of the surgeon manipulating the laparoscope. However, tumor targeting can be continued by changing the irradiation site based on observations of bright-field and thermal images. The main factor responsible for the change in spot size is the displacement of the tip position due to the movement of the organ and the shaking of the operator’s hand. It is desirable that the change in beam size be as small as possible, and a possible measure to achieve this is to fix the endoscope with the apparatus, and not with the operator’s hand. In the future, if an image-tracking based endoscope tip compensation system is built, dimensional changes associated with organ movement can be minimized.
In terms of future prospects, since LTT has been reported for other types of cancer, TC-LTT based on thermal endoscopy may be applicable to other types of cancer in the future. Because the thermoendoscope-based TC-LTT system can be applied to lesions in a non-contact manner, it may be a good indication for mucosal resection or endoscopic mucosal resection in intraepithelial lesions of the gastrointestinal tract or in lesions with a high risk of bleeding that are resistant to endoscopic treatment submucosal dissection28,29,30.
Finally, we constructed the TC-LTT laparoscopic system with a thermal endoscope equipped with an ultra-compact thermal sensor, an additional metal oxide semiconductor (CMOS) camera, a fiber optic channel, and an automatic laser control system. exit. Using this system, non-contact TC-LTT was performed laparoscopically in a rat model of orthotopic hepatocellular carcinoma, and the cancer was successfully eradicated. The results indicate that non-contact TC-LTT can be performed laparoscopically and may be an effective approach in the treatment of solid organ cancer.
The designed thermal endoscope consists of a rigid endoscope (maximum diameter 14 mm, length 288 mm) (serial number 11499, Shinko Koki, Japan), an ultra-compact infrared thermal imaging sensor (HTPA32×32d L2.1, Sea Mann sensor, Germany), and channel 17 for introduction of an optical fiber for laser irradiation (Fig. 5A).
(A) Aerial view of the laparoscopic thermal imager and its handpiece components (lower right). The handpiece assembly consists of a fiber optic channel, a rigid endoscope, an air nozzle, and a thermal sensor. (B) Temperature controlled laser hyperthermia system configuration. The system consists of a laparoscopic thermal endoscope (bottom right), a laser generator (top left), a control PC (top right) and a microcontroller (bottom left). (C) Light source and laparoscope inflation system.
The thermal sensor visualizes a 2D temperature distribution at a frame rate of 8.3 frames per second and a spatial resolution of 32×32 pixels (temperature range 20-80°C linearly corresponds to pixel values ​​0-255). When connected to a rigid endoscope, brightfield images were acquired with a CMOS SLR camera (EO-1312C, Edmund optics, Barrington, NJ, USA).
The laser beam was guided by an optical fiber (NA 0.22, Ceramoptec, Bonn, Germany) and launched from the endoscope tip through the endoscope channel. The appendix shows the change in the size of the beam spot at the irradiation site as the distance between the tip of the light guide and the irradiation site changes. Figure S5. The temperature information received by the infrared thermal imaging sensor is transmitted to the microcontroller. According to the temperature information, the corresponding laser output is calculated in order to maintain a constant temperature of the irradiation target.
The TC-LTT laparoscopic system was used with a laparoscopic insufflation device (PNEUMO SURE, Stryker, San Jose, CA, USA) equipped with a light source unit (L10000, Stryker, San Jose, CA, USA) (Figure 5C).
The temperature distribution of the region observed by the thermal imaging sensor and the bright-field image of the region observed by the CMOS camera were controlled separately. The temperature monitor displays the hottest pixel as a red or green dot: a red dot when the laser is on and a green dot when the laser is off. In addition, the 9×9 pixels surrounding the red/green pixels are automatically extracted, and the four vertices of the square formed by the 9×9 pixels are displayed as blue dots (video 1). In this case, the average temperature value of 81 pixels (9 × 9 pixels) was automatically calculated, and we defined the average temperature as “the temperature of the irradiated target.”
The surgeon confirmed the location of the tumor on the bright field monitor and advanced the fiber through the canal until the tip of the fiber appeared on the bright field monitor. Next, the tumor is irradiated contactlessly through an optical fiber. According to the “irradiation target temperature”, the target tumor is heated while maintaining the temperature by automatically calculating the appropriate laser power.
During laser irradiation, the position of the laparoscopic endoscope was adjusted manually to ensure that the laser irradiation site was slightly separated from the tumor. When the irradiated area is far from the tumor, the laser radiation stops.
TC-LTT was performed in a rat orthotopic hepatocellular carcinoma model (method of preparation will be described later). Experimental animals were randomly divided into two groups: experimental (n = 6) and control (n = 7). After induction of general anesthesia, a 15 mm trocar (VersaOne Optical Trocar 15 mm, COVIDIEN, Norwalk, CT, USA) was inserted through a 1.5 cm skin incision into the abdominal cavity. A thermal imaging laparoscopic camera was inserted through the trocar and carbon dioxide was injected (injection pressure 3 mm Hg). In the treatment group, laser irradiation was carried out at a temperature of 70°C for 300 seconds. According to the addition, the size of the beam spot in this study was taken to be approximately 10 mm. Figure S5. In previous experiments, we have confirmed that this heating regimen (70°C for 300 seconds) has a therapeutic effect on the entire tumor area (Supplementary Figure S1).
Rats were sacrificed one week after hyperthermia. Liver lobes were removed and fixed in 10% formaldehyde solution, followed by GE staining. Tumor size was measured with a digital caliper after tumor removal, and the estimated volume was calculated as follows: (length) × (width) × (height) × 1/6π.
Statistical analysis was performed using the Mann-Whitney U test. The statistical package JMP 14 (SAS Institute Inc., Cary, North Carolina, USA) was used. P < 0.05 was considered to be statistically significant. P < 0.05 was considered to be statistically significant. Р < 0,05 считалось статистически значимым. P < 0.05 was considered statistically significant. P < 0.05 被认为具有统计学意义。 P < 0.05 P <0,05 считался статистически значимым. P<0.05 was considered statistically significant.
Rat hepatocellular carcinoma cells N1-S1 (CRL-1604, ATCC, Manassas, Virginia, USA) were used. Medium was Dulbecco’s modified Eagle culture supplemented with 10% FBS, penicillin (100 U/ml) (Thermo Fisher, Waltham, MA, USA), streptomycin (100 µg/ml) (Thermo Fisher, Waltham, MA, USA), and amphotericin. B (0.25 µg/ml) (Sigma-Aldrich, St. Louis, MO, USA). Culture the cells in an incubator at 37 °C with 5% CO2 and 95% air.
Eight week old female SD rats (Japan SLC, Hamamatsu, Japan) were used in this study. Rats were kept 3-4 per cage at controlled temperature (23-25°C) and relative humidity (50%) and 12 hours of light (7:00-19:00). All animal procedures were performed in accordance with guidelines approved by the National Defense Medical College Animal Care and Use Committee (license number: 19009).
SD rats were intraperitoneally injected with a mixture of anesthetics: medetomidine (0.3 mg/kg) (Nippon Zenyaku Kogyo Co., Ltd., Japan), midazolam (4.0 mg/kg) (Sandz Corp., Japan) and butor forphanol (5 0 mg/kg) (Meiji Seika Pharmaceutical Co., Ltd., Japan). After a small laparotomy, the left lobe of the liver was removed from the body and using a 30 G needle, 20 μl of PBS-based cell suspension (3.5 × 104 cells/μl) was injected under the liver capsule. One week after transplantation of the cell suspension, it was used as a liver tumor model in rats.
The datasets analyzed in the current study are available from the respective authors upon reasonable request.

Post time: Nov-01-2022