## 4 Discussion  
Temperature-time responses of samples vary significantly within the same application group as depicted in Figures 2 and 3. These results confirm our initial assumptions regarding the non-uniformity of optical and thermal characteristics among subjects and show that temperature control is necessary for photothermal medical laser applications. As illustrated in Figure 2, our system can successfully control the tissue surface temperature during laser exposure. Mean tissue surface temperatures during the 5s temperature controlled period do not deviate from the target value more than 0.6 \({}^{\circ}\)C on average. The maximum and minimum temperatures observed during this period also remain within a narrow range (4.1 \({}^{\circ}\)C on average) demonstrating the temperature stability our system can provide. This finding is also supported by the small magnitudes of the standard deviation values in Table 2.  
Among a total of 144 samples in the temperature-controlled applications, only two samples of \(P_{\mathrm{laser}}=10.2\) W & \(T_{\mathrm{target}}=75\) \({}^{\circ}\)C group displayed anomalies. One sample contained fascia tissue underneath the target area, and the temperature increase caused it to contract, moving the laser irradiated surface in and out of the measurement spot. The fact that our system is a confocal _but_ not a coaxial optical setup caused measurement errors due to the alteration of the target surface distance to the probe. Off-axis alignment requires a precise and constant application distance. Use of a mechanical spacer or another technology (e.g., a distance sensor) to ensure the required target to probe distance is necessary for handheld operation. The other sample experienced a high rate of temperature change (>20 \({}^{\circ}\)C/s) and has undergone partial carbonization at the center of the laser spot, even though the highest measured temperature was below 90 \({}^{\circ}\)C. Considering the Gaussian laser beam generates higher temperatures at the center of the spot, partial carbonization at that location is natural. Figure 4: Box and whiskers plot of the tissue surface temperature reached at the end of exposure in the fixed-dose application. The red horizontal lines are the group medians. The blue boxes mark the interquartile ranges. The whiskers indicate the whole data range. On the other hand, fixed-dose application results display increasing tissue surface temperatures until the end of the procedure (cessation of irradiation), as shown in Figure 3, resulting in a wide range of final temperatures within each group, as presented in Figure 4. These results prove that a fixed-dose cannot be relied upon to ensure treatment success (reaching desired temperature) while preventing excessive thermal damage in the clinical conditions due to the high temperature sensitivity of physiological effects generated in PTT. Arrhenius temperature integral suggests that the progression of thermal damage is linear for constant tissue temperatures and exponential for linearly increasing tissue temperatures [6]. As a consequence, swiftly reaching a target value (depending on the desired physiological effect) and maintaining the tissue temperature at that value renders the progression of thermal damage a linear function of time. Our system can be used to constrain tissue surface temperature during laser exposure to control progression of thermal damage while guaranteeing the induction of desired physiological effects and preventing undesirable outcomes (e.g., carbonization, heat fixation, coagulative necrosis). As the tissue is exposed to laser light, its temperature increases and a multitude of phenomena occur simultaneously [29]. Coagulation and denaturation of proteins change the optical properties of the tissue [6; 19; 30]. Moreover, the absorption coefficient of water at the ytterbium fiber laser wavelength (\(\lambda_{center}\approx 1070\,\mathrm{nm}\)) decreases with increasing temperature [31].