Measurements of Tissue Temperatures During Semiconductor Laser Welding Process

Andrzej Zając, Henryk Polakowski,

Military University of Technology, Institute of Optoelectronics

Warszawa, Poland e-mail azajac@wat.edu.pl

Tadeusz Piatkowski, Mariusz Kastek Military University of Technology,

Institute of Optoelectronics Warszawa, Poland

e-mail mkastek@wat.edu.pl

Abstract— During the research the effect of coagulation of biological tissues was investigated, caused by pulse SSL and CW semiconductor laser illumination (λ ≈ 1000 nm). The state and temperature of soft biological tissues exposed to laser radiation was observed in real time, using thermographic and optical methods. The temperature changes, including maximal values, were evaluated by analyzing recorded thermal image. The measured scattered radiation of diagnostic laser (λ = 635 nm) was used to assess the tissue condition during the laser welding process. Exemplary results of the laboratory research on these methods and an interpretation of the results are presented.

Keywords-component; laser therapy, medical lasers, tissue welding, temperature measurements, thermography

I. INTRODUCTION The idea of the biological tissue laser welding is known for

a few years. However, there is still the lack (except the reference [1]) of reports confirming the implementation of this method in clinical practices. It is commonly known that the basic problem here is to achieve the precise control of physical processes occurring in tissues welded during an operation. Tissue welding relies on obtaining a local coagulation focus on both sides of the tissue incision. Two fundamental phenomena play a very important role here: the union of the tissue heated up and an increase in collagen fibers within a coagulated region. This process is very complicated and currently it is not fully described. Obtaining the correct thermal union and not big coagulated region at the same time requires meeting very critical conditions concerning the laser radiation power, the selection of the proper work regime of a device and, what is particularly important, the assurance of a very accurate temperature control of an irradiated tissue.

The possibility to apply different lasers for several medical procedures was considered and analyzed for many years [2, 5]. Currently pulsed Nd:YAG lasers are commonly used for tissue welding. However the development of high output semiconductor lasers (lased diodes, LD) with pulsed operation regime now also allows to apply such lasers for tissue welding. Available LD devices have power output of 100W in typical spectral range used in surgery (800-1000 nm). The temperature of a laser-treated tissue is a function of the power of a laser source, which is around 1 kW (peak pulse power) and several watts in CW mode for currently used solid state lasers. A 100 W semiconductor laser seems to have completely different characteristics than and it does not look like a direct replacement. Semiconductor laser offers, however, a very

attractive set of features, such as small size, simple and safe power supply, easy to obtain beam modulation, lower cost and longer exploitation times. As a result it can beneficial use such laser In medical procedures like tissue welding.

II. CHOICE OF LASER TYPE Analysis of laser parameters reveals that in CW mode both

laser types: Nd:YAG (currently used in medicine) and semiconductor lasers offer similar power output figures. The spectral ranges are also similar. The significant difference is in pulsed mode, which in case of LD devices is achieved by beam modulation, whereas Nd:YAG lasers used for tissue welding are inherently pulsed devices working on free generation and they offer pulse powers in the several kilowatts range. The comparison of basic parameters of both aforementioned laser types is presented in Table 1.

TABLE I. COMPARISON BETWEEN ND:YAG AND LD LASERS

Parametrs Laser Nd:YAG LD Laser max. P CW Up 100 W up 100 W max. Ei Pulse Up to 2 J Several mJ

Wavelength λ1 = 946 nm, λ2 = 1064 nm, λ3 = 1338 nm.

λ = 800 – 1000 nm

pulse duration ti = several ms Modulated beam Spectral bandwith Δλ = 1 nm Δλ = several nm Price 10 000 $ 3 000 $ Dimension 50 x 50 x 350 mm 20 x 20 x 10 mm Cooling power 1÷6 kW 100 W

III. MEASUREMENTS OF TISSUE TEMPERATURE The welding process realized correctly requires the control

of the tissue temperature in real time and, according to [3, 5], it should be realized with the accuracy of ± (1÷3)ºC around the coagulation temperature of the tissue protein. The absolute value of the coagulation temperature depends on the tissue type, the exposure time and equals from 65ºC to 95ºC (for human tissues) [4]. The welding process control is carried out in practical systems by measuring the tissue temperature during its welding. This measurement can be carried out by means of contact methods (offering low accuracy in case of the laser radiation exposition) or non-contact methods. The contact methods, for the sake of the limited accessibility to a relatively small operating region and the need of measuring probe insertion on a tissue surface near the place of a laser radiation interaction, are inconvenient for medical applications and cause additional probe sterilization issues These methods are mainly used in laboratory research on a laser radiation interaction with

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segments of biological tissues. The non-contact temperature measurement of a tissue can be done with the use of MRI technique [6], special pyrometers [7] or thermal imaging systems [6]. The assessment of the actual condition of an internal tissue can be also done by monitoring the spatial distribution of scattered laser radiation. Local change of light scattering properties of a tissue in the area of laser exposure is strongly related to the absorbed laser energy. The resulting optical signal which can be used to monitor the condition of a tissue can be observed in a wide angle α=± (5º÷85 º). In the presented work the additional laser was applied, operating in the visible spectral region.

IV. TISSUE WELDING PROCESS The main issue in the welding of biological tissues is their

non-homogenous nature and variability of their optical properties [3, 8]. Absorption of a laser radiation by a tissue causes local temperature increase. In each case the specific coagulation temperature has to be determined for every welded tissue, which may vary from 65°C to 95°C, depending on the kind of the tissue. A-priori assessment of the temperature of laser-illuminated tissue is currently not possible. As a result the temperature in the volume of a tissue has to be precisely controlled (within ± 2°C margin) to achieve proper coagulation, which results in uniform tissue connection and reduces the wound healing period. For a proper tissue welding, the following conditions must be assured:

• constant temperature monitoring (real-time operation required) of the tissue area exposed to laser radiation,

• determination of necessary (preferably minimal) amount of laser energy required to initiate coagulation,

• application of compounds enhancing the laser welding process (dyes like isocyanine, absorptive nanoparticles, absorptive layers) [9, 10].

The laser beam during welding is usually focused in such a way, that the focal point is in the volume of a tissue. In such a setup the spatial distribution of laser energy can be changed by altering the distance between lens and the tissue. For a fixed laser energy level the coagulation depends on the absorptive properties of a tissue.

Figure 1. Temperature distribution inside uniform tissues with different absorption

Calculated vertical temperature distribution inside the tissue for different values of absorption coefficient of an isotropic medium is shown in Fig.1. The conditions suitable for tissue welding are achieved for the following coefficient values: α1 = 0.2 cm-1 and α2 = 1 cm-1. The coagulation areas for both

aforementioned values is also presented in Fig.1. The tissue temperature of 80 °C ± 2°C is maintained in the coagulation area. For the α1 coefficient the uniform tissue bonding occurs at the depth between 4.2 mm and 8.1 mm, whereas for α2 coefficient the welding depth is between 5.6 mm and 9 mm. It should be mentioned that for the same beam properties the tissue temperature is much lower for α1 coefficient and as a result coagulation will not occur. It is caused by different penetration depths of laser radiation for different absorption coefficients. The analysis of temperature distribution were conducted for typical values of absorption coefficient, characteristic to biological tissues. However, in real case of tissue welding this coefficient is not known and because of that the tissue temperature control must be applied [11].

V. TEST STAND FOR TISSUE WELDING Block diagram of the laboratory stand used for the tissue

welding is presented in Fig.2.

Figure 2. Diagram of a test stand for LD laser tissue welding

The analysis of temperature distribution were conducted for typical values of absorption coefficient, characteristic to biological tissues. However, in real case of tissue welding this coefficient is not known and because of that the tissue temperature control must be applied [11]. Semiconductor laser operating in CW mode at a wavelength λ =635nm (power output P=5mW) was used as a diagnostic device. Welding laser was infrared (NIR) semiconductor laser because such radiation is strongly absorbed by biological tissues. In our system it was laser diode LIMO90-F200-DL980 operating at λ = 970 nm with power output up to 100 W in CW mode. The beam could be modulated with a frequency up to several hundred Hz. Laser diode was controlled by electronic circuit comprising of power supply and generator. This setup allows for the control of beam intensity both in CW (current) and pulse mode (duty cycle). Laser radiation was transferred to operation probe by an optical fiber. The tissue during tests was placed on X-Y table for a precise positioning. The temperature distribution during welding process was visualized by FLIR SC5600 thermal camera with dedicated software for the recording of image sequences. Simultaneously the scattered radiation of diagnostic laser was observed. For the applied wavelengths there are significant differences in absorption coefficients for different biological tissues, reaching up to three orders of magnitude [8].

VI. INVESTIGATION OF THE INTERACTIONS OF LASER BEAM WITH BOLOGICAL TISUES

The test stand described above was used to investigate the interactions between tissues and laser beam using CW and pulse operating modes, at different power outputs of 0.74 W, 3.74 W, 6.4 W and 9.2 W. Resulting thermal effects observed

in muscular tissue were recorded as a sequence of the thermal images. By analyzing of those images the temperature values were determined. The resulting temperature plot versus time obtained for 0,74 W power output (CW mode) is presented in Fig.3 and the plots for the other used power outputs are shown in Fig.4.

Figure 3. Temperature changes on tissue’s surface caused by laser beam (CW; 0.74W)

For the lowest power output the tissue is slowly heated and its temperature rises by 6 °C after about 60 seconds. Even such low power causes tissue coagulation which can be seen as a white ring-shaped area (Fig.3).

Figure 4. Changes of temperature on tissue’s surface caused by laser beam (CW; 3.7W; 6.4W; 9.2W)

Temperature rise rate is increased with the output power of a laser. For 3.7 W power output the coagulation temperature is achieved after 20 seconds, and after 40 seconds the tissue temperature reaches 138 °C. At this level the tissue starts to burn. Just before the burning point the temperature drop can be observed, which is caused by a rapid water vaporization in the illuminated area. After that the muscular tissue is being carbonized which is shown in Fig.5. Around carbonized area the bright ring of coagulated tissue can be seen, about 3.5 mm in diameter.

Figure 5. Sample thermal image of the illuminated area (CW; 3.7W), result of interaction

Two fold increase of the output power (to 6.4W) brings the coagulation temperature in 6 s (Fig.4). Further exposure causes tissue carbonization. At final value of power output (9,2 W) the coagulation temperature is reached in just 3 seconds and at

118 °C rapid temperature drop (by 30°C) occurs due to water vaporization. At this exposure parameters the diameter of the coagulation ring area is only 2.5 mm.

It can be seen from the presented results that for increased power output of a laser beam the time needed to achieve coagulation temperature decreases and the size of coagulation area decreases as well. In case of CW laser the tissue surface temperature increase is quasi-linear for every tested power level. The results may be helpful in determining the optimal beam properties for laser tissue welding.

Figure 6. Changes of temperature on tissue’s surface caused by a laser operating in pulse mode

During the next test phase the interaction of pulsed laser with biological tissue was investigated. First measurement was conducted for the following parameters: diode current: I = 15 A, pulse repetition f=1000 Hz, duty cycle ti= 0.5 ms and pulse energy up to Ei = 1.85 mJ. The tissue temperature versus illumination time is presented in Fig. 6. After the exposure time t = 30 s the temperature increase of 7 °C was observed. In case of pulsed laser operation the temperature rise curve is non-linear, contrary to quasi-linear relation observed when the CW laser was used. It means that the laser energy flows across the tissue volume around the point of interaction.

For increased diode current I = 20 A the pulse energy was Ei = 3.2 mJ for the same pulse frequency. The corresponding plot in Fig.6 shows the temperature increase by 13°C after 60 seconds of exposure. However the temperature increase after 30 seconds was 9 °C, merely 2 °C more than for the previously used pulse energy of 1.85 mJ. Fig.7. presents the temperature profiles on the surface of muscular tissue induced by a pulsed laser beam (0.5 ms duty cycle and repetition rate 1000 Hz).

Figure 7. Temperature profiles of muscular tissue as a function of total laser energy

During three exposures the total energy introduced to treated area was 18.5 J, 57.6 J and 78.2 J respectively. It can be seen that for total energy E = 18.5 J the area with elevated temperature is smallest, only 2.5 mm in size. The tissue temperature rise is ΔT = 7.5 °C. By increasing the energy

a b

transferred to the tissue to E = 57.6 J the corresponding temperature increase of ΔT = 14 °C was achieved and the affected area was 4 mm in diameter. Further increase in total energy value resulted only in slightly higher temperature level and a bit larger area size.

VII. MEASUREMENTS OF SCATTERED LASER RADIATION The result of laser coagulation of a muscular tissue is

presented in Fig.8a. The bright area in the middle of the sample is the coagulated tissue. The tissue welding process can be monitored by measuring the intensity changes of scattered radiation of a diagnostic laser pointed at coagulation area. The amplitude of such radiation measured at three angles of observation are presented in Fig.8.b.

Figure 8. Laser tissue coagulation: a) coagulated muscular tissue, b) amplitude of a scattered radiation of a diagnostic laser measured

at three different angles

The characteristic depicted in Fig. 8 illustrating the changes of the optical signal, obtained by means of three optical fibers. For the one coagulation point of the muscular tissue was split into four time intervals determined by the number of generated pulses. The first interval shows the change dynamics of the radiation scattered in the tissue with the lack of energetic radiation exposure. The second interval shows the initial exposure effect using the LD laser radiation. A strong increase in the amplitude value of the optical signal is seen here. The successive changes of the tissue parameters being the result of the thermal changes induced by the laser radiation have an impermanent character. The discontinuation of the exposure before reaching the maximum value of the signal amplitude in this range permits the withdrawal of changes in a tissue structure. This effect is impossible to observe visually. The third interval includes the preservation of changes in the tissue. The laser radiation absorption continues at a constant level. This time the tissue surface (as a result of the local temperature increase in the tissue volume caused by the laser radiation) becomes drained within the small depth causing the time decrease of the optical signal value. In this interval, the visible evaluation of the magnitude of the coagulation focus is possible. The amount of the energy absorbed in the tissue depends on the focus magnitude. The fourth interval presents a slow increase in the radiation absorption (an increase in the tissue temperature) allowing for instance carrying out the procedure of the tissue laser welding. Above this range the out of control phenomena in the tissue manifesting in its damage (including a carbonization) may occur.

VIII. CONCLUSIONS Analysis of the process of adiabatic tissue heating by means

of a solid state laser and experimental test results revealed that currently available semiconductor lasers can be successfully applied for laser tissue welding. Of two analyzed operating modes the constant wave operation yielded better results. Proper application of solid state lasers require, however, more complicated operating probes and strict temperature control because of high power densities on the surface of treated tissues. Both presented methods of process temperature control can be used. Even though thermal camera records and visualizes surface temperature only, it can be applied because temperature differences between surface and internal tissue layers are within acceptable margin. Initial test and research results obtained with the presented method are promising. Its simplicity and universal character allow for automation of the process of laser tissue coagulation.

ACKNOWLEDGMENT The result presented in paper are supported by realization

of the Project is co-financed by the European Regional Development Fund within the framework of the 2. priority axis of the Innovative Economy Operational Programme, 2007-2013, submeasure 2.1. "The development of centres with high research potential". Contract no. POIG.02.01.00-14-095/09

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