Analysis of Thermally Denatured Depth in Laser Vaporization for Benign Prostatic Hyperplasia using a Simulation of Light Propagation and Heat Transfer (secondary publication)

Abstract

Background and Aims: Laser vaporization of the prostate is expected as a less invasive treatment for benign prostatic hyperplasia (BPH), via the photothermal effect. In order to develop safer and more effective laser vaporization of the prostate, it is essential to set optimal irradiation parameters based on quantitative evaluation of temperature distribution and thermally denatured depth in prostate tissue.

Method: A simulation model was therefore devised with light propagation and heat transfer calculation, and the vaporized and thermally denatured depths were estimated by the simulation model.

Results: The results of the simulation were compared with those of an ex vivo experiment and clinical trial. Based on the accumulated data, the vaporized depth strongly depended on the distance between the optical fiber and the prostate tissue, and it was suggested that contact laser irradiation could vaporize the prostate tissue most effectively. Additionally, it was suggested by analyzing thermally denatured depth comprehensively that laser irradiation at the distance of 3 mm between the optical fiber and the prostate tissue was useful for hemostasis.

Conclusions: This study enabled quantitative and reproducible analysis of laser vaporization for BPH and will play a role in clarification of the safety and efficacy of this treatment.

Introduction

According to the Statistic and Information Department of the Minister's Secretariat at the Japanese Ministry of Health, Labour and Welfare, there are 418,000 men in Japan who suffer from benign prostatic hyperplasia (BPH) as of the year 2011 ¹⁾. BPH is a very common condition among middle-aged to elderly men. It is said that 50% of men at the age of 50 have hyperplastic changes of the prostate tissue, while 50% of these men manifest some type of symptoms ²⁾. Presently the mainstream treatment method for BPH is the transurethral resection of the prostate (TURP) ^{3, 4)}. TURP is a highly effective treatment method for BPH, performed widely across Japan and is considered to be the standard treatment method for BPH. However there are some drawbacks of TURP such as post-operative hemorrhage and hyponatremia (TUR syndrome) which require careful attention as this is a treatment of a condition which affect the elderly ^5–8). Presently, different approaches using specific laser devices have been investigated as an alternative for TURP and are showing promising results. Especially so, are the holmium laser enucleation of the prostate (HoLEP) which uses the Ho:YAG laser at the wavelength of 2.1 µm, and the photoselective vaporization of the prostate (PVP) which uses the lithium triborate (LBO) laser which emits laser energy at 532 nm, the same wavelength as the second harmonic of the 1064 nm Nd:YAG laser achieved via frequency doubling by a KTP crystal. These two laser-based approaches are representative treatments using lasers and have been approved by the Japanese regulatory agency for pharmaceutical affairs in 2011. HoLEP is being acknowledged as a safe and effective treatment method because of its decreased hemorrhage ^{9, 10)}. However, because the technical and handling skills associated with this procedure are highly clinician-dependent and demanding, the acquisition of such skills is considered to be difficult ¹¹⁾. PVP which utilizes laser vaporization is known for its high hemostasis resulting from the tissue coagulation surrounding the vaporized area. It has procedural similarities with TURP and is considered much easier to learn compared to HoLEP ^{12, 13)}. However, it has been reported that treatment efficacy is considered to be lower than that of TURP ¹⁴⁾. In recent years, besides the Ho:YAG and LBO lasers, prostate vaporization using a diode laser with the wavelength of 980 nm has shown excellent clinical results ¹⁵⁾. Laser energy of Ho:YAG laser wavelength of 2100 nm is highly absorbed by water and its penetration depth is thus limited to no more than approximately 400 µm ¹⁶⁾. Laser irradiation of water-rich prostatic tissue with this wavelength will result in vaporization and coagulation of only the tissue surface. Any thermal effect that could reach deeper tissue is limited and easily controlled, which makes this laser the most suitable for laser enucleation of the prostate. As for the LBO laser, its wavelength of 532 nm is strongly absorbed by hemoglobin. The penetration depth into blood is approximately 800 µm and the LBO laser is absorbed efficiently by blood in the vascular rich prostatic tissue ¹⁶⁾. During a PVP procedure, removal of prostatic tissue is achieved through vaporization and efficient removal of tissue mandates a high vaporization capability of the laser device being applied. The residual thermal damage seen in the coagulation layer has a hemostatic effect and is an important factor for the suppression of intra-operative hemorrhage. In a PVP procedure using the LBO laser, the high absorption of laser light by hemoglobin results in its high hemostatic effect. In contrast, laser energy at the wavelength of 980 nm is absorbed by both water and hemoglobin. The penetration depth is greater than 1 mm and the deeper penetration results in an even greater coagulation layer which leads to higher hemostasis ^{14, 16, 17)}. It has also been reported that higher vaporization capability can be attained at this wavelength ^{14, 17)}. Recent advances in high output diode lasers have led to the development of a diode laser with a maximum output of 300 W ^{12, 15)}. The 980 nm diode laser is anticipated to be the laser with a high vaporization capability. Laser treatment of BPH is performed trans-urethrally using a rigid endoscope. An optical fiber is inserted into the endoscope and laser energy is guided to the prostate tissue. Therefore, the optical fiber used for the procedure directly affects the safety and efficacy of the treatment. In a PVP procedure using the LBO laser, the most popular optical fiber is a side–firing type fiber used in a non-contact mode to the prostate. However from the viewpoint of maneuverability and durability, contact mode optical fibers are gaining attention ^{12, 15)}.

The factor which has the most impact on the safety and efficacy of laser vaporization of the prostate is the depth and intensity of tissue damage caused by the laser irradiation in the prostate tissue beyond that which is vaporized. The temperature of this layer of remnant tissue is raised, and the elevated temperatures induce coagulation, then apoptosis in varying numbers of the deeper target cells which eventually die, with the delayed occurrence of sloughing off of the affected tissue. Thus these photothermal effects will appear from the surface going deep, with the respective depths depending on the power density and irradiation time of the laser ¹⁶⁾. The tissue layers of coagulation and denaturation with apoptotic cell death will slough off from the remaining prostate tissue post-operatively, and may lead to post-operative perforation ¹⁸⁾. At the time of the treatment the depth of the coagulated and denatured apoptotic layers cannot be ascertained with certainty. Therefore, in order for a safe treatment to be performed, elucidation of the relation between laser output and irradiation time, and a quantitative evaluation of irradiation parameters in relation to the size of the irradiation target, are warranted. One way to solve this problem is the accumulation of data through multiple non-clinical investigations. However, this method would require a tremendous amount of time and money. Data accumulation through clinical trials is dependent on the skill level and experience of the attending surgeon. A more objective method of evaluation of safety and efficacy is required.

In recent years, the importance of regulatory science is growing in the field of medical device development. Regulatory science is a scientific method for the evaluation of safety, efficacy and quality of a novel medical device and which is required for the practical usage and promotion of the device ¹⁹⁾. The induction of regulatory science not only may shorten the duration of development and review but is also anticipated to lessen the cost of clinical trials. This present investigation conforms to the precepts of regulatory science and laser vaporization for BPH was investigated through non-clinical methods and simulations, as a precursor study to clinical trials. The current study was able to search for the optimum laser treatment parameters through non-clinical means and simulations in a relatively short period of time, in a highly quantitative and reproducible manner which otherwise would have been time and cost consuming, The current investigation is considered to play a role in the clarification of safety of the treatment through laser irradiation for even severe conditions not possible in clinical settings ²⁰⁾.

The authors have previously reported on the relation of the depth of vaporization and coagulation to laser irradiation parameters using a high intensity diode laser (Ceralas HPD 300, Biolitec, Germany) at the wavelength of 980 nm with a maximum output of 300 W, in combination with a contact mode optic fiber (Twister fiber, Twister™Large Fiber, Biolitec ) through ex-vivo experimentation ²¹⁾. However, in the previous study, the apoptotic cell death layer, sloughing off of which appeared post-operatively, could not be examined experimentally. The occurrence of the apoptotic cell death layer associated with the protein denaturation layer may lead to the perforation of the prostatic capsule through tissue sloughing at some stage postoperatively, and therefore the examination of laser parameters and their relation to the creation of the apoptotic cell death layer is very important.

Purpose

The purpose of this study was to propose safe laser settings for the vaporization of the prostate by estimating the depth and extent of the apoptotic cell death layer through a simulation model constructed by combining light propagation simulation using the Monte Carlo method and heat transfer calculations.

This study was comprised of the following components:

Firstly, measurements of the angular distribution of laser energy were made and the beam diameter at the prostate surface was calculated. A simulation model of light propagation and heat transfer was constructed using this beam diameter.

Secondly, the estimated depths and extent of the protein denaturation and apoptotic cell death layer were compared to that of an ex-vivo prostate laser vaporization experiment. Through this comparison, the validity of the simulation was assessed.

Thirdly, the depths and extents of vaporization layer, and the protein denaturation layer/apoptotic cell death layer were measured based on the simulation model for various laser settings.

Fourth and finally, the results were compared with and studied on MRI images of actual patients from clinical trials. In this study the sum of the depths of the denaturation layer with apoptotic cell death was defined as the depth of thermal denaturation.

Materials and Methods

Measurement of angular distribution of the laser energy emitted from the Twister fiber.

The setup of the goniometer for the measurement of angular distribution of laser energy emitted from the fiber is shown in Figure 1. The position of the tip of the fiber was adjusted to match the center of the rotating breadboard (RBB300/M, Thorlabs). The fiber was fixed to a stage adjustable along the y-axis with a clamp. The laser energy emitted from the fiber was magnified by 5 times with the objective lens (M PLAN APONIR 5X, Mitutoyo) and imaging lens (MT-L, Mitutoyo) and guided through an aperture of 100 µm diameter (P100S, Thorlabs) to a standard photo-diode power sensor (S122C, Thorlabs) where the output was measured. The objective lens and imaging lens were used to improve the angular resolution and to lower the power density at the aperture. From calculations based on geometric optics, the angular resolution of this goniometer optic system is approximately 1°. As shown in Figure 1 the objective lens, imaging lens, aperture and photo-diode power sensor were fixed as a unit and rotated to measure the angular distribution by measuring the output. The axis of the fiber was deemed to be 0°. Angular distribution was measured along the longitudinal and transverse directions. The angular distribution of the longitudinal direction is depicted in Figure 2(a). The angle of maximum power density was deemed to be 0°, while the transverse direction was defined to be perpendicular to the longitudinal direction. The angular distribution of the transverse direction is depicted in Figure 2 (b). The range of measurement of the angular distribution is defined as angle θ = 0∼50° for the longitudinal direction and ϕ = −25∼25° for the transverse direction.

Fig. 1:

Goniometer setup for measurement of the angular distribution of the laser power density.

Fig. 2:

Measurement of the angular distribution of the laser power density in (a) the longitudinal and (b) traverse directions.

Light propagation and heat transfer simulation

In this simulation, calculations of light propagation within the prostatic tissue were made using the Monte Carlo method, and calculations of heat transfer were made using the heat conduction equation. The results were combined for the estimation of the temperature gradient within the prostatic tissue.

The gradient was used in the Arrhenius equation to determine the distribution and ratio of the protein denaturation apoptotic cell death layer. In order to calculate the light propagation within the prostatic tissue, a simulation code based on the Monte Carlo method, which was developed by Wang and released to the public on the website of the Oregon Medical Laser Center, was used ²²⁾. The parameters and their values used for the simulation of light propagation and heat transfer are listed in Table.1. The prostatic tissue model used for the Monte Carlo simulation was based on the assumption that the tissue is flat surfaced with a uniform structure, with a tissue thickness of 10 mm and tissue diameter of 20 mm ^{21, 23)}. The absorption coefficient and scattering coefficient of prostatic tissue at the wavelength of 980 nm have been measured directly in previous studies and were 0.66 cm⁻¹ and 58.8 cm⁻¹ respectively. The refractive index was 1.38 and the anisotropy factor was 0.862 ²¹⁾. The resolution of the calculation was 50 µm for both depth and horizontal directions. The laser light energy assumed to be incident perpendicularly to the tissue surface and the number of trials of the Monte Carlo simulation was set at 1 million times. The absorbed energy densities of each and every coordinates were calculated for when Gaussian laser energy was incident on the prostatic tissue at a certain beam diameter ²²⁾. The beam diameter used for the simulation was determined from the measurements of angular distribution of the laser energy in section Measurement of angular distribution of the laser energy emitted from the Twister fiber.

Table 1: Principal input parameters.

Parameter	Value	Parameter	Value
Tissue thickness	10 mm	Thermal conductivity λ	0.501 W/(m·K)
Tissue diameter	20 mm	Gas constant R	8.314 J/(mol·K)
Absorption coefficient	0.66 cm−1	Frequency factor A
Scattering coefficient	58.8 cm−1	Coagulation	7.39 × 1037 s−1
Refractive index	1.38	Cell death	2.99 × 1080 s−1
Anisotropy factor	0.862	Activation energy ΔE
Resolution	50 µm	Coagulation	2.577 × 105 J/mol
Density ρ	1020 kg/m3	Cell death	5.064 × 105 J/mol
Specific heat Cp	3450 J/(kg·K)

For the calculation of heat conduction, the rise in temperature, derived from the energy densities calculated from the Mont Carlo simulation, and the conduction of heat as time progressed were numerically analyzed. In the two-dimensional heat conduction equation shown below, the energy densities gained from the Monte Carlo simulation were input to the energy density S [W/m³] and the temperature T [°K] at any given time t was calculated.

In this equation ρ is density [kg/m³], C_p is specific heat [J/(kg•K)], λ is thermal conductivity [W/(m•K)]. The above equation was spatially discretized using the finite volume method, and the temperature of the coordinates for each calculation step (Δt = 10 ms) were determined by a implicit method of numerical analysis, the successive overrelaxation method known for its low rate of error. The thermophysical properties of the prostatic tissue are as follows. Density ρ = 1020 kg/m³, specific heat C_p = 3450 J/(kg•K), thermal conductivity λ = 0.501 W/(m•K) ²⁴⁾.

Thermal effect on living tissue is categorized into three thermodynamic processes. The first process is described as the reaction velocity process where heat damage accumulates at a relatively low temperature range of 43°C ∼ 150°C. The second process occurs at temperatures over 100°Cwhere confinement and release of water vapor due to vaporization is seen. The third process occurs at temperatures 300°C ∼ 1000°C where transpiration, combustion, dissociation of molecules and plasma formation are seen ²⁵⁾. With such processes in mind, in this simulation, the maximum temperature was set at 300°C where ablation of tissue starts to occur. Of the coordinates that exceeded 300°C, the temperature was fixed at 300°C and included in the equation as so. The boundary conditions were such that the base of the tissue and point of laser incidence were thermally insulated and the temperature of the tissue lateral boundaries was fixed at 37°C. The depth of ablation was defined as the deepest point exceeding 300°C after the duration of laser irradiation.

Based on the calculations of the temperature gradient within the prostatic tissue, the rate and distribution of protein denaturation and apoptotic cell death were calculated using the Arrhenius equation as cited below. The process of a normal state protein changing to a denatured state can be described as an equation with the protein concentration C.

k is the denaturing speed coefficient, while A [s⁻¹] is the frequency factor, ΔE [J/mol] is the activation energy, R is the gas constant 8.314 J/(mol•K) and T [K] is the absolute temperature. In the case of protein denaturation, A and ΔE are considered to be 7.39 × 10 ³⁷ s⁻¹ and 2.577 − 10⁵ J/mol respectively and in the case of apoptotic cell death, A and ΔE are considered to be 2.99 × 10⁸⁰ s⁻¹ and 5.064 × 10⁵ J/mol respectively ^{27, 28)}. The damage index Ω used to quantify thermal damage is expressed as the equation below.

At any time t, the proportion of denatured protein is expressed as the equation below.

Protein denaturation depth is defined as the difference between the maximum depth where the proportion of protein denaturation volume was greater than 1 − e⁻¹ (∼63.2%) and the vaporization depth while apoptotic cell death depth is defined as the difference between the maximum depth where the volume of cell deaths was greater than 1 − e⁻¹ and the maximum depth where protein denaturation volume is greater than 1 − e⁻¹.

Laser light – prostatic tissue interaction was studied comprehensively on each case when the bottom edge of the side-firing fiber tip was placed at 0, 1 and 3 mm (non-contact distances) from the surface of the prostate. For each non-contact distance, the beam diameters derived from the measurements of angular distribution of laser light in section Measurement of angular distribution of the laser energy emitted from the Twister fiber. were used for the simulation. The non-contact distance of 0 mm meant that the tip of the side-firing fiber was touching the tissue surface. Laser irradiation time was set at 1 s and the incident powers 200 W, 250 W and 300W were simulated. For the non-contact distances of 1 and 3 mm, laser energy attenuation due to light absorbance by water according to Lambert-Beer's law was taken into account 29).

The measurements of vaporization and coagulation depths in ex-vivo experiments

Actual ex-vivo laser irradiation experiments were performed using the 980 nm diode laser at the maximum output of 300 W in combination with the Twister fiber ²¹⁾. The samples used were bovine prostate tissue. The bovine prostates were taken from livestock processed for meat and otherwise would have been thrown away as waste. The number of samples was 5. In order to recreate the clinical settings and conditions, an ex-vivo laser irradiation system was constructed, as shown in Figure 3. In an actual clinical laser vaporization of the prostate, the fiber fixed to the rigid endoscope is inserted trans-urethrally and the prostate surface is irradiated with laser energy via the fiber held in contact with the tissue, which is is irrigated with normal saline solution. In this ex-vivo experimental setup, the fiber was passed through a stainless steel tube with the same channel diameter as the rigid endoscope (9 G, internal diameter 3.19 mm , external diameter 3.75 mm). The fiber was placed so as the tip of the fiber came in contact with the bovine prostate tissue surface. The target tissue was irradiated with laser energy while immersed in normal saline and hence the clinical setting was recreated. The stainless steel pipe was inserted through the side of a plastic case, while the prostatic tissue sample was tacked to a corkboard with pins and sunk to the bottom of the case filled with normal saline solution using weights. The Twister fiber was inserted into the stainless steel tube and the tube was adjusted so that the tip of the fiber contacted the prostatic tissue sample. The height of the sample tissue was adjusted by inserting rubber boards underneath the corkboard. Laser energy was delivered at output powers of 200, 250 and 300 W for an irradiation time of 1 s, in continuous wave mode. After the laser irradiation, the tissue samples were cut in a perpendicular direction to the surface and the depths of vaporization and coagulation were measured.

Fig. 3:

Schematic of the ex vivo experimental setup for contact laser irradiation.

The evaluation of the depths of thermal damage with magnetic resonance imaging (MRI)

MRI images were studied of actual patients taken 2 weeks after being treated with a high intensity 980 nm diode laser, maximum output 300W, in combination with the Twister fiber, (images were provided by the Japan Community Healthcare Organization (JCHO) Sendai Hospital, of patients who had consented to the use of the images). MRI images of 3 patients were used in the present study. The parameters of the laser vaporization are listed in Table. 2. The ethics committee of the JCHO Sendai Hospital granted permission for this clinical study.

Table 2: Patient data and peri-operative data.

Patient No.	Prostate volume [mL]	Power setting [W]	Irradiated energy during operation [kJ]	Irradiation time [min]
1	111	200	305	28
2	85	200	274	23
3	42	200	90	7

MRI imagery was performed with an MRI system (MAGNETOM Symphony 1.5 T, Siemens) under the conditions of repetition time 6.17 ms, echo time 3.09 ms, slice thickness 5 mm and spatial resolution 0.47 mm. From the differences of signal intensity, the widened post-operative urethra, thermally damaged prostatic tissue and normal prostatic tissue were distinguished. Signal intensities of T2 images of MRI are highly dependent on the water content of the tissue. Prostatic tissue thermally damaged by laser irradiation would have less water content compared to normal tissue and thus the image of thermally damaged regions would show as darker or hypointense images. Signal intensities of the widened urethra, thermally damaged prostatic tissue and normal prostatic tissue were measured by averaging the intensities of 5 circular regions of interest (ROI) of each area. The median values between the average intensities of the widened urethra and thermally damaged prostate and of the thermally damaged prostate and normal prostate were considered threshold values. These values were used to reconstruct a binarized image that enhanced the boundaries between the regions. The binarized images were smoothed out with a median filter. Using the binarized images, the thickness of the thermally damaged prostate was measured by measuring the distance from a point of urethra-thermal damage boundary to a point of thermal damage-normal tissue boundary located horizontally from the original point. Fifteen different measurements were made and the average was considered the depth of thermal damage. The average signal intensities of the 3 regions were divided by the standard deviation of signal intensities from noise to calculate the signal to noise ratio (SNR).

Results

The angular distribution of laser light emitted from the Twister fiber

The normalized angular distributions in both the longitudinal and transverse directions of laser light emitted airborne from the Twister fiber are shown in Figure 4. It was assumed that l the point 0 mm from the center to a depth of 0 mm was irradiated with laser energy. In the longitudinal direction, maximum power density was seen at θ = 21° while 1/e² of maximum power density were seen at θ = 9 and 38° and the spread angle of the laser light was 29°. In the transverse direction, maximum power density was seen at ϕ = −1° while 1/e² were seen at ϕ = −14. 12° and the spread angle was 26°. From these results, the cross sectional areas of this elliptic laser beam at non-contact distances 0, 1 and 3 mm were 0.27, 35.6 and 127.2 mm² respectively. The points of 1/e² of maximum power density were considered the ends of the long and short axes of the ellipse. If the shape of the beam were a true circle with the same area as the ellipse and if the laser were to be aimed perpendicular to the prostate surface, the beam diameter at non-contact distances 0, 1 and 3 mm would have been 0.6, 6.7 and 12.7 mm respectively.

Fig. 4:

Angular distributions of the laser power density (a) in the longitudinal and (b) traverse directions.

The results of the ex-vivo experiment and the light propagation and heat transfer simulation

The intra-tissue temperature distribution after laser irradiation at a setting of contact irradiation, 1 s irradiation time and output of 200 W calculated from the simulation and the concentration ratios derived from the Arrhenius' equation of coagulation and cell death are shown in Figure 5. The vaporization depths of both the simulation and ex-vivo experiments at laser setting of contact mode, outputs of 200, 250 and 300 W are shown in Figure 6(a). Since no obvious difference exists between the vaporization depth of both the simulation and experiment, it is highly suggestive that the simulations are as credible as the experiments. The tendency of greater vaporization depth at higher outputs is seen in both the simulation and in the experiment. The coagulation depths under the same conditions are shown in Figure 6(b). The estimated coagulation depth of the simulation was greater than the actual experiment. The differences ranged from 0.6∼0.8 mm. The simulated vaporization depths at the non-contact distance settings of 0, 1 and 3 mm, irradiation time of 1 s and output powers of 200, 250 and 300 W, are shown in Figure 7(a). No vaporization occurred for non-contact distance 1 mm at 200 W output and non-contact distance 3 mm for any output. The changes in coagulation and cell death depths under the same laser settings are shown in Figure 7(b). Coagulation depth is strongly dependent on non-contact distance and was the greatest at the non-contact distance of 1 mm. The thermally denatured depth, which is the sum of the depths of the coagulation and cell death, was greatest at the non-contact distance 1 mm and in the case of 200 W laser output the depth was 4.55 mm.

Fig. 5:

Distributions of (a) temperature and concentration ratios of tissues (b) in coagulation and (c) in cell death to natural tissue.

Fig. 6:

Relationships between (a) the vaporized, or (b) coagulated depth and the power setting of 200, 250, and 300 W in the simulation and ex vivo experiment. Laser irradiation time was set as 1 s. Data of the ex vivo experiment are expressed as means, and the error bars are the standard deviations of the mean.

Fig. 7:

Relationships between (a) the vaporized, or (b) the sum of coagulated and thermally denatured depths and the distance of from the fiber to tissues for each power setting. Laser irradiation time was set as 1 s.

Measurements of thermally denatured depths using MRI images

The T2 weighted image of the laser irradiated prostate and its periphery, its enlarged image and the binarized image reconstructed from the threshold values calculated from the signal intensities are shown in Figure 8. The average SNR of each region was 54 for the urethral space, 27 for the thermally affected region and 37 for normal prostate tissue. It was confirmed that the SNR of the thermally affected region was the lowest. The average depth of the thermally affected region in the images from the 3 patients was 5.1 ± 1.8 mm.

Fig. 8:

(a) MRI image around prostate tissue, (b) magnified image, and (c) binarized image of prostatic lesion in (a).

Discussion

Comparison between the light propagation and heat transfer simulation and the ex-vivo experiment

The comparison of vaporization depths of both the simulation and the ex-vivo experiment is shown in Figure 6. The difference in depth between the two ranged between 0.1 – 0.4 mm. This range is small enough to suggest that the vaporization depths seen in the ex-vivo experiments could be deduced through simulations. The tendency of irradiation with greater outputs of laser energy to produce deeper vaporization depths was also seen in both the simulation and the experiment in a similar manner. The same trend was seen in the coagulation depths of the simulation and the experiment; however the estimated depths by the simulation were 0.6 – 0.8 mm greater than the experiment. One plausible explanation for this phenomenon is that in the simulation, the laser incident surface was set to be thermally insulated. In the experiment, the convective current of the normal saline, in which the tissue was immersed, was flowing while the tissue was being irradiated with the laser, and heat may have been conducted away into the saline by the convection current. Fluid dynamics were not taken into consideration in this simulation and the heat trapped within the insulated surface may have caused the simulation of more severe conditions. However, since the diameter of the prostates of the patients was approximately 60 mm, the difference of coagulation depths was only 1 – 1.3% of the prostate diameter, which is dismissible in any clinical setting. From the above, the results of the simulation and the results of the ex-vivo experiment were practically the same and estimating the effect of laser irradiation of prostatic tissue through simulation is a valid method.

In the experiment, the laser energy emitted from the Twister fiber was incident on the tissue surface at an angle as shown in figure 9, whereas the setting of the simulation was that the laser energy was incident on the surface perpendicularly. This had minimal effect on the results due to the fact the light intensity distribution of the laser energy emitted from the Twister fiber was hemispherical as shown in Figure 5. Laser energy propagates within the tissue through scattering randomly. The optical characteristics of the simulated prostate were set at an absorption coefficient of 0.66 cm⁻¹ and at a scattering coefficient of 58.8 cm⁻¹, meaning that scattering was the predominant factor. The inverse of the scattering coefficient is the average distance where scattering occurs (mean free path) and in the case where the scattering coefficient is 58.8 cm⁻¹, light scatters approximately every 170 µm, leading to a hemispherical light intensity distribution. Therefore, the difference between the calculation of the Monte Carlo simulation where the laser energy was incident perpendicular to the surface, and the actual laser energy incident at an angle, was very small.

Fig. 9:

Schematic of the relationship between the distance from the Twister fiber and laser power density.

Comparison between the light propagation and heat transfer simulation and the clinical trial

The estimated thermally denatured depth resulting from the simulation at the same setting as the clinical trial at laser output of 200 W and laser irradiation in the contact mode was 3.3 mm. However, the measurements of the MRI images revealed thermally denatured depths of 5.1 ± 1.8 mm. Three particular reasons are thought to be the cause of this discrepancy. One is the difference in the method of laser irradiation. In the simulation, laser was aimed at a single point continuously from a fixed fiber, however in a real clinical setting, the laser is never stationary and is moved from point to point. The prostatic tissue is irradiated with the laser several times and thermal damage due to absorption and heat conduction occurs several times. Heat accumulation will therefore gradually be conducted deeper into the tissue creating deeper coagulation and apoptotic cell death depths when compared to that of the simulation. In addition, although the clinical trial was performed with laser in the contact mode, inadvertent non-contact laser irradiation may have occurred which may lead to deeper thermal damage. Second is the technical error inherent in MRI imagery. Since the depths that were measured included regions of irreversible degradation and the regions that would heal to normal tissue post-operatively, a partial volume effect may have affected the measurements. The partial volume effect occurs when multiple tissue types are present within a voxel of the sliced image of the MRI. The signals of the multiple tissue types become averaged and the signal intensity of the tissue borders becomes ambiguous. In the case of laser vaporization of the prostate, the coagulated and cell death layers slough off post-operatively. However, in the regions where thermal damage had not reached the extent of irreversible degradation, it is thought that there exists a layer that will heal to normal and will therefore not become necrotic, and slough off. Since the borders of the urethral space, thermally affected region and normal prostatic tissue are all ambiguous, the depth of thermal damage is a value with inherent error and the actual depth of irreversible denaturation is thought to be smaller than the 5.1 ± 1.8 mm. Thirdly, blood flow was not taken into consideration in the simulation. In the clinical trial, prostatic tissue with blood flow was irradiated with the laser. Some of the photothermal effect during tissue irradiation may have been attenuated via the cooling effect of the blood flow. If clinical laser treatment were to be performed at the same setting as the simulation, single point, continuous laser irradiation, the denaturation depth would be smaller than that of the simulation due to the heat and energy loss through blood flow.

The effect of laser irradiation according to the irradiation parameters in laser vaporization of the prostate.

From the results of the measurements of the angular distribution of laser energy, that energy emitted from the Twister fiber is presumed to spread as is shown in Figure 9. The angles of distribution were θ = 29° for the longitudinal direction and ϕ = 26° for the transverse direction. The laser was incident on the prostate surface at an angle of 69°. The power densities of each non-contact distance were calculated with absorbance by water taken into consideration. The power density at the prostate surface will decrease in an inverse square ratio as the non-contact distance increases as the beam diameter increases, and energy is absorbed by water. Therefore, the reaction induced at the prostate surface is predicted to be strongly dependent on the non-contact distance. In the estimation of the thermally denatured depth by the simulation, the greatest thermally denatured depth was seen at the non-contact distance of 1 mm and the least at the non-contact distance of 3 mm as shown in Figure 7. At the non-contact distance of 1 mm, the power density at the surface of the prostate was not enough for vaporization of the tissue and was limited to coagulation and apoptotic cell death and resulted in the deepest thermally denatured depth. For the simulation, coagulation depth was defined as the difference between the vaporization depth and the maximum depth of coagulation where the rate of coagulation was greater than 1 − e⁻¹ (∼63.2%). At the non-contact distance of 1 mm, the decrease of the vaporization depth led to the increase of the coagulation depth. At the non-contact distance of 3 mm, the enlarged beam diameter led to a substantial decrease in power density at the prostate surface which led to only shallow coagulation and cell death depths. From these findings, it is suggested that the hemostatic capability strongly depends on the non-contact distance. From the above, for the efficient vaporization of the prostate, contact laser irradiation is the best approach. However, when vaporizing the prostate in areas close to the sphincter muscle, contact irradiation may increase the possibility of collateral thermal damage ¹²⁾. In order to avoid such damage, it is suggested that lowering the output when vaporizing near the sphincter muscle may lessen the risk. Situations where non-contact irradiation may be beneficial would be a situation where only hemostasis is required. Creating only the coagulation layer with no vaporization of the prostate is possible. However, close attention must be paid the fact that thermally denatured depth is strongly affected by the non-contact distance.

Finally, based on the above arguments, it may be suggested that laser-prostate tissue interaction at any parameter and setting can be estimated by using the simulation constructed in this study and proposals of safe and effective laser parameters for the treatment can be made.

Conclusion

In the present study, simulation of light and heat transfer within the prostatic tissue was performed and the vaporization, coagulation and apoptotic cell death depths caused by laser irradiation were determined. The results of the simulation were compared to clinical results and studied. The simulation and ex-vivo experiment conducted in this study allowed for a quantitative and highly reproducible analysis of the effect of laser vaporization of the prostate in a short period of time. As a result, comprehensive changes in thermally denatured depths according to non-contact distances and outputs were successfully estimated. From these results, safe and effective laser irradiation methods can be proposed and information to be used in the clinical setting can be provided.