Morphological Characterization of Tissue Destruction According to the Distance Between Holmium:YAG Laser Tip and Tissue Surface

Article information

Int Neurourol J. 2024;28(3):196-206

Publication date (electronic) : 2024 September 30

doi : https://doi.org/10.5213/inj.2448332.166

Dongho Kang^1,*, Ohbin Kwon^2,*, Seong-Chan Kim², Kanghae Kim³, Youngseok Seo⁴, Seung-June Oh^,1,5, Min-Joo Choi^2,6

¹Department of Medical Device Development, Seoul National University College of Medicine, Seoul, Korea

²Interdisciplinary Postgraduate Program in Biomedical Engineering, Jeju National University, Jeju, Korea

³Department of Biomedical Science, Dankook University College of Medicine, Cheonan, Korea

⁴Wontech Co., Ltd., Daejeon, Korea

⁵Department of Urology, Seoul National University College of Medicine, Seoul National University Hospital, Seoul, Korea

⁶Department of Medicine, School of Medicine, Jeju National University, Jeju, Korea

Corresponding author: Seung-June Oh Departments of Urology and Medical Device Development, Seoul National University College of Medicine, Department of Urology, Seoul National University Hospital, 101 Daehak-no, Jongno-gu, Seoul 03080, Korea Email: sjo@snu.ac.kr

Co-corresponding author: Min-Joo Choi Interdisciplinary Postgraduate Program in Biomedical Engineering, Jeju National University, Department of Medicine, School of Medicine, Jeju National University, 102 Jeju Daehak-no, Jeju 63243, Korea Email: mjchoi@jejunu.ac.kr

*Dongho Kang and Ohbin Kwon contributed equally to this study as co-first authors.

Received 2024 July 26; Accepted 2024 September 6.

Abstract

INTRODUCTION

Surgery for urological diseases such as urinary stones, benign prostatic hyperplasia (BPH), and urethral stricture is conducted widely in a less invasive manner using an endoscope and laser in a liquid medium. Among these, endoscopic surgery for urinary stones and holmium laser enucleation of the prostate (Ho-LEP) [1] for BPH using a holmium:yttrium aluminum garnet (Holmium:YAG) laser has brought about revolutionary changes in urological surgeries. This is due to the unique energy characteristics of the holmium laser. The basic mechanism of urinary stone destruction is relatively well known.

A localized laser energy delivered through the optical fiber tip located in a liquid medium produces a heated gas bubble that encircles the tip. The bubble grows when the laser emits, and it collapses after the laser pulse is off. The collapse generates a pressure wave which interacts with urinary stones to destroy [2-4]. HoLEP requires the implementation of various surgical modes, such as incision of prostate tissue, separation between the prostatic adenoma and prostatic capsule, and hemostasis [5]. In actual clinical practice, these modes are implemented by manually adjusting the distance between the laser fiber tip and the tissue surface. The mechanism of the tissue destruction by the laser pulse in HoLEP is presumed to differ from that for breaking urinary stones. However, the detailed mechanism by which the holmium laser destroys prostatic tissue is largely unknown.

The present study was to perform a morphological study using a soft tissue phantom (STP) to understand the impact of the distance between the laser fiber tip and the phantom surface (DLP) on the tissue destruction by a Holmium:YAG laser and to discuss how the surgical modes are related to the DLP. To our best knowledge, this is the first study that addresses the issue.

MATERIALS AND METHODS

Experimental Setup

Holmium:YAG laser (2,100 nm) equipment (Holinwon Prima, Wontech, Daejeon, Korea) that is commercially available for clinical use was employed in the present experiment [6]. The laser fiber tip used a core diameter of 550 μm. To examine the reproducibility of the laser pulses, the 10 repeated emissions irradiated at 2 J in a water bath were measured. As shown in Fig. 1B, the laser pulse started to appear from a time delay of ~0.34 msec after the laser module was switched on (t=0), and the full width at half maximum was measured to be 0.3 msec. The temporal structure of the laser pulses remained unchanged but the (peak) power was shown to fluctuate within +/- 15% of the mean value of ~1.5 W.

Fig. 1.

(A) Experimental setup which consists of Holmium:YAG laser (2,100 nm), soft tissue phantom (STP) submerged in a water bath, and an optical system to view the responses of the STP to the laser pulse. (B) Temporal beam profiles of the Holmium:YAG laser pulse measured 10 times at the output setting of 2 J. Note that the time zero (t=0) represents the moment when the laser module is switched on, and the FWHM stands for full width at half maximum which was measured to be about 0.3 msec.

The STP was made of a polyacrylamide (PAA) gel with a polymer concentration of 16% and a water concentration of 84% [7] (Table 1). The PAA gel is optically transparent and has mechanical properties [8,9] similar to those of prostate tissue, and its structure decomposes at temperatures above 100°C. The laser irradiation results in cracks or tears in the STP, depending on the DLP. An additional STP was considered to which bovine serum albumin (BSA) was added. The BSA STP is denatured when its temperature is above a threshold temperature, and it becomes opaque in a temperature range of 60°C–80°C [10,11].

Table 1.

Composition of the soft tissue phantoms considered inthis study for the volume of 50 mL

Experimental setup is schematically drawn in Fig. 1A. In a water bath (Acrylic Tank, Onda USA LLC, Delaware, OH, USA), the laser fiber tip and the STP (15 mm×35 mm×15 mm, Jeju BME Lab, Korea=BME Lab, Jeju, Korea) were fixed with a transparent acrylic base and an XYZ positioner (3-Axis Positioner, Onda USA LLC), so that the magnitude of the DLP was accurately controlled. The water temperature was measured to be 20°C±1°C and the laboratory room temperature was 22°C± 1°C. The laser-induced cavitation bubble and the resulting tissue destruction in the STP and were recorded at 50,000 FPS (80× 108 pixels) using a high-speed camera (V642, Vision Research lnc., Wayne, NJ, USA), as the DLP was altered from 0 to 4 mm. A lamp (LT-V9-15, GsVitec GmbH, Bad Soden Salmünster, Germany) was used to continuously illuminate the other side of the camera to capture contrast images of either the bubble or the destruction.

For the STP taken from the water bath after the laser irradiation, the depth and width of the cracks (or destruction) formed in the STP were photographed with a camera (EOS 5D Mark III, Canon, Tokyo, Japan) in the surface and lateral views (camera settings: 1/25s, F2.8 and ISO100). LED lights were shone above and below the STP to visualize the destruction in the transparent, while the camera lens focused on the center of the STP.

RESULTS

A Laser Pulse-Induced Bubble and Its Penetration Into STP

When the single laser pulse (2 J) was irradiated into water, a heated gas bubble was produced and grew at the tip of the optical fiber during the laser emission for a little longer than the laser pulse width (~0.3 msec). The bubble reached its maximum size larger than ~4 mm in diameter at t=~0.7 msec when the laser irradiation was terminated. Process of the bubble growth is not presented here but the fully grown bubble would be similar to that for the largest DLP of 4 mm shown in the first row of the last column in Fig. 2A. The fully grown bubble is large and unstable so that it rapidly collapses when the laser irradiation was terminated.

Fig. 2.

Effects of a single laser pulse (2 J) to the soft tissue phantom (STP) at the different distances (0, 1, 2, 3, 4 mm) between the laser fiber tip and the phantom surface (DLP), observed by a high-speed camera: (A) The cavitation bubbles generated by the laser and their penetration into the STPs. The images in the first row were obtained at t=0.7 msec at the moment when the laser emission stopped and the bubble was grown to be the largest. Those in the second and third row were captured at t=1.5 msec and 3.0 msec, respectively under the condition without the laser irradiation. (B) The depth (left) and width (right) of the bubble penetration measured on the images captured at t=0.7 msec are plotted against the DLP.

The laser produced bubbles exhibited different shapes, depending on the DLP (Fig. 2). The more distortion was observed from a spherical shape at the smaller values of the DLP. At the largest DLP of 4 mm, the bubble generated at the laser tip was elongated toward the STP, which was measured to be 7.4±0.2 mm in the length and 6.1±0.08 mm in the width. It is of interest to note that, even though the bubbles collapsed and disappeared, the penetrated parts were dynamically active (oscillating) for a long time which was observed from the video images.

The penetration depth and width of the bubble into the STP were measured at t=0.7 msec (illustrated in the second column of the first row in Fig. 2A) and they were shown to vary with the DLP. At the DLP of 0 mm, the penetration depth and width were 4.5±0.1 mm and 2.8±0.3 mm, respectively. As the DLP increased, the penetration depth and width decreased (P<0.05). At the DLP lager than 3 mm, no penetration was observed.

Destruction in STP by a Single Laser Pulse

The dynamic bubble penetration into the STP may result in the permanent destruction in the STP. The destruction formed in the STPs by the single laser pulse was photographed with a camera after the laser irradiation (Fig. 3A and B for the lateral and surface views, respectively). At the DLP of 0 mm, the destruction depth was measured to be 3.0±0.2 mm, and the width was 0.6±0.2 mm in the lateral view of the photographic images of the STP (Fig. 3C). The destruction depth was the greatest at the DLP of 0 mm and decreased significantly with the DLP (P<0.05). As expected, the surface damages were observed to be larger than the internal destruction width for the same value of the DLP. The maximum surface damage was measured to be 1.4±0.4 mm (Fig. 3C and D). As the DLP increased, the extent of the destruction decreased (P<0.05), and there was no destruction in the STP beyond the DLP of 3 mm (Fig. 3). With the single laser pulse, the extents of the surface damage did not change significantly from the DLP less than 2 mm (P>0.05) but they decreased significantly beyond the DLP of 3 mm (P<0.05).

Fig. 3.

Permanent destruction formed in the soft tissue phantoms (STPs) by a single laser pulse (2 J) at a different DLP of 0, 1, 2, 3, and 4 mm, photographed with a camera after the laser irradiation: (A) the lateral views for 5 STPs, (B) the surface views for 5 STPs, (C) the depth and width (manually measured from the 5 lateral view images) against the DLP, and (D) the surface damage size versus the DLP measured from the 5 surface view images. DLP, distance between the laser fiber tip and the phantom surface.

Destruction of STP by Multiple Laser Pulses

As seen in Fig. 4, the depth and width of the damage by the 60 laser pulses were observed to decrease significantly (P<0.05) as the DLP increased. Unlike the single laser pulse scenario, the surface damages by the multiple laser pulses decreased significantly with the DLP (P<0.05). Note that, at the DLP of 0 mm, damages in the optical fiber tip (burn-back phenomenon) occurred in the 2 experimental cases (at the second and the seventh laser pulse, respectively, in each case).

Fig. 4.

(A) Typical shapes of the destruction formed in the soft tissue phantom (STP) by the 60 laser pulses (2 J) repeated at 12 Hz at the DLP of 0, 1, 2, 3, and 4 mm, photographed with a camera after the laser irradiation (the upper panels for the lateral views and the lower for the surface views), (B) the depth and width (manually measured for 5 STPs) against the DLP, and (C) the measured surface damage size for 5 STPs versus the DLP. DLP, distance between the laser fiber tip and the phantom surface.

Assessment of the Thermal Effects by Multiple Laser Pulses on BSA STP

Multiple laser pulse irradiation to the BSA STP is expected to result in discoloration in the adjacent area to the destruction (Fig. 5). This is because the BSA contained in the optically transparent STP is denatured when the temperature rise up over a threshold level of about 60°C as described in the experimental setup (MATERIALS AND METHODS). The discoloration (dark brown in Fig. 5) was observed at the value of DLP up to 3 mm, which occurred in a circular region surrounding the destruction (black on the images). The circular discolored areas are clearly visible on the surface view of images. The discolored area was shown to decrease in size as the DLP increased (Fig. 5). No discoloration was observed at the largest value (4 mm) of the DLP. At the DLP of 3 mm, there is a discolored concentric circle with an annular ring shown without any destruction. Note that the discoloration is interpreted to represent the thermal coagulation in the tissue closely related to hemostasis. This underpins that the DLP 3 mm would be an optimized condition for hemostasis.

Fig. 5.

The destruction and the thermal effects produced in the BSA STP by the 60 laser pulses (2 J) repeated at 12 Hz at the DLP of 0, 1, 2, 3 and 4 mm, photographed with a camera after the laser irradiation (the upper panels for the lateral views and the lower for the surface views). Note that the dark brown color represents the protein denaturation in the BSA STP due to thermal effects by laser pulses. BSA, bovine serum albumin; STP, soft tissue phantom; DLP, distance between the laser fiber tip and the phantom surface.

DISCUSSION

In urology, less invasive endoscopic surgical methods for treating urinary stones and BPH have been established. To achieve the destruction of urinary stones and endoscopic removal of prostatic adenomas, an independent energy source is required. Energy sources such as electric energy and lasers have been used for this purpose. Endoscopic urinary tract surgery differs from surgery in other areas because it is performed in a liquid medium. Basic research on the medical use of holmium lasers has focused on urinary stones; therefore, the research results on urinary stones are relatively well known. The mechanism of urinary stone destruction by a holmium laser involves the cavitation effect [12,13]. The shock wave generated via cavitation crushes the urinary stones through a mechanical effect. In addition, the heat energy generated in the plasma of the bubble destroys urinary stones through a thermal effect [3,4,14].

HoLEP enables surgeons to separate the plane between the prostate capsule and the enlarged prostatic adenoma using a holmium laser. Although the clinical use of holmium lasers for BPH has increased over the past 20 years, little is known about the underlying mechanisms of their action. Our present findings showed that, as with urinary stones, bubbles are formed at the laser tip and penetrate the STP, causing its rupture. In other words, direct rupture of the STP occurred due to the mechanical effect of cavitation. The thermal effect also appeared to be complex. Our data demonstrated that the extent of mechanical rupture decreased with increasing DLP values between the laser tip and the target tissue.

Clinicians use holmium lasers on soft tissues, such as prostate tissue, in ways other than simply destroying solid objects, such as urinary stones. Urologists use holmium lasers for tissue incision, hemostasis, and tissue separation [5]. A surgical plane can be developed between the prostatic adenoma and the prostatic capsule by mechanical effects such as pressure generated by a holmium laser. In HoLEP, this separation effect is termed enucleation and it constitutes the most important surgical principle of HoLEP surgery.

Clinicians empirically implement various laser action modes by adjusting the distance between the laser tip and prostatic tissue (analogous to DLP) rather than changing settings such as Joules or pulse frequency. Urologists generate an incision effect by directly contacting the prostate tissue with the holmium laser tip under water, a hemostatic effect by distancing it a little from tissue, and an earthquake effect by moving it further away [5]. When the laser tip directly contacts the prostate tissue (analogous to a DLP of 0 mm), a thermal effect materializes, but the pressure effect is much stronger and causes tissue incision rather than vaporization or coagulation [15]. To achieve hemostasis in HoLEP, surgeons empirically increase the distance between the laser and tissue (analogous to DLP) by about 2 mm, termed ‘defocusing’ [16,17] and the temperature effect becomes greater than the mechanical effect, maximizing the coagulation effect [5]. According to previous reports [18,19], the laser-generated bubble turns into a plasma state and the temperature at the tip of the bubble rises exponentially. Therefore, it can be assumed that the location where the edge of plasma is formed is located about 2 mm from the laser tip at the typical setting of 80 W, 2 J, 40 Hz [1]. These clinical findings are consistent with the phenomena observed in our experiments. When the laser tip is close to tissue (DLP less than 1 mm), direct tissue destruction occurs leading to incision. As the DLP gradually increases farther to 2–3 mm, the destruction effect weakens but thermal coagulation is still effective, shifting to hemostasis. In real tissue, this distance is expected to be somewhat different from that in the STP, but this phenomenon and its explanation will still hold true.

When the laser tip is spaced further apart from tissue, the temperature effect is attenuated and pressure is transmitted, creating a separation between the prostatic capsule and the adenoma [13]. In other words, when the laser tip is located far from tissue (DLP >3 mm in our experiment), the repeated radiation forces of the pressure waves resulting from the repetitive bubble collapses induce the separation being vibrational at the repetitive frequency. It is assumed that the capsular plane is developed through this effect. Separation occurs even when the distance between the laser and the tissue is very large because the pressure effect can be propagated from a distance when the medium is water [13]. In real-world clinical practice, in a narrow space such as the capsular plane where the prostate is to be separated, the pressure is further amplified, and a vibration effect is expected to occur. This is because the prostate is a closed space rather than an open space as in the experimental design of the present study. In the present study, we found no mechanical destruction or thermal damage at a distance >4 mm.

Since incision or enucleation is possible due to the mechanical effect caused by the shock wave generated when the laserinduced bubble collapses, theoretically, enucleation can be achieved by increasing the size of the bubble. Previous research has already shown that tissue destruction depends on the size of the bubble. As energy (J) rises, the size of the bubble increases [20]. The size of the fiber tip also greatly affects the size of the bubble. It has been shown that as the size of the fiber tip increases, the bubble size also increases [21]. However, stronger pressure waves may be generated with larger bubbles when hemostasis is needed, and this may have the disadvantage of not being effective in hemostasis. Therefore, some urologic surgeons [5,22,23] hold the view that it is preferable to lower the laser energy for hemostasis. Additionally, the biggest parameter that determines the shape of the bubble is pulse duration [24]. When the pulse duration is short (short pulse), small circular bubbles are generated in the water, while longer pulses generate long oval or pear-shaped bubbles [8]. Empirically and clinically, even if the amount of energy is the same, it can be observed that when the pulse is long, the thermal effect is greater than when the pulse is short, making hemostasis easier. More research will be needed on this phenomenon in the future.

The main issue of this experimental study was the effect of DLP on the STP. To our knowledge, there has been no other study on this topic yet. However, our study has one limitation. The STP we designed may not be completely identical to actual prostate tissue. In this study, the STP was made to resemble the optical and thermal properties of the prostate as much as possible. However, there is still a possibility that it may not mimic all properties, and additional research using biological tissues such as liver tissue [25] may be necessary in the future. We hope that this experimental study will not only be useful in helping clinicians understand the mechanism of HoLEP but will also serve as basic data for improving surgical equipment for the treatment of BPH.

In conclusion, this study illustrated that soft tissue destruction was associated with the cavitation bubble generated by the holmium laser. For the first time reported experimental evidence that various surgical modes in HoLEP, such as incision, hemostasis and separation, could be achieved by controlling the DLP.

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