Open Access
Add to BibSonomyAbstract
An optical diffuser was developed to achieve radially uniform light irradiation by micro-machining helical patterns on the fiber surface for endoscopically treating urethral stricture. Spatial emission from the diffuser was evaluated by goniometric measurements. A computational model was developed to predict spatio-temporal heat distribution during the interstitial coagulation. The fabricated diffuser yielded circumferential light distribution with slightly concentrated energy at the proximal end. Both simulation and tissue testing demonstrated approximately 1-mm coagulation thickness at 6 W for 10 sec with 1470 nm. The proposed optical diffuser may be a feasible tool to treat the urethral stricture in a uniform manner.
© 2015 Optical Society of America
1. Introduction
Urethral stricture is scarring of urethral pathway involving a fibrotic process of spongy erectile tissue (i.e., corpus spongiosum) [1]. The corpus spongiosum lies under urethral epithelium, of which fibrosis and contraction often lead to scarring and eventual reduction in a urethral luminal diameter (i.e., spongiofibrosis). According to the data from Veterans Affairs (VA) in 2003, a prevalence rate of the urethral stricture was estimated to be 0.2% (i.e., 193 per 100,000 diagnoses) in US [2]. Among the diagnosed population, approximately 20% of patients develop symptomatic strictures due to iatrogenic causes, such as traumatic catheterization, penal inflammation, congenital disease, and urological injuries to genitourinary tracts [3]. The patients with the urethral stricture typically experience obstructive voiding symptoms including weak urinary stream, hematuria, pain, incomplete bladder emptying, and burning urinary infection [3]. Depending upon characteristics of the stricture (i.e., location, etiology, size, and density of spongiofibrosis), several invasive treatments have been performed [4]. For instance, dilation with balloon or catheter is a temporary treatment that stretches fibrotic tissue to widen the urethral lumen [3]. Stent is another mid-term treatment option that places a small tube in a urethra to keep the stricture open. Urethrotomy is a non-thermal procedure (called “cold knife”) performed under anesthesia that surgically removes the urethral tissue to relieve the stricture. However, these treatment techniques have demonstrated high complication and recurrence rates associated with local pain, discomfort, and stone formation due to incomplete removal of the scarring tissue [3].
Lasers have been used for a variety of medical applications, such as tissue ablation, blood coagulation, surgical resection, and optical imaging [5–9]. Due to short optical penetration depth associated with absorption and scattering events, non-invasive applications of lasers have often been limited. Thus, optical fibers have been used instead to incise, coagulate, and vaporize the diseased tissues. In particular, for photothermal treatment of urethral stricture, an optical fiber (end- or side-firing) in conjunction with a rigid or semi-rigid cystoscope is typically inserted into a genital to internally ablate the targeted scarring tissue (called “hot knife”) [9–13]. However, since the fiber transmits high intensity light in a forward or lateral direction, it can hardly achieve uniform light distribution in tissue, which may cause non-homogenous coagulation and even carbonization during the thermal treatment. Furthermore, different wavelengths achieve various tissue responses in terms of coagulation and ablation, depending on the degree of light absorption. Thus, in order to obtain a more precise depth profile in tissue, a 1470 nm wavelength can often be employed due to its optical penetration depth combined with increased absorption in the tissue water [5]. In the current study, to compensate directional limitations of the fibers for treating tubular structure tissue, a diffusing optical fiber was designed and developed to cylindrically transmit uniform laser light (1470 nm) with low irradiance to the urethral stricture. It was hypothesized that based upon photothermal interactions, a relatively wide range of temperature distribution and heat accumulation in the urethral tissue could induce coagulative necrosis to selectively treat the stricture with minimal thermal injury to the underlying corpus spongiosum. Unlike the cold and hot knife applications associated with high complications, the proposed technique as a thermal procedure may immediately treat a wide extent of the urethral tissue without bleeding, eventually facilitating healing process and enhancing the procedure success.
Various biomedical applications of diffusing optical fibers have been investigated in the field of phototherapeutics including photodynamic therapy (PDT) [14–16], laser hair removal [17], laser lipolysis [18], and bladder carcinoma treatment [8]. To fabricate the optical diffusers, a number of manufacturing processes have been proposed; introduction of scattering media [19], melt-drawing technique [20], acid-etching [18, 21], and UV micro-drilling [21]. For PDT applications, a thin layer of scattering media (e.g., raw quarts and a mixture of optical epoxy and Al2O3 powder) was added to the tip of an optical fiber to achieve volumetric light diffusion merely at low power levels [19, 22–24]. Micro-needles were fabricated for light diffusion by using either melt-drawing or acid-etching techniques to have extremely a small diameter of shafts (33-125 µm) [18, 20]. Thus, the microneedles were pertinent to the limited fabrication size and lengths (i.e., shorter than 1 cm) [18, 20]. In the case of UV micro-drilling technique, the irradiance emitted from the diffuser presented a ripple profile due to the gaps among multiple arrays of micro-holes on the fiber surface. In addition, the micro-holes in the diffuser induced low structural strength, leading to micro-cracking and variations in the output intensity [21]. Furthermore, the proposed fabrication techniques have accompanied multiple complex procedures that take at least a few hours [18, 21] to complete the fabrication process in a reliable manner. Additionally, the techniques were often difficult to control spatial emission profiles from the diffuser tip in terms of uniform light distribution and minimal forward propagation [18, 21].
In order to overcome the previous fabrication limitations, we developed a new technique to fabricate diffusing optical fibers by means of laser-assisted micromachining. For homogeneously circumferential light distribution with delivery of high power, a flat-cut ended bare fiber (flat fiber) was micro-machined by laser to engrave patterns on the fiber surface. The transmitted light could be diffused by varying the critical angles inside the fiber to reduce the probability of the total internal reflection. The fabricated fiber tips were evaluated in terms of microscopic imaging and goniometric measurements to characterize spatial uniformity of light propagation. Spatio-temporal heat distribution in tissue was also modeled to predict the degree of thermal coagulation for optimizing the treatment parameters during the interstitial irradiation of 1470 nm. Then, in vitro tissue testing with a 1470-nm laser along with histological analysis was performed to construct thermal dosimetry to optimize the laser parameters for irreversible tissue coagulation.
2. Methods
Figure 1 presents a set-up for laser micro-fabrication of a diffusing optical fiber. A 30-W CO2 laser system (λ = 10.6 µm, Synrad, Mukilteo, WA, USA) was employed for the fabrication process. A laser beam in 3-mm diameter was expanded to approximately 7.5 mm by a combination of concave lens (f1 = −25 mm) and convex lens (f2 = 75 mm). Then, the expanded beam was focused on the fiber surface by another convex lens (f3 = 25 mm). The focal spot was estimated to be around 45 µm in diameter, and the corresponding irradiance was 3.1 × 105 W/cm2 for glass ablation. An air cooling system was used to maintain clean surface on the machining area by removing any particles and melted parts. Multimode silica optical fibers with a 600-µm core diameter (NA = 0.48, Thorlabs Inc., Newton, New Jersey, USA) were used for the current study, and the distal end of each fiber was solely fabricated for radial emission. After removal of buffers with a mechanical stripper, an optical fiber was positioned in a fiber holder that was fixed at a motorized-stage system consisting of both translational and rotational stages as shown in Fig. 1. To obtain helical groove patters, the translational stage moved the fiber along the x-axis direction at the speed of 2~7 mm/s while the rotational stage was simultaneously implemented to rotate the fiber at 240~480 rpm. Each movement speed was deliberately controlled and selected to avoid any overlapping between two consecutive grooves during the fabrication process. In turn, the fabrication angle between the fiber axis and the groove path was adjusted to be 45°, and the total length of the diffusing tip was 10 mm. The depth of each groove was fabricated to be around 20 µm, which was significantly smaller than the core-diameter of the fiber (600 µm) to maintain the structural integrity. Thus, the damage threshold of the diffuser could be roughly 250 kW/cm2 for continuous wave or 1 GW/cm2 for 10-ns pulsed mode. All the motorized-stages were driven by a motion control system in conjunction with LabView software (National Instrument Corp., Austin, Texas, USA). To validate the helical patterns induced by laser fabrication, the surface of the optical diffuser was imaged with scanning electron microscope (SEM). A HeNe (λ = 632 nm, Thorlabs Inc., Newton, New Jersey, USA) laser was also employed to transmit laser light through the fabricated fiber and to roughly visualize the degree of light diffusion for swift evaluations on the fiber. The diffused light was imaged with a digital camera (D5100, Nikon Co., Ltd, Tokyo, Japan) in a dark room. To protect the fabricated part from any mechanical damage, the diffuser tip was covered with a glass cap (i.e., length = 13 mm, outer diameter = 1.4 mm, and thickness = 0.2 mm) and sealed with epoxy.
Fig. 1 Illustration of laser fabrication set-up for optical diffuser (H: fiber holder, T: translational stage, and R: rotational stage).
Spatial distribution of light intensity emitted from a fabricated diffuser tip was evaluated with a goniometric system, as shown in Fig. 2. The measurement set-up consisted of a photodiode head (PD-300-3W, Ophir, Jerusalem, Israel) and a power meter (Nova II, Ophir, Jerusalem, Israel), based upon the methodology from Vesselov et al. [25]. A HeNe laser was used as a light source for the goniometric measurements. The photodiode with a sensing area of 1 cm2 was situated on a rotational rail platform with a radius of 57 mm to measure the power propagated from the fiber. For polar emission measurements, the diffuser was vertically positioned, and the photodiode sensor was rotated around the axis of the diffusing fiber at the middle point of the diffusing tip (5 mm from distal end) in Fig. 2(a). Regarding azimuthal emission measurements, the sensor was rotated around the tip of the horizontally-located fiber, of which the axis of rotation was perpendicular to that of the fiber in Fig. 2(b). A step size for both the polar and the azimuthal measurements was 10°, and the spatial resolution of the emission point on the diffuser was approximately 50 µm. Lastly, the sensor was moved by 0.5 mm along the fiber to measure longitudinal light emissions at various diffuser positions. The photodiode was positioned 1 mm above the diffuser, and a 0.5 × 0.5 mm2 diaphragm slit was placed between the photodiode and the diffuser to prevent the detection of any diffusive light. All the measured intensities were normalized to minimize the effect of any fluctuations in the input power on the measurements.
Fig. 2 Schematic of goniometric measurements: (a) polar and (b) azimuthal emission measurement setups. Note a dash-dot line represents an axis of rotation for a photodiode.
Numerical simulations were conducted to estimate temperature distribution and the degree of thermal damage during laser coagulation, by using a finite-element method (FEM) software (v3.7, COMSOL Inc., Burlington, Massachusetts, USA). A model was based upon a glass-capped diffuser (13 mm long and 1.4 mm in outer diameter) that was inserted into a tissue specimen of 30 × 30 × 30 mm3. A 0.2-mm air gap existed between the diffuser and the glass cap. Porcine liver was used as a target tissue and irradiated by 1470-nm light at 3, 6, and 9 W for 15 sec: the total energy delivered was 45, 90, and 135 J respectively. Thermal properties of each material were considered temperature-independent and listed in Table 1. During the laser irradiation, heat transport within the tissue was described by using a bio-heat transfer equation as follows [17]:
where ρ (kg/m3), c (J/kg·K), and k (W/m·K) are density, specific heat, and thermal conductivity of the liver tissue respectively, T (K) the local tissue temperature, and Qext (W/m3) the heat source associated with light absorption. Neither metabolic heat generation nor blood perfusion was considered in the current model owing to in vitro conditions. Due to application of a near-infrared wavelength of 1470 nm, no scattering effect was assumed, and light propagation in the liver was primarily governed by Beer’s law [26]. In the current model, the laser intensity was simplified to consist of two directions of laser power, based upon goniometric measurements: radially emitted power from the diffusing part, P1 (72% of the incident laser power) and forward emitted power from the fiber tip, P2 (18% of the incident laser power). In turn, for the radial direction, the heat source was estimated as follows:where μa (cm−1) is the absorption coefficient of the tissue, μs (cm−1) the scattering coefficient of the tissue, r (m) the radial distance from the diffuser surface, and l (m) the diffuser length. μa of porcine liver tissue was 19 cm−1 [27] (no scattering assumed; μs = 0). For the forward direction, the light was diverged from the fiber tip (NA = 0.48), and the spatial beam profile was assumed to be Gaussian. Thus, the heat source was a function of both the radial and the axial distances:where σ (µm) is the deviation of Gaussian distribution of the laser beam (i.e., 100 µm resulting from 99% of the output laser power with 300-µm beam radius) and z (m) the axial depth in tissue. The initial temperature was set to 293 K, and the external tissue surface was insulated.
Table 1. Thermal Properties of Materials Used in Simulation Model
Lastly, the extent of thermal damage in tissue was described by a first-order thermal-chemical rate equation, which represents the rate when molecules become thermally denatured [30]. Damage was quantified by using a single parameter, Ω that was calculated from the Arrhenius equation [30]. Ω is dimensionless, exponentially dependent on the temperature and the exposure time as follows:
where Af (1/s) is the frequency factor, Ea (J/mol) the denaturation activation energy, R (J/mol·K) the universal gas constant of 8.314, T (K) the absolute temperature in tissue, and τ (sec) the duration of laser irradiation, which was 10 sec for the current damage model. Ω = 1 represents the extent of the irreversible thermal damage that occurs in 63% of tissue due to temperature increase up to approximately 338 K [30]. All the physical properties of the liver were assumed to be constant during the tissue coagulation and summarized in Table 2.Table 2. Physical Properties of Liver Tissue Used for Thermal Damage Model
In an attempt to determine the thermal dosage for irreversible tissue denaturation, porcine liver tissue was tested with an optical diffuser in vitro. A 10-W 1470-nm laser system (Veincare, Wontech, Daejeon, Korea) was implemented as a light source for the tissue testing. The liver tissue procured from a local abattoir was selected as a tissue model due to high light absorption features at 1470 nm (µa = 19 cm−1) [27]. Each tissue specimen was prepared in size of 3 × 3 cm2 with flat surface and stored at 277 K prior to the tissue tests. For interstitial application of the optical diffuser, a 15-mm long cylindrical hole (1.4 mm in diameter) was initially created inside the tissue. Then, the glass-capped optical diffuser was inserted into the hole, and the 1470 nm laser light was interstitially and circumferentially irradiated to the liver tissue under various conditions. Three different power levels (3, 6, and 9 W) were implemented for the in vitro tissue tests at various irradiation times from coagulation threshold to 8 sec after the threshold with an increment of 2 sec (i.e., five test conditions). The total energy delivered was 48, 72, and 81 J respectively. Irreversible thermal denaturation was evaluated in light of discoloration on the tissue surface. The coagulation threshold was thus defined as the onset time of any evident discoloration on the tissue surface after the laser irradiation at each power level. Each condition ran five times (N = 5) for the entire testing. Upon the irradiation tests, each specimen was cross-sectioned along the axis of the optical diffuser, and the cross-sectional area was imaged with a portable microscope. Coagulation depth was defined as the distance from the axis of the diffuser to the outer boundary of the discolored tissue region. Five measurements were conducted from the proximal to the distal end of the diffuser with Image J (National Institute of the Health, Bethesda, MD). For statistical analysis, Students’ t-test was performed, and p < 0.05 indicated significance.
To explore the feasibility of transurethral treatment for urethral stricture, porcine urethral tissue (1.5 cm long and 1.5 mm in inner diameter) was experimented with a glass-capped optical diffuser in vitro. The urethral tissue was obtained from a local slaughter house and stored in isotonic saline at 277 K to prevent dehydration and structural deformation. Prior to coagulation tests, the diffuser was inserted through the urethral tissue, which was placed on a microscope slide in air. Due to the size of glass cap, the diffuser slightly made contact with the urethral wall. Based upon dosimetric studies with porcine tissue, the urethral tissue was irradiated for 10 sec at 6 W (i.e., total energy delivery of 60 J) with 1470-nm laser in order to achieve coagulation depth of approximately 1 mm (N = 3). The average irradiance on the tissue surface was calculated to be approximately 11 W/cm2. Upon the laser treatment, the tested tissue was fixed in 10% neutral-buffered formalin (VWR International, Westchester, PA, USA) for five days and embedded in paraffin. Then, the treated region was sectioned into 4 µm thick specimens and stained with hematoxylin and eosin (H&E). The prepared slides were imaged with an optical transmission microscope, and the extent of thermal denaturation was circumferentially measured around the urethral tissue ten times with Image J (N = 8).
3. Results
Figure 3(a) demonstrates a SEM image of the fiber surface fabricated at 45° with a CO2 laser system. The image showed uniform helical groove patterns on the surface in a rhombic shape with a 100-µm side. The groove width between two consecutive patterns was measured to be around 25 µm, indicating no overlapping between the patterns. The angle between the fiber axis and the helical patterns overtly validated the fabrication angle of 45°. Figure 3(b) presents the gross light distribution emitted from the diffuser tip by using a HeNe laser. Overall, photons were radially propagated along the optical diffuser without any significant of light delivery at the distal end. It was observed that relatively higher intensity was accumulated near the proximal end of the diffuser.
Fig. 3 Images of optical diffuser fabricated at 45°: (a) SEM image (15 kV, × 250) of fiber surface and (b) HeNe light distribution from fiber tip (P: proximal and D: distal ends)
For quantitative evaluations on light diffusion, Fig. 4 exhibits spatial distribution of the normalized intensity from a diffuser tip in conjunction with HeNe light. Polar emission results demonstrated almost isotropic radiation with less than 5% deviation from the average angular intensity over 2π (i.e., 0.94 ± 0.03; Fig. 4(a)). The difference between the highest and lowest intensities was 11% (i.e., 0.89~1). Thus, the irradiance distribution on the plane normal to the diffuser axis was associated with a concentric circle without any distortion, indicating homogenous and circumferential radiation from the diffuser tip. In Fig. 4(b), the overall shape in azimuthal emission shows two symmetric side-lobes at the angles of 60° and 300°, representing backward light diffusion. The lateral light distribution (i.e., 60~150° and 210~300°) also seemed rather uniform. The emission at the tip (i.e., 150~210°), however, revealed relatively lower light intensity distribution. In Fig. 4(c), longitudinal emission demonstrated an asymmetric and relatively less homogeneous profile of the light distribution with less than 10% deviation along the diffuser (i.e., 0.80 ± 0.07). The highest intensity occurred near the proximal end of the diffuser, and the intensity immediately dropped at the location 3 mm away from the proximal end, which corresponded to the qualitative observation in Fig. 3(b). Then, the light intensity gradually increased along the diffuser tip toward the distal end. The difference between the highest and lowest intensities was estimated to be less than 30%. It was noted that no significant difference was found in the emission profiles between 632 and 1470 nm. The power measurements validated that 18% and 72% of the input power were transmitted in forward and radial directions respectively whereas 10% loss occurred due to insertion and transmission.
Fig. 4 Normalized light intensity measured from 10-mm optical diffuser: (a) polar, (b) azimuthal, and (c) longitudinal emission profiles.
Figure 5(a) demonstrates the temporal development of temperature at the glass-tissue interface under various power levels. Overall, the temperature rise increased with the applied power and irradiation time. As tissue coagulation typically commences at the temperature of around 338 K [30], the heating rates to initiate the irreversible coagulation were estimated to be approximately 6, 11, and 16 K/sec for 3, 6, and 9 W respectively. Figure 5(b) presents the spatial distribution of tissue temperature along the fiber after 10-sec irradiation at 6 W. The temperature was developed symmetrically around the diffuser axis. Due to a combination of forward and radial emissions, both the fiber tip and the diffusing part yielded two distinctive regions of the heat distribution. Due to the forward emissions, heat generation was diverged from the fiber tip in a conical shape (around 57°). The peak temperature at the tip was found to be 333 K, and then, the temperature rapidly decreased on account of light absorption and scattering by the tissue during the beam divergence. On the other hand, the radial emissions from the diffusing parts yielded almost symmetrical distribution in a flat-top shape. The peak temperature at the interface between the glass-cap and the tissue was 373 K, which was higher than that at the distal end. The temperature at the inner surface of the glass cap still remained high (around 372.5 K) due to its higher thermal conductivity than that of tissue (i.e., kglass = 1.3 W/m·K vs. ktissue = 0.513 W/m·K in Table 1).
Fig. 5 Simulation results: (a) temporal development of temperature at tissue interface at 3, 6, and 9 W and (b) spatial distribution of temperature after 10-sec irradiation at 6 W. A line represents the threshold temperature for coagulation.
Figure 6(a) demonstrates a cross-sectional image of the temperature distribution from the middle of the fiber after 10-sec irradiation at 6 W. Apparently, the temperature achieved circumferential expansion around the diffuser. Figure 6(b) represents the extent of irreversible tissue damage (Ω = 1) corresponding to the temperature distribution in Fig. 6(a). The thermal denaturation was almost uniformly distributed around the diffuser axis with coagulation thickness of 1.3 ± 0.2 mm.
Fig. 6 Simulated thermal responses of tissue after 10-sec irradiation at 6 W: (a) cross-sectional view of temperature distribution and (b) thermal damage with log10 of Arrhenius integral.
Figure 7(a) shows consecutive events of thermal denaturation during 6-W irradiation with a fabricated diffuser on liver tissue in vitro. Each image shows the tissue dissected along the diffuser axis after coagulation. The onset of coagulation (i.e., discoloration) was found to be at 4 sec after laser irradiation. The proximal end initially yielded more denatured area (top image in Fig. 7(a)). Upon the coagulation, the denatured region apparently grew with irradiation time, eventually forming a narrow lopsided funnel shape (bottom image in Fig. 7(a)). It should be noted that the coagulation pattern was similar to the emission profiles in Figs. 3(b) and 4(c). The effect of power on coagulation depth was also investigated in terms of irradiation time in Fig. 7(b). Overall, the depth linearly increased with both the irradiation time and the applied power (R2 = 0.95~0.99). Higher power levels resulted in lower coagulation thresholds (e.g., 1 sec at 9 W vs. 8 sec at 3 W), and the coagulation rates were estimated to be 0.06, 0.08, and 0.10 mm/sec at 3, 6, and 9 W respectively. Particularly, the coagulation depth measured at 6 W for 10 sec irradiation was approximately 20% thinner than the simulation results in Fig. 6(b) (i.e., 1.1 ± 0.1 mm for tissue vs. 1.3 ± 0.2 mm for simulation p < 0.001).
Fig. 7 In vitro tissue coagulation with optical diffuser: (a) development of irreversible tissue denaturation at 6 W with 1470 nm (P: proximal and D: distal ends; bar = 5 mm) and (b) variations in coagulation depths as function of time for three different power levels (N = 5).
Figure 8 presents H&E-stained histological images of urethral tissue treated at 6 W for 10 sec. The treatment conditions were selected to attain approximately 1 mm of thermal denaturation in the tissue. The images correspond to the coagulated region that was positioned at the middle point of the tested diffuser. In Fig. 8(a), coagulative necrosis was vividly observed in the discolored and almost circumferential region. The extent of coagulation thickness was measured to be 1.3 ± 0.2 mm, which was slightly thicker (18%) than the liver testing under the same conditions (i.e., 1.1 ± 0.1 mm in Fig. 7(b); p < 0.001). In addition, the coagulation depths measured at the proximal and distal ends of the diffuser were 1.5 ± 0.3 and 1.3 ± 0.1 mm, respectively. Thus, the quantitative evaluations confirmed that almost uniform coagulation patterns were created along the axis of the urethra (p = 0.29~0.78) in the current study. In Fig. 8(b), the irreversible denaturation was clearly evidenced with the presence of epithelial cellular death, endothelial injury, and formation of vacuolation. The preserved tissue still contained layers of smooth muscle cells without any cellular deformation, which is observed in a native urethral tissue.
Fig. 8 H&E-stained histological images of urethral tissue treated at 6 W for 10 sec: (a) gross image ( × 12.5; bar = 2 mm) and (b) magnified image of transition zone ( × 100; bar = 200 µm).
4. Discussion
The goal of this study was to develop a novel optical fiber to deliver laser light radially to treat a cylindrical tissue such as urethra stricture. Laser-assisted micromachining was implemented to create periodic structural patterns on the fiber surface for light diffusion. For the sake of simplicity and reproducibility of fabrication process, helical groove patterns were selected and continuously engraved onto the surface of the fiber tip. The proposed fabrication typically requires less than five minutes creating the entire superficial deformation with almost unvarying fiber diameter (cylindrical geometry). In turn, the current process could provide a considerably faster and simpler fabrication procedure than the previous other techniques [18, 20, 21]. According to Figs. 4(a) and 4(c), both polar and longitudinal emission profiles quantitatively evidenced that the fabricated surface patterns in Fig. 3(a) yielded the normalized deviation of 3% (polar) and 9% (longitudinal) whereas the previous diffusers showed roughly 5~11% (polar) and 20~28% (longitudinal) [25]. Thus, the current diffusers were able to obtain a better homogeneity in both directions. In addition, the azimuthal emission shown in Fig. 4(b) indicated that less amount of the light with lower irradiance could be delivered in a forward direction, compared with conventional flat fibers. Accordingly, it is conceived that the proposed optical fiber could ensure favorable light delivery during endoscopic photothermal treatments.
According to the numerical simulations in Fig. 5(b), spatial distribution of temperature largely reflected a comparable shape of tissue coagulation in Fig. 7(a), where the tissue temperature experienced around 338 K or higher [30]. However, due to 18% of laser power emitting in a forward direction, the temperature distribution at the distal end was diverged in a hemi-spherical manner as shown in Fig. 5(b) whereas in vitro testing demonstrated a funnel shape of coagulation continuously developed in the tissue as shown in Fig. 7(a). This discrepancy could be explicated by spatial light distribution at the diffuser tip. Although the forward beam was assumed to be Gaussian in the simulations, the tested diffuser could have transmitted the non-Gaussian beam due to partial ablation at the tip circumference. In turn, the light intensity delivered merely from the center could narrow the extent of coagulation necrosis. In addition, both the decreased µa and the increased µs during the photocoagulation [27] could further restrict temperature development, unlike the hemispherical heat diffusion in the simulations. Furthermore, the in vitro liver testing yielded the overall coagulation depth approximately 20% thinner than the simulation model at 6 W for 10-sec irradiation (i.e., 1.1 ± 0.1 mm for tissue testing vs. 1.3 ± 0.2 mm for simulation; p < 0.001). The thinner coagulation depth in the tissue could also result from variations in optical properties. Unlike the assumption in the simulations, both absorption and scattering events are typically inconstant. Thus, during the tissue denaturation, the value for µa could be overestimated for the numerical analysis while the scattering events (µs) could be concomitantly augmented due to temperature dependence [27, 31]. Another explanation for the difference could be pertaining to the empirical values of Af and Ea as well as dynamic effects of tissue properties (optical and thermal), which could affect thermal sensitivity during the interstitial photocoagulation. Thus, incorporation of dynamic optical/thermal responses of tissue will improve computation accuracy as well as specifically predict the extent of thermal necrosis in the tissue [27, 31]. Moreover, in vitro urethral tissue testing confirmed an almost circular form of coagulative necrosis in Fig. 8(a), which shows a good agreement with the degree of irreversible damage in Fig. 6(b). Unlike the liver testing, almost equivalent thermal coagulation (p = 0.99) was found between the numerical model and the treated urethral tissue, which could be attributed to slightly higher light absorption characteristics of the urethral tissue at 1470 nm (µa,urethra = 22.5 cm−1 = 90% of interstitial water in urethra 25 cm−1 of water absorption coefficient [32, 33]) than the liver (µa,liver = 19 cm−1). It was noted that both the tissues had the comparable thermal diffusivity (αurethra = 1.410−7 m2/s [34] vs. αliver = 1.310−7 m2/s [35]).
In vitro tissue experiments in Fig. 8(a) demonstrated almost concentric coagulation in urethra with the eccentricity of 0.3 mm (i.e., thicker at 5 o’clock) despite the fact that the urethral wall was circumferentially coagulated. Although the longitudinal emission in Fig. 4(c) presented less homogeneity than the radial emission, the urethra testing still exhibited that the extent of thermal denaturation was quite consistent along the diffuser in Fig. 8(a) (i.e., 1.3 ± 0.1~1.5 ± 0.3 mm; p = 0.29~0.78). In particular, almost 30% more intensity was radially delivered from the proximal end whereas the coagulation thickness increased merely by 10%. Both the out-of-center phenomenon and the discrepancy could be explained by the blind deployment of the optical diffuser during the tests. Since the inner structure of the tissue was irregular, the diffuser was difficult to be precisely located at the center of the urethra and to be in contact with the entire tissue surface for securing consistent thermal denaturation. In fact, the end-point of laser stricture treatment is to extend the narrowed urethral passage for easy urination without development of inflammation or infection. Hence, the integration of an inflatable stent or catheter with the proposed diffuser is currently under consideration to deploy the transurethral device in a reliable and concentric manner. The dilation of the integrated device will also apply mechanical pressure to the urethral wall in a radial direction and conceivably induce structural deformation during the irradiation. Furthermore, in vivo experiments using a mini-pig model [36] will be conducted to identify acute and chronic responses of the treated tissue as well as healing process associated with infiltration of macrophages. The feasibility of achieving the complete stricture treatment should thus be evaluated terms of various disease geometries and therapeutic parameters including diffuser length, wavelength, pulse configurations, and irradiance.
In spite of reproducible and circumferential tissue coagulation, experimental limitations still remain in photon distribution. For instance, azimuthal emissions presented that a certain amount of light was still transmitted in a forward direction (i.e., 150~210° in Fig. 4(b)), which could be associated with safety issues during transurethral treatment. To further minimize the forward propagation, a conical tapering method is currently integrated into the fabrication process, in that a reduction in the cross-sectional diffuser area can increase the probability of the photons being diffused laterally. Due to precise material removal, acid etching will also be combined with the tapering process in order to improve directivity performance of the diffusers. Another improvement is related to deletion of the abrupt intensity peak near the proximal end as shown in Fig. 4(c). According to microscopic evaluations, the observed discontinuities were associated with fabrication delays at the interface due the reverse change of machining directions. Thus, the protracted interaction time between the machining laser and the fiber surface led to formation of relatively deeper and wider holes near the proximal end, accompanying more light diffusion. Therefore, the concomitant light defocusing technique with the use of high resolution stages will be incorporated into the laser machining process in order to enhance the fabrication precision as well as to attain smooth machining transitions particularly at the boundary. Although all the experiments were performed in air, the diffuser can be surrounded by irrigation saline during the transurethral procedures. Thus, the increased refractive index could reduce the overall divergence angle from the diffuser, eventually varying the applied irradiance and the spatial extent of coagulation in tissue. Accordingly, the effect of the surrounding media on the light emission should be evaluated quantitatively for clinical applications. Moreover, the current fabrication process merely implemented the fixed machining parameters (i.e., 45° for fabrication angle, 100-µm segment size, and 25-µm groove width in Fig. 3(a)). Since the flat-top lateral emission was the eventual goal of the current fiber machining, further dosimetric investigations will be performed to optimize a variety of fabrication parameters including patterning angle, groove width, and segment size. Various coating materials can also be applied to the glass cap not only to protect the fiber tip from mechanical shock and but also facilitate the light diffusion at the cap surface.
5. Conclusion
The current study demonstrated a simple and effective method to fabricate a diffusing optical fiber. The cylindrical light diffusion achieved radially uniform coagulative necrosis in urethral tissue. Further developments on fabrication techniques will be pursued to entail a flat-top profile of longitudinal emissions. The proposed diffusing fiber can be a feasible delivery device to treat urethral stricture in a safe and predictable manner.
Acknowledgments
The authors thank H. Shin for his help on tissue testing and M. Ahn for tissue preparation. This research was supported by a grant from Marine Biotechnology Program (20150220) funded by Ministry of Oceans and Fisheries, South Korea.
References and links
1. C. Chapple, D. Andrich, A. Atala, G. Barbagli, A. Cavalcanti, S. Kulkarni, A. Mangera, and Y. Nakajima, “SIU/ICUD Consultation on Urethral Strictures: The management of anterior urethral stricture disease using substitution urethroplasty,” Urology 83(3Suppl), S31–S47 (2014). [CrossRef] [PubMed]
2. R. A. Santucci, G. F. Joyce, and M. Wise, “Male urethral stricture disease,” J. Urol. 177(5), 1667–1674 (2007). [CrossRef] [PubMed]
3. A. Mangera and C. R. Chapple, “Urethral stricture disease,” Surgery 32(6), 304–309 (2014). [CrossRef]
4. R. Mathur, D. Nayak, G. Aggarwal, A. Shukla, F. Khan, and S. Odiya, “A retrospective analysis of urethral rtrictures and their management at a Tertiary Care Center,” Nephrourol. Mon. 3, 109–113 (2011).
5. M. Amzayyb, R. R. van den Bos, V. M. Kodach, D. M. de Bruin, T. Nijsten, H. A. Neumann, and M. J. van Gemert, “Carbonized blood deposited on fibres during 810, 940 and 1,470 nm endovenous laser ablation: thickness and absorption by optical coherence tomography,” Lasers Med. Sci. 25(3), 439–447 (2010). [CrossRef] [PubMed]
6. H. W. Kang, J. Kim, and J. Oh, “Enhanced photocoagulation with catheter-based diffusing optical device,” J. Biomed. Opt. 17(11), 118001 (2012). [CrossRef] [PubMed]
7. R. G. Kolkman, W. Steenbergen, and T. G. van Leeuwen, “In vivo photoacoustic imaging of blood vessels with a pulsed laser diode,” Lasers Med. Sci. 21(3), 134–139 (2006). [CrossRef] [PubMed]
8. K. Miyazaki, Y. Morimoto, N. Nishiyama, Y. Maekawa, W. Z. Hu, K. Nakatate, K. Kaneda, N. Shinomiya, and K. Kataoka, “A novel homogeneous irradiation fiber probe for whole bladder wall photodynamic therapy,” Lasers Surg. Med. 44(5), 413–420 (2012). [CrossRef] [PubMed]
9. R. K. Singal, J. D. Denstedt, H. A. Razvi, and S. S. Chun, “Holmium:YAG laser endoureterotomy for treatment of ureteral stricture,” Urology 50(6), 875–880 (1997). [CrossRef] [PubMed]
10. M. Atak, H. Tokgöz, B. Akduman, B. Erol, I. Dönmez, V. Hancı, O. Türksoy, and N. A. Mungan, “Low-power holmium:YAG laser urethrotomy for urethral stricture disease: comparison of outcomes with the cold-knife technique,” Kaohsiung J. Med. Sci. 27(11), 503–507 (2011). [CrossRef] [PubMed]
11. B. A. Kamal, “The use of the diode laser for treating urethral strictures,” BJU Int. 87(9), 831–833 (2001). [CrossRef] [PubMed]
12. M. Gürdal, A. Tekin, E. Yücebaş, S. Kireççi, and F. Sengör, “Contact neodymium: YAG laser ablation of recurrent urethral strictures using a side-firing fiber,” J. Endourol. 17(9), 791–794 (2003). [CrossRef] [PubMed]
13. H. Hibi, K. Mitsui, T. Taki, H. Mizumoto, Y. Yamamda, N. Honda, and H. Fukatsu, “Holmium laser incision technique for ureteral stricture using a small-caliber ureteroscope,” JSLS 4(3), 215–220 (2000). [PubMed]
14. M. F. Marqa, P. Colin, P. Nevoux, S. R. Mordon, and N. Betrouni, “Focal laser ablation of prostate cancer: numerical simulation of temperature and damage distribution,” Biomed. Eng. Online 10(1), 45 (2011). [CrossRef] [PubMed]
15. P. Saccomandi, E. Schena, F. M. Di Matteo, M. Pandolfi, M. Martino, R. Rea, and S. Silvestri, “Laser Interstitial Thermotherapy for pancreatic tumor ablation: theoretical model and experimental validation,” Conf. Proc. IEEE Eng. Med. Biol. Soc. 2011, 5585–5588 (2011). [PubMed]
16. H. Wenger, A. Yousuf, A. Oto, and S. Eggener, “Laser ablation as focal therapy for prostate cancer,” Curr. Opin. Urol. 24(3), 236–240 (2014). [CrossRef] [PubMed]
17. K. Hashimoto, M. Kogure, T. L. Irwin, K. Tezuka, T. Osawa, K. Kato, and T. Ebisawa, “Permanent hair removal with a diode-pumped Nd:YAG laser: a pilot study using the direct insertion method,” J. Am. Acad. Dermatol. 49(6), 1071–1080 (2003). [CrossRef] [PubMed]
18. M. A. Kosoglu, R. L. Hood, J. H. Rossmeisl Jr, D. C. Grant, Y. Xu, J. L. Robertson, M. N. Rylander, and C. G. Rylander, “Fiberoptic microneedles: Novel optical diffusers for interstitial delivery of therapeutic light,” Lasers Surg. Med. 43(9), 914–920 (2011). [CrossRef] [PubMed]
19. M. H. Niemz, Laser-tissue Interactions: Fundamentals and Applications (Springer, 2004).
20. M. A. Kosoglu, R. L. Hood, Y. Chen, Y. Xu, M. N. Rylander, and C. G. Rylander, “Fiber optic microneedles for transdermal light delivery: ex vivo porcine skin penetration experiments,” J. Biomech. Eng. 132(9), 091014 (2010). [CrossRef] [PubMed]
21. V. V. Volkov, V. B. Loshchenov, V. I. Konov, and V. V. Kononenko, “Fibreoptic diffuse-light irradiators of biological tissues,” Quantum Electron. 40(8), 746–750 (2010). [CrossRef]
22. H. Shangguan and L. W. Casperson, “Estimation of scattered light on the surface of unclad optical fiber tips: a new approach,” Opt. Commun. 152(4-6), 307–312 (1998). [CrossRef]
23. H. Shangguan, T. E. Haw, K. W. Gregory, and L. W. Casperson, “Novel cylindrical illuminator tip for ultraviolet light delivery,” in Proceedings of SPIE Conference on Diagnostic and Therapeutic Cadivovascular Intervention (SPIE, 1993), pp. 167–182. [CrossRef]
24. V. V. Tuchin, “Lasers and fiber optics in medicine,” in Proceedings of SPIE Conference on Optical Methods of Biomedical Diagnostics and Therapy (SPIE, 1993), pp. 2–16. [CrossRef]
25. L. M. Vesselov, W. Whittington, and L. Lilge, “Performance evaluation of cylindrical fiber optic light diffusers for biomedical applications,” Lasers Surg. Med. 34(4), 348–351 (2004). [CrossRef] [PubMed]
26. G. V. G. Baranoski and J. G. Rokne, “An algorithmic Reflectance and Transmittance model for plant tissue,” Comput. Graph. Forum 16(3), C141–C150 (1997). [CrossRef]
27. J. P. Ritz, A. Roggan, C. Isbert, G. Müller, H. J. Buhr, and C. T. Germer, “Optical properties of native and coagulated porcine liver tissue between 400 and 2400 nm,” Lasers Surg. Med. 29(3), 205–212 (2001). [CrossRef] [PubMed]
28. K. J. Bergemann, R. Krasny, and S. R. Forrest, “Thermal properties of organic light-emitting diodes,” Org. Electron. 13(9), 1565–1568 (2012). [CrossRef]
29. J. Kwon, C. Y. Lee, J. Oh, and H. W. Kang, “Computational analysis of endometrial photocoagulation with diffusing optical device,” Biomed. Opt. Express 4(11), 2450–2462 (2013). [CrossRef] [PubMed]
30. B. Cox, “Introduction to laser-tissue interactions,” PHAS 4886, 1–61 (2007).
31. A. N. Bashkatov, E. A. Genina, V. I. Kochubey, A. A. Gavrilova, S. V. Kapralov, V. A. Grishaev, and V. V. Tuchin, “Optical properties of human stomach mucosa in the spectral range from 400 to 2000 nm: Prognosis for gastroenterology,” Med. Laser Appl. 22(2), 95–104 (2007). [CrossRef]
32. S. L. Jacques, “Optical properties of biological tissues: a review,” Phys. Med. Biol. 58(11), R37–R61 (2013). [CrossRef] [PubMed]
33. R. R. van den Bos, P. W. van Ruijven, C. W. van der Geld, M. J. van Gemert, H. A. Neumann, and T. Nijsten, “Endovenous simulated laser experiments at 940 nm and 1470 nm suggest wavelength-independent temperature profiles,” Eur. J. Vasc. Endovasc. Surg. 44(1), 77–81 (2012). [CrossRef] [PubMed]
34. J. A. Balderas-López, A. Mandelis, and J. A. Garcia, “Thermal-wave resonator cavity design and measurements of the thermal diffusivity of liquids,” Rev. Sci. Instrum. 71(7), 2933–2937 (2000). [CrossRef]
35. J. W. Valvano, J. R. Cochran, and K. R. Diller, “Thermal conductivity and diffusivity of biomaterials measured with self-heated thermistors,” Int. J. Thermophys. 6(3), 301–311 (1985). [CrossRef]
36. K. D. Sievert, C. Selent-Stier, J. Wiedemann, T. O. Greiner, B. Amend, A. Stenzl, G. Feil, and J. Seibold, “Introducing a large animal model to create urethral stricture similar to human stricture disease: a comparative experimental microscopic study,” J. Urol. 187(3), 1101–1109 (2012). [CrossRef] [PubMed]
