Optical clearing of vaginal tissues, ex vivo, for minimally invasive laser treatment of female stress urinary incontinence

Chun-Hung Chang; Erinn M. Myers; Michael J. Kennelly; Nathaniel M. Fried

doi:10.1117/1.JBO.22.1.018002

12 January 2017 Optical clearing of vaginal tissues, ex vivo, for minimally invasive laser treatment of female stress urinary incontinence

Chun-Hung Chang, Erinn M. Myers, Michael J. Kennelly, Nathaniel M. Fried

Author Affiliations +

Funded by: National Institutes of Health

Journal of Biomedical Optics, Vol. 22, Issue 1, 018002 (January 2017). https://doi.org/10.1117/1.JBO.22.1.018002

Abstract

Near-infrared laser energy in conjunction with applied tissue cooling is being investigated for thermal remodeling of the endopelvic fascia during minimally invasive treatment of female stress urinary incontinence. Previous computer simulations of light transport, heat transfer, and tissue thermal damage have shown that a transvaginal approach is more feasible than a transurethral approach. However, results were suboptimal, and some undesirable thermal insult to the vaginal wall was still predicted. This study uses experiments and computer simulations to explore whether application of an optical clearing agent (OCA) can further improve optical penetration depth and completely preserve the vaginal wall during subsurface treatment of the endopelvic fascia. Several different mixtures of OCA’s were tested, and 100% glycerol was found to be the optimal agent. Optical transmission studies, optical coherence tomography, reflection spectroscopy, and computer simulations [including Monte Carlo (MC) light transport, heat transfer, and Arrhenius integral model of thermal damage] using glycerol were performed. The OCA produced a 61% increase in optical transmission through porcine vaginal wall at 37°C after 30 min. The MC model showed improved energy deposition in endopelvic fascia using glycerol. Without OCA, 62%, 37%, and 1% of energy was deposited in vaginal wall, endopelvic fascia, and urethral wall, respectively, compared with 50%, 49%, and 1% using OCA. Use of OCA also resulted in 0.5-mm increase in treatment depth, allowing potential thermal tissue remodeling at a depth of 3 mm with complete preservation of the vaginal wall.

1. Introduction

1.1.

Female Stress Urinary Incontinence

Over 6.5 million women in the United States suffer from stress urinary incontinence (SUI), but only 3% of these women choose to have surgery.¹^,² General anesthesia, long recovery time, surgical incisions, potential treatment failures associated with future pregnancy, and morbidity of the procedure are several common reasons provided by patients who are reluctant to have surgery. The vast majority of women instead use disposable absorbable products to cope with their symptoms.³ A nonsurgical method that improves patient quality of life, that can be rapidly performed, has minimal morbidity, and only a short recovery time may prove to be a desirable alternative for treatment of female SUI.

1.2.

Preliminary Laser Studies

Our preliminary studies showed that subsurface thermal tissue denaturation can be achieved using deeply penetrating near-infrared laser energy in combination with applied cooling, to preserve 1 to 2 mm of tissue surface from thermal insult.⁴^,⁵ For SUI treatment, the goal is to thermally remodel and tighten endopelvic fascia without damaging surrounding tissue layers (e.g., urethra or vaginal wall). The thickness of the vaginal wall, endopelvic fascia, and urethral wall has been measured with ultrasound to be 2.7, 4.3, and 2.4 mm, respectively.⁶^,⁷ The vaginal wall may potentially be preserved while submucosal tissue is thermally remodeled, using an endoscopic laser probe with integrated cooling for minimally invasive treatment of SUI.

Recent computer simulations of light transport, heat transfer, and tissue thermal damage have demonstrated that a transvaginal approach is more feasible than a transurethral approach.⁸ However, results were suboptimal, and some undesirable thermal insult to the vaginal wall was still predicted.

1.3.

Optical Clearing Agents

Optical clearing agents (OCAs) have been used for reducing light scattering and increasing optical penetration depth in tissues for optical diagnostics (e.g., imaging and spectroscopy) and laser therapy.⁹^–¹⁵ Optical clearing is achieved through both refractive-index matching and tissue dehydration.¹⁶ The OCA induces water flux from the interstitial space to the tissue surface, and the increase in osmolarity of interstitial fluid draws water from the cell and leads to dehydration.¹⁷ These processes provide additional refractive-index matching and a decrease in the reduced scattering coefficient for skin. Decreased scattering may significantly enhance optical transmission, which in turn translates into improved optical penetration depth.

Optical clearing has been successfully demonstrated in a wide range of tissues including skin,¹⁵^,¹⁸^,¹⁹ muscle,¹⁶ collagen,²⁰ cartilage,²¹ and bone²² using different agents and concentrations. For example, recent studies in skin, in vitro, have shown a 53% increase in transmission.¹⁸ In general, higher concentrations and longer immersion times yield higher optical transmission. In some cases, however, improved optical transmission does not correspond to highest concentration. For example, 70% glycerol concentration has been reported to provide deeper light penetration than 90% glycerol in porcine skin.²³ Similarly, an optimal diffusion concentration of 40.5% with glucose and ethylene glycol (EG) mixture instead of higher concentration of EG has been reported. This superior result obtained with a lower OCA concentration is likely to do with the better equilibrium between tissue free-water content and the water in the treatment solution, which yields a maximum agent diffusion rate.²⁴

The primary objective of this study is to determine whether application of an OCA will allow deeper targeting of endopelvic fascia for laser thermal remodeling with complete preservation of the vaginal wall from thermal insult.

2. Methods

Several experimental methods were used in our experiments. In those methods, which are described in detail in the following sections, we have used porcine vaginal tissues with the physical data provided in Table 1.

Table 1

Optical, thermal, and damage parameters for computer simulations of transvaginal and transurethral laser therapy. A laser wavelength of 1064 nm, incident power of 5.0 W, spot diameter of 5.0 mm, and irradiation time of 15 s was used.

Optical properties (phase 1: MC)	Vaginal wall	Endopelvic fascia	Urethral wall
Absorption coefficient ( $μ_{a}$ ) ${cm}^{- 1}$	0.43	0.35	0.5
Scattering coefficient ( $μ_{s}$ ) ${cm}^{- 1}$	21.6 → 5.9	484	239
Anisotropy factor (g)	0.9	0.9	0.9
Index of refraction ( $n$ )	1.38	1.39	1.39
Tissue thickness (mm)	2.7	4.3	2.4
References	6,7,25	6,7,26,27	6,7,28
Thermal properties (phase 2: heat transfer)	Vaginal wall	Endopelvic fascia	Urethral wall
Thermal conductivity ( $κ$ ) $W / m - k$	0.54	0.47	0.46
Specific heat ( $c$ ) $J / kg - k$	3655	3200	3306
Density ( $ρ$ ) $kg / m^{3}$	1088	1085	1102
References	29	29	29
Damage parameters (phase 3: damage integral)	Vaginal wall	Endopelvic fascia	Urethral wall
Frequency factor ( $A$ ) $s^{- 1}$	$5.73 \times 10^{34}$	$1.61 \times 10^{45}$	$5.60 \times 10^{63}$
Activation energy ( $E_{a}$ ) $J / mol$	$2.40 \times 10^{5}$	$3.06 \times 10^{5}$	$4.30 \times 10^{5}$
$T_{crit}$ (°C)	88.0	80.4	78.9
References	28	28	28

2.1.

Optical Clearing Agents

OCAs are commonly mixed with other solutions, for example, dimethyl sulfoxide (DMSO), to enhance the rate of perfusion into tissue.³⁰^–³² Our laboratory reported a 26% increase in transmission through canine skin using a $3 ∶ 1$ mixture of glycerol/DMSO.³³ Other mixtures of OCA’s and transport agents have also been tested, including propylene glycol (PG) and EG. For these studies in porcine vaginal tissue, a minimum of three measurements were performed for each type of the following OCA mixtures tested: 100% glycerol ( $n = 6$ ), 100% PG ( $n = 3$ ), 100% EG ( $n = 3$ ), 100% DMSO ( $n = 3$ ), $3 ∶ 1$ glycerol/DMSO ( $n = 3$ ), $3 ∶ 1$ glycerol/PG ( $n = 3$ ), $3 ∶ 1$ glycerol/EG ( $n = 3$ ), $3 ∶ 1$ PG/DMSO ( $n = 3$ ), and $3 ∶ 1$ EG/DMSO ( $n = 3$ ).

2.2.

Experimental Setup for Measuring Percent Change in Optical Transmission

Fresh porcine vaginal tissues were obtained from Spear Products, Inc. (Coopersburg, Pennsylvania). The tissue was stored in saline and refrigerated for use within 48 h. A continuous-wave ytterbium fiber laser at 1075-nm wavelength and with a collimated beam diameter of 5.2 mm ( $1 / e^{2}$ ) was used. A heat chamber maintained consistent temperature (37°C), and initial intensity was measured before OCA was applied. After OCA application to porcine vaginal tissue, transmitted power was recorded from a detector (PM10, Molectron, Portland, Oregon) connected to a power meter (EPM1000, Molectron) every 2 min for a period of 30 min.

2.3.

Optical Coherence Tomography

Optical coherence tomography (OCT) of vaginal tissue with and without OCA was performed with an OCT system (Niris, Imalux, Cleveland, Ohio) and a handheld 8 Fr probe. The OCT system was a time-domain system based on a common path all fiber interferometer design and operated at a wavelength of 1300 nm, acquiring images with $15 - μ m$ axial resolution and $25 - μ m$ lateral resolution. OCT images of vaginal tissue before and after application of glycerol were acquired to determine increased image depth based on decreased tissue scattering. Measurements were performed using imageJ software (National Institutes of Health, Bethesda Maryland) to analyze the 256 gray scale images.

2.4.

Reflection Spectroscopy

A spectrometer was used to quantify the change in reflection spectrum of porcine vaginal wall as a function of OCA application time. A fiber optic $Y$ cable (RP21, Thorlabs, Newton, New Jersey) was used to connect both the white light source (MI-150, Edmund Optics, Barrington, New Jersey) and the spectrometer (USB 4000-VIS-NIR, Ocean Optics, Dunedin, Florida). The fiber tip was in direct contact with the sample, and reflectance data were collected using the proprietary software, Spectra Suite (ocean optics). By comparing the baseline spectrum from native tissue at 0 min with the spectrum from optically cleared tissue, the percentage change in reflection spectrum due to OCA was calculated.

2.5.

Computer Simulations

Using experimental data on change in transmission by OCA in a known thickness of porcine vaginal tissue, change in scattering and absorption can be determined. Using Beer–Lambert law,

Eq. (1)

I_{1} = I_{0} e^{- μ_{t} x},

Eq. (2)

I_{2} = I_{0} e^{- μ_{t} x},

Eq. (3)

\frac{I_{2}}{I_{1}} = \frac{e^{- μ_{t 2} x}}{e^{- μ_{t 1} x}},

Eq. (4)

μ_{t} = μ_{a} + μ_{s} (1 - g),

where

μ_{t}

is the attenuation coefficient (

{cm}^{- 1}

),

μ_{a}

is the absorption coefficient (

{cm}^{- 1}

),

μ_{s}

is the scattering coefficient (

{cm}^{- 1}

),

g

is the anisotropy factor measuring directional scattering,

x

is the tissue thickness (cm),

I_{0}

is the initial laser intensity (

W / {cm}^{2}

),

I_{1}

is the intensity measured through porcine vaginal tissue at 0 min, and

I_{2}

is the intensity measured through tissue after 30 min of OCA topical application.

A decrease in the reduced scattering coefficient, defined as $μ_{s} (1 - g)$ , is the dominant OCA effect.³⁴ The simulations assume that the absorption coefficient and anisotropy factor remain constant and that only the scattering coefficient changes upon OCA application. This calculation reduces the scattering coefficient from 21.6 to $5.9 {cm}^{- 1}$ as shown below and used in Table 1. From Eq. (2) above, the following calculation is performed:

\frac{I_{2}}{I_{1}} = \exp [(- μ_{t 2} x) - (- μ_{t 1} x)],

\ln {\frac{I_{2}}{I_{1}}} = x [μ_{t 1} - μ_{t 2}],

\frac{1}{x} \ln {\frac{I_{2}}{I_{1}}} = [μ_{t 1} - μ_{t 2}] .

Combining this expression with Eq. (3) results in the following expression:

μ_{t 1} = μ_{a} + μ_{s} (1 - g), g = 0.9,

\frac{1}{x} \ln {\frac{I_{2}}{I_{1}}} = [μ_{a 1} + 0.1 μ_{s 1} - μ_{a 2} - 0.1 μ_{s 2}] .

Since absorption coefficient and anisotropy factor are unchanged, the equation simplifies to

\frac{1}{x} \ln {\frac{I_{2}}{I_{1}}} = 0.1 [μ_{s 1} - μ_{s 2}],

where

μ_{s 1} = 21.6 {cm}^{- 1}

,

x = 0.3 cm

,

I_{1} = 1

,

I_{2} = 1.6

.

Substituting these values into the equation yields a scattering coefficient of $μ_{s 2} = 5.9 {cm}^{- 1}$ .

Optical clearing potential (OCP) is defined as a ratio of $μ_{s}^{'}$ (before) and $μ_{s}^{'}$ (after) [Eq. (4)]. For human skin, glycerol was previously reported to have an average OCP of 3;¹⁹ in this study, glycerol applied to vaginal wall produced a similar OCP of 3.6.

Eq. (5)

OCP = \frac{μ_{s}^{'} (before)}{μ_{s}^{'} (after)},

where

μ_{s}^{'}

is reduced scattering coefficient (

{cm}^{- 1}

) given by

μ_{s}

(

1 - g

).

A Monte Carlo (MC) program was adapted to simulate the distribution of 1 million photons absorbed in tissue layers.³⁵^–³⁷ A convolution program fitted results to a laser beam of known power, profile, and diameter.³⁷ Values for tissue optical properties were compiled based on the Nd:YAG laser wavelength used in previous experimental studies (Table 1).

Temperature simulations were conducted using finite element software (ANSYS 14.5). A heat transfer model was imported from MC model dimensions with mesh of $350 \times 350$ elements ( $28.5 μ m \times 28.5 μ m$ ). Tissue layers were represented in the mesh using the thermal properties from Table 1. Absorption data from MC simulations were convolved and converted into the input of the heat transfer model to simulate heat created by laser irradiation. The initial tissue temperature was 37°C (310 K) with contact probe cooling temperature set to 0°C (273 K).

A standard Arrhenius integral model was employed to predict thermal injury using published values.²⁸ A single damage parameter, $Ω (t)$ , quantified thermal damage to tissue [Eq. (6)]

Eq. (6)

Ω (t) = A \int_{0}^{τ} \exp [- \frac{E_{a}}{R T (t)}] d t,

where

A

(

s^{- 1}

) is the experimentally derived frequency factor,

τ (s)

is the total heating time,

E_{a}

(

J / mol

) is the experimentally derived activation energy of transformation (Table 1),

R

(

8.32 J / K mol

) is the universal gas constant, and

T (t)

is the absolute tissue temperature (kelvin). Thermal damage corresponds to

Ω (t) = 1

. Damage rate increases exponentially when temperature exceeds

T_{crit}

.

Simulations computed thermal damage based on optical, thermal, and damage parameters for the tissues (Table 1). Due to limited availability of damage parameter data, vaginal wall was modeled as smooth muscle, endopelvic fascia as collagen, and urethral wall as aorta tissue.

3. Results

3.1.

Comparison of Different Optical Clearing Agent Mixtures

Figure 1 shows optical transmission increases in vaginal tissue at $λ = 1075 nm$ , as a function of time for various OCAs and mixtures. 100% glycerol provided the best result, an optical transmission increase of 61% after 30 min, compared to 100% EG, PG, and DMSO, which yielded increases of 42%, 41%, and 35%, respectively. Vaginal tissue lacks the barrier properties of skin, and hence the additional mixture of DMSO to other OCA solutions tested, which facilitates the perfusion transport process, did not provide improved optical transmission.

Fig. 1

(a) Experimental setup to measure optical transmission through porcine vaginal tissue, (b) percent increase in optical transmission through porcine vaginal tissue at 37°C, as a function of OCA and application time.

3.2.

Optical Coherence Tomography

Comparison of OCT images taken with and without OCA, and at room (22°C) and body temperatures (37°C), shows how diffusion of glycerol in tissue reduces scattering and allows imaging of deeper tissue layers (Fig. 2). For porcine vaginal tissue at 22°C, after 30 min of glycerol application, a 32% increase in optical transmission was measured using the following method. Since gray scale intensity is directly proportional to photon penetration depth, the $1 / e$ intensity decrease is equal to a gray scale value of 94. At 22°C, in the initial OCT image ( $t = 0 \min$ ), a $1 / e$ intensity decrease was measured at a depth of 0.484 mm. However, after OCA application ( $t = 30 \min$ ), the new $1 / e$ , gray scale value of 94 increased to a depth 0.639 mm, which translates to a 32% increase in optical penetration depth with OCA. Using a similar method, at 37°C, an increase in optical transmission of 59% was measured. Improved optical transmission at higher tissue temperatures is achieved from greater permeability of glycerol due to higher molecular kinetic energy, in agreement with previous studies.¹⁸

Fig. 2

Optical coherence tomography images of porcine vaginal wall in its (a) native state without glycerol, at room temperature, 22°C; (b) after topical application of 100% glycerol for 30 min, at room temperature, 22°C, transmission increased by 32%; (c) native state without glycerol, at body temperature, 37°C; (d) after topical application of 100% glycerol for 30 min, at body temperature, 37°C, transmission increased by 59%; (e) increase in optical transmission using glycerol, for 22°C ( $n = 8$ ) and 37°C ( $n = 6$ ); where higher tissue temperature yielded greater perfusion rate of glycerol in tissue and hence increased optical transmission as well.

3.3.

Reflection Spectroscopy

After OCA treatment, the tissue reflection spectrum was altered, as observed by a change in color at the tissue surface. For porcine vaginal tissue wall, after 30 min of glycerol, the rapid decrease in reflection in green and blue spectrum and slower decrease in reflection in red spectrum was measured, corresponding to a progressively darker red appearance of the vaginal wall (Fig. 3). This is due to major hemoglobin absorption peaks at 420, 540, and 580 nm, which reduce the amount of reemitted light from tissue compared to the red part of the spectrum.

Fig. 3

Porcine vaginal wall in its (a) native state and (b) after topical application of 100% glycerol for 30 min, (c) percent decrease in reflection across the visible and near-IR spectrum. Note that the lowest decrease in reflection occurs in the red (600–700 nm), accounting for the change in appearance of the tissue.

3.4.

Computer Simulations

MC simulations predicted that 50% of energy is absorbed by vaginal wall, 49% is absorbed in endopelvic fascia target layer, and only 1% is deposited in urethral wall. This corresponded to a 12% decrease in energy absorbed by the vaginal wall and 12% increase in energy absorbed in the endopelvic fascia layer using OCA, versus without OCA application.

After applying glycerol for 30 min, scattering in vaginal wall decreased, enabling photons to penetrate deeper into endopelvic fascia where scattering coefficient is greater (Table 1). Energy deposition as a function of depth is shown in Fig. 4. The initial subsurface peak is due to additional contributions of photons through subsurface scattering. The large peak between the vaginal wall and the endopelvic fascia is due to an increase in the scattering coefficient across the interface. The rapid decline in energy absorbed in the endopelvic fascia is due to the high scattering coefficient.

Fig. 4

One-dimensional absorbed energy distribution by porcine vaginal tissue at the center of laser irradiation at 5 W, both with OCA (blue) and without OCA (red).

Heat transfer simulations were also conducted. For similar laser power simulated (5 W, 15 s), use of OCA resulted in higher energy deposition in the endopelvic fascia, which translated into a 7°C higher peak temperature than without OCA [Figs. 5(a) and 5(c)]. However, 15 s after laser was turned off, peak tissue temperatures decayed to a similar value of $\sim 43 ° C$ , with tissue including OCA only 0.5°C higher than tissue without OCA [Figs. 5(b) and 5(d)].

Fig. 5

(a) Two-dimensional temperature distribution (depth x width) after laser irradiation for 15 s. Internal tissue temperature reached 71°C (344°K) at boundary between vaginal wall and endopelvic fascia, (b) temperature distribution in tissue 15 s after laser irradiation, for transvaginal approach. Peak temperature decreased to 42.8°C (316.0°K) at end of 15 s postoperative cooling time period, (c) temperature after laser irradiation for 15 s with OCA reached 78.1°C (351.3°K), (d) peak temperature decreased to 43.3°C (316.5°K) at end of 15 s postoperative cooling time period.

For comparison, the same laser power and time were initially used to contrast thermal damage with and without OCA. Arrhenius integral calculations [Eq. (6)] indicated higher damage in vaginal tissue with OCA ( $Ω = 1.4$ ) compared to without OCA ( $Ω = 0.18$ ) (Fig. 6), due to the higher peak temperature in the endopelvic fascia layer with OCA ( $Δ T = 7 ° C$ ). Previous clinical SUI studies using radiofrequency (RF) energy targeted 65°C in endopelvic fascia for 30 s, based on temperature versus time data for collagen thermal denaturation and shrinkage.³⁸ This corresponds to a damage parameter of $Ω = 0.28$ . Hence, the tissue model without OCA required a power increase from 5.0 to 5.35 W, while the OCA model required a power decrease from 5.0 to 4.45 W, to achieve $Ω = 0.3$ . In practice, OCA treatment is preferred, because it not only enables laser treatment at a lower and hence safer power, but also improves optical penetration depth by 0.5 mm, thus preserving the entire ( $\sim 2.7 - mm$ thick) vaginal wall during the procedure (Fig. 7).

Fig. 6

(a) 5 W without OCA yields $Ω = 0.18$ in endopelvic fascia, (b) 5 W with OCA yields $Ω = 1.4$ in endopelvic fascia.

Fig. 7

(a) Laser power of 5.35 W to thermally remodel endopelvic fascia without OCA ( $Ω = 0.3$ ), (b) laser power of 4.45 W to thermally remodel endopelvic fascia with OCA ( $Ω = 0.3$ ).

4. Discussion

RF energy has previously been used for transurethral thermal shrinkage and remodeling of submucosal collagen without tissue necrosis, in bladder neck and proximal urethra as a safe, nonsurgical SUI treatment with improvement in patient quality of life.³⁹^,⁴⁰ However, RF heating is limited and decays rapidly with depth as $1 / r^{4}$ ( $1 / r^{2}$ term for resistive heating and $1 / r^{2}$ term for electric field strength), resulting in a superficial penetration depth due to direct heating of only 1 to 2 mm.⁴¹^,⁴² RF probes for SUI therapy therefore require needles inserted through the urethral wall and into submucosal tissue for localized heating, denaturation, and shrinkage of collagen, with saline irrigation of mucosa to prevent overheating. RF therapy is thus more invasive than desired by patients seeking a minimally invasive, nonsurgical treatment.

Lasers have recently been used for subablative resurfacing of atrophic vaginal tissue in postmenopausal women, and for treatment of SUI. Both the ${CO}_{2}$ laser at $λ = 10.6 μ m$ ⁴³^–⁴⁵ and erbium:YAG laser at $λ = 2.94 μ m$ ⁴⁶^–⁵⁰ have been used. Short-term studies appear more promising for vaginal rejuvenation, but also show moderate improvements for SUI. However, these laser wavelengths, adapted from the skin resurfacing field, are limited to an optical penetration depth of tens of micrometers and a thermal treatment zone of less than $500 μ m$ .

Preliminary studies in our laboratory have demonstrated that subsurface thermal denaturation of tissues can be achieved using a deeply penetrating laser wavelength ( $λ = 1064$ or 1075 nm) in conjunction with applied surface cooling, preserving 1 to 2 mm of the tissue surface from thermal necrosis.⁴^,⁵ Thus, laser denaturation and shrinkage of the endopelvic fascia may potentially produce improved tissue remodeling results similar to the RF approach, but in a less invasive manner with preservation of the vaginal mucosa.

Computer simulations utilizing optical, thermal, and damage parameters for vaginal wall, endopelvic fascia, and urethral wall were conducted for a transvaginal endoscopic approach to minimally invasive laser treatment of female SUI, with and without an OCA. By combining contact cooling with OCA, it is possible to preserve the entire 2.7-mm thick vaginal wall from thermal insult. In addition, OCA application translated into use of lower laser power, which may provide a safer and less expensive procedure. Glycerol used in these studies is not only biocompatible, but also OCA effects can be easily reversed with tissue returning to its original state after rehydration. The 100% glycerol solution used in this study is an oversaturated solution, which produces tissue transparency primarily by dehydration. The improved transparency is mainly due to a decrease in tissue thickness and better ordering of the tissue components, rather than refractive-index matching between tissue solid components and interstitial fluids.²⁴

These simulations were based on several assumptions. First, it was assumed that experimental OCA results applied to porcine vaginal tissue will be similar to human vaginal tissue. Second, optical clearing was assumed to produce a decrease only in scattering coefficient, but not absorption coefficient or anisotropy factor. Other studies have reported that the decrease in reduced scattering coefficient is primarily due to an increase in the anisotropy factor ( $g$ ) due to structural changes in average collagen fiber size and packing of collagen fibers, rather than an increase in the scattering coefficient.⁵¹ We therefore simulated this alternative case as well. However, the results were very similar. MC simulations predicted that 51% of energy is absorbed by vaginal wall, 48% is absorbed in endopelvic fascia target layer, and only 1% is deposited in urethral wall. During thermal simulations, the peak temperature in the endopelvic fascia reached 79.5°C. An incident laser power of 4.43 W was necessary to achieve a damage parameter of $Ω = 0.3$ .

Third, thermal property and damage parameters used in Table 1 were compiled from different sources, introducing a small error in our simulation results. For example, 5 W with OCA produced a peak temperature of 78.1°C in the endopelvic fascia layer during simulations, slightly below the critical damage threshold of 80.4°C provided in Table 1. However, the damage parameter was $Ω = 1.4$ , a value greater than $Ω = 1$ , which indicates irreversible thermal damage to the tissue.

Finally, dynamic optical property values due to changes in temperature and pressure and changes in thermal properties due to tissue perfusion were not included in this model. Despite these limitations, the simulations demonstrate that use of an OCA may further improve treatment depth for thermal remodeling by $\sim 0.5 mm$ . This should allow preservation of the entire (2.7-mm thick) vaginal wall during subsurface laser targeting and thermal remodeling of the endopelvic fascia for minimally invasive treatment of female SUI.

Future studies will involve histological examination (e.g., hematoxylin and eosin stains) of thermally denatured and remodeled tissue, as well as pivotal in vivo animal studies with functional tests (e.g., Valsalva leak point pressure measurements) to test feasibility of minimally invasive laser therapy for SUI, similar to what has been reported for RF studies.³⁸^,⁴⁸^,⁵² The simulations and results from this study can be utilized as a guide for such studies.

5. Conclusions

This study used experimental studies and computer simulations to explore whether application of an OCA to vaginal wall can further improve optical penetration depth and completely preserve vaginal wall during subsurface laser treatment of endopelvic fascia. Optical transmission studies showed that glycerol produced a 61% increase in transmission through vaginal wall at body temperature (37°C) versus 32% at room temperature (22°C) over 30 min. MC models of light transport in the tissue layers showed improved energy deposition in endopelvic fascia using glycerol. Without OCA, 62%, 37%, and 1% of energy was deposited in the vaginal wall, endopelvic fascia, and urethral wall, respectively, compared with 50%, 49%, and 1% when using an OCA. Application of an OCA also resulted in a 0.5-mm increase in depth for thermal tissue remodeling, allowing 3 mm of tissue surface to be preserved.

Disclosures

None of the authors has a financial interest in this paper.

Acknowledgments

This research was supported by the National Institutes of Health, Grant No. 1R15DK099774.

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Biography

Chun-Hung Chang is currently a PhD student in the Optical Science and Engineering Program at the University of North Carolina at Charlotte.

Erinn M. Myers is currently a physician specializing in female pelvic medicine and reconstructive surgery with the Women’s Center for Pelvic Health at Carolinas Medical Center.

Michael J. Kennelly is currently the director of the Women’s Center for Pelvic Health at Carolinas Medical Center.

Nathaniel M. Fried is currently a professor in the Department of Physics and Optical Science at the University of North Carolina at Charlotte and an adjunct faculty member in the McKay Department of Urology at Carolinas Medical Center. His research interests include therapeutic and diagnostic applications of lasers in urology.

Citation Download Citation

Chun-Hung Chang, Erinn M. Myers, Michael J. Kennelly, and Nathaniel M. Fried "Optical clearing of vaginal tissues, ex vivo, for minimally invasive laser treatment of female stress urinary incontinence," Journal of Biomedical Optics 22(1), 018002 (12 January 2017). https://doi.org/10.1117/1.JBO.22.1.018002

Received: 5 October 2016; Accepted: 22 December 2016; Published: 12 January 2017

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KEYWORDS

Tissues

Optical clearing

Scattering

Laser therapeutics

Transmittance

Computer simulations

Optical coherence tomography

1.

Introduction

1.1.

Female Stress Urinary Incontinence

1.2.

Preliminary Laser Studies

1.3.

Optical Clearing Agents

2.

Methods

Table 1

2.1.

Optical Clearing Agents

2.2.

Experimental Setup for Measuring Percent Change in Optical Transmission

2.3.

Optical Coherence Tomography

2.4.

Reflection Spectroscopy

2.5.

Computer Simulations

Eq. (1)

Eq. (2)

Eq. (3)

Eq. (4)

Eq. (5)

Eq. (6)

3.

Results

3.1.

Comparison of Different Optical Clearing Agent Mixtures

Fig. 1

3.2.

Optical Coherence Tomography

Fig. 2

3.3.

Reflection Spectroscopy

Fig. 3

3.4.

Computer Simulations

Fig. 4

Fig. 5

Fig. 6

Fig. 7

4.

Discussion

5.

Conclusions

Disclosures

Acknowledgments

References

Biography

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