1.

Introduction

Lung cancer is the leading cause of cancer-related mortality worldwide, with nearly 1.4 million deaths each year.^1–3 Lung cancer is diagnosed at an advanced stage in a majority of patients, which is the primary reason behind the high mortality rate associated with this disease. Early detection continues to be an elusive goal, and substantial numbers of patients diagnosed with localized disease are often unsuitable for curative surgical procedures due to concomitant medical illness. Non-small cell lung cancer (NSCLC), which includes adenocarcinoma, squamous cell carcinoma, large cell carcinoma, and bronchioloalveolar carcinoma, accounts for nearly 85% of all cases of lung cancer.^2,3 Cigarette smoking is the most common etiological factor, accounting for nearly 85% of patients with lung cancer.^2,3

White light reflectance (WLR) color imaging endoscopy is the standard for clinical detection and localization of lung cancer. However, it suffers from poor sensitivity for early cancer detection. The diagnostic sensitivity of WLR imaging for early lung cancer is only about 9%.⁴ Preinvasive lesions such as carcinoma in situ (CIS) and moderate or severe dysplasia are only a few cell layers thick and may not show enough morphological changes or color variations detectable by a WLR image. As a scattering medium, tissue shows optical effects that are characteristic of turbid physical systems.⁵ Much of the light scattering in biological tissues is due to variations in polarizability, which can be characterized by variations in the index of refraction.

In addition, the shape, size, and distribution of tissue constituents play an important role in the overall scattering properties of tissue.⁶ The refractive index mismatch of biological tissue components as a highly scattering medium has a strong influence on optical transmittance and reflectance, as well as on polarization and coherence of the scattered light.⁷ Light-based therapeutic and diagnostic techniques could be improved if scattering within tissue were temporarily reduced. Laser surgery and therapy would benefit from the increased depth of penetration. Imaging modalities, such as fluorescence microscopy, confocal imaging, and optical coherence tomography (OCT), would also benefit from an increased penetration depth.⁶ It is well known that the turbidity of a dispersive physical system can be effectively controlled using an immersion effect (matching the refractive indices of scatterers and the base material) as well as by changes in the size of scatterers and their packing (the volumetric arrangement affects the spatial correlation of the scatterers).⁵ The optical (scattering) properties of living tissue can be controlled by using various physical and chemical reactions such as compression, stretching, dehydration, coagulation, ultraviolet (UV) irradiation, low-temperature application (reversible cold cataract), and addition of chemicals.^5,8–13

Several recent studies have demonstrated that sonophoretic (SP) delivery, as a noninvasive physical method, exhibited an enhanced tissue-clearing effect when applied topically with optical clearing agents (OCAs).^14–17 The significant ultrasound-induced enhancement in OCT imaging depth and contrast of in vitro porcine skin and in vivo human skin was found.^18,19 Several imaging techniques have been used to study the permeation of analytes in biological tissues,²⁰ including ultrasound,²¹ magnetic resonance imaging (MRI),²² optical projection tomography (OPT),²³ and OCT.^24–28 OCT is particularly attractive because it is a nondestructive, noninvasive technique that provides high-resolution, real-time images of biological tissues. OCT has been used extensively in tissue studies^24–28 to measure permeation of different molecules and materials through various epithelial tissues, such as rabbit sclera,²⁵ monkey skin,²⁶ rabbit cornea,²⁷ porcine aorta,^26,28 human skin,²⁹ and human esophagus.³⁰

Ultrasound-mediated bioeffects are generally believed to be caused by cavitation.³¹ Acoustic cavitation involves the creation and oscillation of gas bubbles in a liquid.³² Cavitation bubbles may exhibit sustained growth and oscillations over several acoustic cycles (stable cavitation) or violent growth and collapse in less than a cycle (transient or inertial cavitation).³² Potentially, both stable and transient cavitation may induce membrane permeabilization.³³ A large number of studies have reported on the SP delivery of small molecules through the skin in transdermal drug delivery (TDD).³⁴ It represents a useful noninvasive physical method to enhance epidermal permeability. In spite of the wide use of SP delivery in TDD, there are few studies to develop SP delivery for the effects of glucose permeability of normal, benign, and cancerous human lung tissues. The objectives of our study were to evaluate the effects of glucose mediated with ultrasound on permeability of normal, benign, and cancerous human lung tissue, to monitor and quantify the differences in permeability coefficients of ultrasound-OCAs combination in human normal lung (NL), lung benign granulomatosis (LBG), lung adenocarcinoma tumor (LAT), and lung squamous cell carcinoma (LSCC) tissue in vitro by the Fourier-domain optical coherence tomography (FD-OCT).

2.

Materials and Methods

3.

Results and Discussion

In this investigation, the main purpose of this pilot study was to evaluate the effects of ultrasound-mediated glucose on permeability of normal, benign, and cancerous human lung tissues, to monitor and quantify the differences in permeability coefficients of ultrasound-OCAs combination in four different kinds of lung tissue by the FD-OCT, and to find a more effective ultrasound-OCAs combination method to distinguish normal and diseased human lung tissues than pure OCAs.

The OCTSS of four kinds of lung tissue were calculated from the 80 μm region at a tissue depth of approximately 230 μm from the tissue surface. The OCTSS graphs were analyzed to obtain the permeability coefficients. First, the lung tissue was imaged for about 10 min to record the OCTSS base line as a control OCTSS with no glucose/SP added. During this time, the time series of OCTSS fluctuation was induced by the sample (NL, LBG, LAT, and LSCC tissue) scanning as shown in Figs. 1, 2, 3, and 4. After the 10 min period, 30% glucose/SP was topically applied to the OCT scanning area where imaging continued for another 2 h. Figure 1 displays a typical OCTSS graph as a function of time for NL tissue with the topical application of 30% glucose/SP. Diffusion of ultrasound-mediated glucose solution inside the lung tissues dynamically changed the local scattering coefficient and it was detected by the OCT system. The OCTSS decreased due to the reduction of scattering inside the tissue caused by the local increase of glucose concentration, and decrease of water content in the interstitial space due to osmotic action of glucose.^24,25,27 The increase in slope at the end of the graph is thought to be the reverse process. Water starts to diffuse into the tissue due to the difference in concentration created by the increase of glucose concentration.

Fig. 1

OCT signal slope as a function of time recorded from NL tissue treatment with 30% glucose and ultrasound.

Fig. 2

OCT signal slope as a function of time recorded from LBG tissue treatment with 30% glucose and ultrasound.

Fig. 3

OCT signal slope as a function of time recorded from LAT tissue treatment with 30% glucose and ultrasound.

Fig. 4

OCT signal slope as a function of time recorded from LSCC tissue treatment with 30% glucose and ultrasound.

In Fig. 1, the diffusion of glucose solution at NL tissue took approximately 52 min to reach an equilibrium state. A smoothed time-signal was obtained by adjacent averaging every five points. The mean permeability coefficient of 30% glucose/SP in NL tissue was measured in 12 independent experiments, which was estimated to be $(2.01\pm 0.21)\times {10}^{-5}\mathrm{cm}/\mathrm{s}$ cm/s. Similarly, Figs. 2, 3, and 4 show the OCTSS graphs as a function of time for LBG, LAT, and LSCC tissue with the topical application of 30% glucose/SP, respectively. The changing trends in the OCTSS of LBG, LAT, and LSCC tissue are also similar to that of NL tissues. In Figs. 2, 3, and 4, the diffusion of glucose solution at LBG, LAT, and LSCC tissue took approximately 40, 37, and 31 min to reach an equilibrium state, respectively. Their mean permeability coefficients of 30% glucose/SP were $(2.75\pm 0.28)\times {10}^{-5}$, $(4.53\pm 0.49)\times {10}^{-5}$, and $(5.81\pm 0.62)\times {10}^{-5}\mathrm{cm}/\mathrm{s}$, respectively.

Figure 5 represents the comparison of permeability coefficients for four different types of lung tissues with the topical application of 30% glucose/SP in vitro. The results shown in Fig. 5 illustrate that the permeability coefficients of 30% glucose/SP increase approximately 36.8%, 125.4%, and 189.1% for the LBG, LAT, and LSCC tissue compared with that for the NL tissue, respectively. There were statistically significant differences in permeability coefficients of 30% glucose/SP between LBG and NL tissue ($p<0.05$), between LAT and NL tissue ($p<0.05$), and between LSCC and NL tissue ($p<0.05$).

Fig. 5

Comparing permeability coefficients of 30% glucose/SP diffusion in NL, LBG, LAT, and LSCC tissue in vitro.

In our previous studies³⁵ the mean permeability coefficients have been reported for pure 30% glucose solution in human NL, LBG, LAT, and LSCC tissue in vitro. Comparison of the results of the current study with our previous studies³⁵ shows a significant increase in the permeability coefficients of 30% glucose/SP compared with that of pure 30% glucose in the NL, LBG, LAT, and LSCC tissue in vitro, respectively. Their increments are 48.9%, 54.5%, 57.3%, and 64.6%, respectively. This phenomenon is consistent with our previous findings that ultrasound caused the increase in permeability of normal breast and breast cancer tissue¹⁷ and that ultrasound caused the increase of cell membrane permeability in TDD.^39–42 The studies have demonstrated the ultrasound-induced tissue optical clearing enhancement with topical application of 30% glucose on in vitro normal, benign, and malignant human lung. Moreover, their optical clearing effects are significantly different. In the last decade, many studies analyzed the relationship between tissue composition, microstructure, and macrophysiology, showing that the lung’s physiological behavior reflects both the mechanical properties of tissue individual components and its complex structural organization.

Different lung pathologies can affect the extracellular matrix (ECM).⁴³ For example, the increased deposition of ECM proteins and fibroblasts in the stroma surrounding the epithelial cells of the breast is the leading risk factor for breast carcinoma (which accounts for 30% of breast cancer).⁴⁴ This and many other diseases might potentially be diagnosed by comparing the permeability rates of several therapeutic or diagnostic agents in normal and abnormal tissues.²⁶ Therefore, precise assessment of the diffusion processes in normal and abnormal tissues with OCT might provide a truly noninvasive way for evaluation of tissue health.²⁸ As yet, however, no published reports have demonstrated the capability of the OCT technique for quantifying the permeability coefficients of the ultrasound-OCAs combination in normal, benign, and malignant human lung tissue. Our experiment is the first reported investigation of ultrasound-mediated glucose diffusion in normal, benign, and malignant human lung tissue.

In our pilot experiments, the ultrasound was selected to enhance the glucose diffusion due to several reasons. First, the technique is noninvasive. Second, no insertion or surgery is needed, and acoustic transducers are placed in contact with the mixed solution that is spread on the tissue. Third, the method is extremely simple to implement (sonicators have been used in beauty parlors and clinics). Additionally, the cost is low. Ultrasound has demonstrated an enhanced transdermal mass transport effect when applied topically in TDD. Its mechanism is believed to be through acoustic cavitation, which is the formation and collapse of gaseous cavities, and has the dominant role in sonophoresis.^45,46 Based on the aforementioned results, it indicated that due to the ultrasound penetration enhancing and convection effect, more glucose can penetrate the lung tissue and glucose can penetrate more quickly to achieve more refractive index matching in human lung tissue in vitro to improve lung tissue optical clearing.

4.

Conclusions

Our results demonstrate that simultaneous application of ultrasound and glucose led to a marked increase in the permeability coefficients for four different types of lung tissues in vitro. The permeability coefficients of 30% glucose/SP increase approximately 36.8%, 125.4%, and 189.1% for the LBG, LAT, and LSCC tissue compared with that for the NL tissue, respectively. The coadministration of OCAs diffusion under ultrasound-assistance is a more effective approach in using OCT imaging to diagnose normal, benign, and cancerous human lung tissues than that of pure OCAs.³⁵