1.

Introduction

Turbid materials and biological tissues scatter light strongly. For this reason, the imaging depth of different optical techniques is quite limited in such media. Light scattering properties are determined by the size, position, and shape of the scattering particles in relation to the wavelength and direction of the probing light. The major cause of scattering is the refractive index mismatch between the scattering particles and the background medium. In tissue, these components(particles) are mitochondria, cytoplasms, cell membranes, extracellular media, and collagen and elastin fibers.^{1, 2, 3} Seventy percent of the dry weight of skin consists of collagen fibers with a high refractive index, and hence the main scatterers are collagen fibers.^{2, 4}

Optical clearing by means of immersion technology is a powerful tool for controlling the optical properties of tissues. It is based on matching of the refractive indices of the bulk material and scattering particles. Several clearing agents with different refractive indices have been used to decrease scattering, enhance light penetration depth, increase image contrast, and overcome the limitations of strong optical scattering of biological samples.⁵

A theoretical model has been developed to describe solvent penetration by diffusion and the matching of refractive indices in fibrous tissue. Permeation of the agents alters the size of scattering centers by contracting or swelling, which controls the rate of optical clearing.⁴

The optical clearing phenomenon also forms the basis for optical scattering-based measurements of different substances. In these experiments, changes in the concentration of a substance result in changes in scattering, and they are further seen in the optical signals. This theoretical background has been studied for noninvasive glucose monitoring.^{6, 7} Increasing the glucose concentration decreases scattering and increases light penetration depth. In recent years, several optical methods, such as spatially resolved diffuse reflectance measurements and optical coherence tomography (OCT), have been used to study these changes in tissue-simulating phantoms and in tissues both in vitro and in vivo. ^{8, 9, 10, 11} However, physiological glucose levels are in the range of 3 to $30\mathrm{mM}$ , which are far below the level of making the tissue transparent.

OCT is a high-resolution imaging technique based on an interferometer. After the invention of OCT in 1991,¹² it has been applied to several imaging applications.^{13, 14} OCT is capable of detecting photons backscattered from different layers in the sample. Thus, changes in the scattering properties can be studied at different depths in the sample.

Investigation of the changes in a sample’s scattering properties with OCT is based on analyzing the depth profile of an attenuating interference signal. Changing optical properties, i.e., in the case of adding glucose to the sample, induce changes in the slope value of a line fitted to the depth profile. The changes in the slope value can be thought to originate from changes in scattering in the following conditions: the scattering coefficient $\left({\mu }_s\right)$ is much higher than the absorption coefficient $\left({\mu }_a\right)$ , and the slope-fitting regime is in the single-scattering part of the profile. Deeper in the sample, especially in a highly scattering medium, a more complicated model is preferred for extracting the value of ${\mu }_s$ .¹⁵

Optical scattering-based techniques, such as OCT, are based on the interaction of light with the object under study. The effect of glucose is measured through changes in the refractive index mismatch between scattering particles and the base medium, and hence in light scattering properties. Due to the indirect nature of the measurements, OCT is not specific for single glucose molecules, and hence all other effects that may change the scattering properties should be tightly controlled.¹⁶ It is easier to study the effects of different factors on the experiments in in vitro conditions. This is the reason for developing different tissue-simulating phantoms and complicated tissue models.¹⁷ Compared with in vitro experiments, in vivo measurements are confronted with different problems. These effects have been studied, e.g., by Larin, ¹⁶ and they include the effect of variation in the concentration of different tissue components, the physiological response of tissues, tissue heterogeneity, motion artifacts, and temperature.¹⁸ Recent progress in developing an OCT-based glucose sensor has been made in improving accuracy and in finding out other possible effects than glucose that would change the OCT signal’s slope value.^{19, 20, 21} Quite recently, studies concerning possible enhancements of OCT to increase selectivity for single molecules have been published.²²

This work concentrates on the differences in the effect of glucose on the optical properties of different materials (Intralipid and skin). The differences between in vitro and in vivo experimental conditions are also considered.

2.

Materials and Methods

2.2.

Optical Coherence Tomography Device

A commercial OCT system (manufactured by the Institute of Applied Physics, Nizhny Novgorod, Russia) was used in the studies. It uses a wavelength of $910\mathrm{nm}$ with resolutions of $<10\mathrm{\mu }\mathrm{m}$ for axial and lateral resolution in air. The spectrum width is $49\mathrm{nm}$ . The Doppler frequency of the OCT device is $1415\mathrm{kHz}$ . The OCT device was controlled with a PC via a USB channel. The signals were measured with an optical probe and saved on the computer. A scanner inside the optical probe scanned along the sample surface. The adjustable scanning range along the $x$ and $y$ axes was between 0 and $2.2\mathrm{mm}$ . The inner diameter of the optical fiber tip was $6\mathrm{\mu }\mathrm{m}$ , the diameter of the cladding was $125\pm 1\mathrm{\mu }\mathrm{m}$ , and the outside diameter was $245\mathrm{\mu }\mathrm{m}\pm 5\%$ . The size of the focus point was $8.4\mathrm{\mu }\mathrm{m}$ when in the center area of the focal adjustment range of the probe, and with the magnification of the optics at 1.4. The numerical aperture of the optical fiber was in the range of 0.13 to 0.15. The lateral scanning range and the diameter of the probing beam increased proportionally to the magnification of the output lens.²⁹ The image acquisition speed of the OCT device was 1 image/s for an image size of $100\times 400$ pixels (100 in the lateral and 400 in the depth direction). The setup used in the experiments is shown in Fig. 1 .

Fig. 1

Setup used in the experiments. (a) Image from the OCT device and probe, (b) schematic diagram of the setup used in the in vitro experiments, and (c) schematic diagram of the setup used in the in vivo experiments. (b) and (c) are from Refs. 28, 30, respectively.

3.

Results and Discussion

The three samples used in the studies, Intralipid, skin tissue samples, and mice in vivo, have different optical properties and physical structures. Figures 2a and 2b show a typical OCT image and a 1-D depth profile from Intralipid, respectively. Figures 2c and 2d show a typical OCT image and a 1-D depth profile from the skin tissue sample, respectively. Figures 2e and 2f show a typical OCT image and a 1-D depth profile from the mouse skin in vivo. Physical depths were calculated by dividing the optical depth values with the average refractive index of the sample. A refractive index value of 1.34 was used for 2% Intralipid in Figs. 2a and 2b, a refractive index value of 1.4 was used for the skin tissue sample in Figs. 2c and 2d, and for mouse skin in Figs. 2e and 2f.^{22, 23}

Fig. 2

(a) OCT image and (b) averaged depth profile from 2% Intralipid. (c) OCT image and (d) averaged depth profile from the skin tissue sample: and (e) OCT image and (f) averaged depth profile from mouse skin tissue in vivo.

Figures 2a and 2b show that Intralipid is quite a homogeneous and turbid material, whereas the refractive index distribution has more variation in the studied tissue samples [Figs. 2c and 2d]. The in vivo results show the highest backscattering intensity when measured with OCT, but also rapid attenuation as a function of depth. The peaks at the beginning of the images (left side) result from reflection from the sample’s surface.

Different concentrations of glucose were added to the samples to study possible differences in matching of the refractive indices in different samples, and hence changes in the optical scattering properties. Figure 3 shows focused views of the fitted lines on the OCT signal depth profiles in different samples with two different glucose concentrations. Physical depths were calculated by dividing the optical depth values with the average refractive index of the sample. A refractive index value of 1.34 was used for 2% Intralipid in Fig. 3a, and a refractive index value of 1.4 was used for the mouse skin tissue sample in Fig. 3b and for mouse skin in Fig. 3c.^{2, 23}

Fig. 3

Fitting of a straight line on the OCT signal profiles at two different glucose concentrations from (a) 2% Intralipid, (b) a skin tissue sample, and (c) mouse skin in vivo.

In the Intralipid samples, the slope values were determined by fitting the slope on the most linear part of the attenuation curve after the sample surface [Fig. 3a]. This was also true for the skin tissue samples in vitro [Fig. 3b].²⁸ The effect was smallest in the Intralipid samples, whereas the matching effect was larger in the real tissue. This indicates that optical clearing with glucose would also be more obvious in vivo. In the in vivo measurements, the slope values were determined at the place where the glucose-induced change in the slope value was at a maximum [Fig. 3c]. This was found to be in the dermis part of the skin, where the network of small capillaries is located. The value of $2.7\%∕\mathrm{mM}$ is an average of four measurements: 1.6, 1.077, 3.36, and 4.85%, which were recently published in Ref. 30. The correlation coefficients of those measurements were 0.24, 0.75, 0.86, and 0.60, respectively. The very pure value of 0.24 was due to some unexpected changes in the slope values at the end of the experiment.

Table 1 shows the average glucose-induced changes in the OCT signal slope value obtained from the different samples. The in vitro tissue experiments consisted of measurements of five different samples. Possible reasons for the different effect of glucose could be the different refractive indices of the scattering centers in Intralipid and in skin tissue. The refractive index of soybean particles is $\approx 1.463$ , and of water $\approx 1.33$ . In tissue samples, $n\approx 1.40$ for mitochondria, 1.46 for collagen fibers, and $\approx 1.362$ for extracellular fluid. Assuming seventy percent of the dry weight of skin consists of collagen fibers and approximating the rest to be formed by other components with $n=1.40$ , it is possible to get the value of $n=0.7\times 1.46+0.3\times 1.4=1.442$ for skin tissue. For the nutrition medium, $n$ was supposed to be the same as for water. ^{2, 23, 31, 32} This gives the refractive index difference between the bulk medium and scattering particles of $\mathrm{\Delta }n\approx 1.442-1.362=0.08$ for skin tissue in vivo, $\mathrm{\Delta }n\approx 1.442-1.33=0.112$ for skin samples in vitro, and $\mathrm{\Delta }n\approx 1.463-1.33=0.133$ for 2% Intralipid. Possible errors may be induced into the calculations when assuming the refractive index of the nutrition medium to be 1.33 instead of a combination of the refractive indices of the nutrition medium and intracellular fluid. Moreover, alteration of the sizes of the scattering centers and cells by contracting or swelling has to be taken into account in the skin experiments.⁴

Table 1

Average glucose-induced changes in the OCT signal slope value in Intralipid and skin samples in vitro, and mouse skin in vivo.

The skin samples and the skin in vivo have a layered structure. The main layers in the mouse skin are the epidermis and dermis, with thicknesses of a few dozen micrometers and between 170 to $500\mathrm{\mu }\mathrm{m}$ , respectively,^{33, 34} in vivo. They differ largely from the Intralipid suspensions. Although Intralipid quite closely imitates the properties of skin in transmittance and reflectance measurements,²⁶ there are no different layers observed in the measured OCT images. In Intralipid, the scattering particles have an almost spherical shape, whereas there are many kinds of components with varying shapes in skin (e.g., collagen fibers, cell nucleus, mitochondria, etc.). The results also show that backscattering of photons is much stronger in real tissue in vivo than in vitro. One explanation is the lack of blood in the tissue samples, but in the in vivo experiments, blood is present.

The basic setup for glucose injection is different in the in vitro and in vivo experiments. In the in vitro case, glucose was added to the dish and the solution was mixed a little. In these experiments, the glucose was thought to be distributed smoothly in the sample after a time period of 15 to $30\min$ after glucose addition. On the contrary, in the in vivo experiments, the glucose was diffused into the tissue from blood. Thus, the original place where the glucose load penetrated the sample was different. Both methods were different from those used in the experiments conducted by Wang and Tuchin,³ where intradermal injection was used. Intradermal injection was used because the epidermis has very slow permeability to glucose.³

The place of measurement has more effect on the measurements in the skin than in the Intralipid. This is due to the inhomogeneous structure of the skin sample surface. Measurement at different places on the rough sample surface leads to a large standard deviation in the measurements. However, the skin tissue samples stayed still during the measurements. The situation is different in the experiments in vivo, where the motion of the mice caused distortions during the 2-h measurement period. The movements took place especially in the cases where blood samples were taken from the feet of the mice to measure blood glucose control values, and when the effect of the anesthetic drug weakened and more of the drug was injected. The motion artifacts were corrected afterward.

The measurement conditions, such as temperature, humidity, and ${\mathrm{CO}}_2$ content, also affect the results. The mouse skin samples were measured in a warm room $(+37°\mathrm{C})$ . This diminished the effect of changing temperature during the experiments. The measurement time in the in vitro measurements was minimized to only about ten seconds for one image to diminish the effect of decreasing ${\mathrm{CO}}_2$ content while the samples were outside the incubator. Because the optical system of the measurement probe has a curved focus plane, the width of the scanned image was decreased from the maximum to avoid harmful effects on the OCT signal depth profile and to increase the accuracy of the experiments.

A comparison of the glucose-induced changes in the Intralipid and other in vitro measurements (wavelength is $1310\mathrm{nm}$ ) shows quite good comparability. The glucose-induced changes vary in the range of 0.023%/mM for polystyrene microspheres in vitro ⁹ to $2.5\%∕\mathrm{mM}$ in Yucatan Micropigs,³⁵ $3.3\%∕\mathrm{mM}$ for rabbit ear,¹⁶ and 3.42%/mM for human skin¹¹ in ivo. The in vivo results from mice are in the same range as the other in vivo experiments. The glucose values in vitro from the skin tissue samples are in the middle of these two categories.

4.

Conclusions

The structures of the studied samples, Intralipid, fetal skin tissue samples, and mouse skin tissue in vivo, differ from each other because Intralipid has spherical scattering particles with varying size and homogeneous particle distribution, while skin tissue consists of several components with different diameters and shapes and a layered structure. OCT signals from the studied samples show that scattering in the backward direction is stronger from skin tissue in vivo than from skin samples in vitro. Glucose is found to have the most pronounced effect in in vivo conditions, and the smallest effect in 2% Intralipid. This can partly be explained by the refractive index differences between the bulk medium and scattering particles ( $\mathrm{\Delta }$ n $\approx$ 0.08 for skin tissue in vivo, $\mathrm{\Delta }n\approx 0.112$ for skin samples in vitro, and $\mathrm{\Delta }n\approx 0.133$ for Intralipid).

These results confirm that a commercial OCT device, designed for imaging purposes, is capable of detecting glucose-induced changes in scattering in different types of samples in very well-controlled conditions. Moreover, cell cultured tissue samples seem to be a viable, although not most stable, option when developing optical sensing techniques from a phantom base toward in vivo conditions. However, if more parameters change simultaneously, this analysis method is not accurate enough. In that case, quantification of the glucose concentration cannot be made from one measurement. Thus, simultaneous measurements with different techniques or improved glucose selectivity with OCT are needed.