1.

Introduction

Optical coherence tomography¹ (OCT) is a high-resolution imaging modality appropriate for assessment of skin in vivo. Structures such as the stratum corneum, epidermis, dermo-epidermal junction, blood vessels, and skin appendages, such as hair follicles, sweat ducts, and sebaceous glands, have been identified using OCT and good correlations with histology have been published.^{2, 3} OCT has been shown to provide an accurate measure of epidermal thickness validated against cryostat histology,⁴ and studies have used OCT to monitor epidermal thickness changes due to UV radiation,^{5, 6} aging,⁷ and skin diseases.⁸ OCT has also been applied to real-time assessment of skin inflammatory disease^{8, 9, 10, 11, 12, 13} and tumors,^{14, 15, 16} the examination of skin vasculature dilation,¹⁷ and the efficacy of laser treatment of such vascular malformations as port-wine stain birthmarks.³

OCT imaging of skin can suffer from significant image artifacts. We consider an image artifact to be any image feature not indicative of a structure in the skin. Image artifacts may present as extraneous, non-existent visual features, or distortions (in shape or intensity). The presence of artifacts produces an inaccurate representation of the underlying skin morphology and complicates the clinical interpretation of the images. Several types of instrument- and sample-induced artifacts have been reported to date for OCT. Instrument-induced artifacts include those due to conjugate image features (or mirror artifact),¹⁸ optical aberration (or fan distortions),¹⁹ and autocorrelation.²⁰ Sample-induced artifacts include those due to horizontal banding (or birefringence artifact),²¹ shadowing,²² refraction,²³ and motion.²⁴ In this study, we have classified two prominent types of a sample-induced artifact in OCT images of skin, which we shall refer to as intensity distortion and morphological distortion. These artifacts are present particularly when the system is used in noncontact mode (i.e., when an air-gap exists between the skin and the focusing optics of the OCT scanner).

Intensity distortions appear as vertical streaks of low-intensity backscatter under skin furrows, which are grooves of variable depth on the skin surface (Fig. 1). The low-intensity streaks are observed to correlate with the location of the surface furrows and broaden with depth. We expect that they are associated with refraction and divergence of the light beam caused by the skin's furrowed surface and the associated abrupt change in refractive index (RI) between the skin and air. These low-intensity streaks cause an imprint of the skin surface topography to appear at all depths of the OCT scan. In en face images (i.e., lateral cross-sectional images) under the skin surface, this imprint creates extraneous features, presenting as lines aligned to the surface furrows.

Fig. 1

OCT images of the skin of the medial calf of a human subject illustrating intensity distortion artifact: (a) 3D solid rendering of the skin showing skin surface topography. The labeled cross-sectional images are shown in (b) and (c); (b) En face image (CS1) at 400-μm depth (uncorrected for RI) below the surface showing evidence of artifactual lines replicating skin surface topography; (c) vertical cross-sectional images (CS2) highlighting the source of the artifact: low-intensity, vertical streaks (indicated by arrows) which emerge at multiple locations directly under the skin furrows. Vertical scale bar in (c) is not corrected for RI.

Two types of morphological distortion are observed. One type occurs because OCT systems generate images based on optical path length, instead of the true physical path length. As optical path length is a function of the RI of the medium, different amounts of air and tissue traversed in reaching a feature will change the optical path length to the feature. Thus, the surface shape and thickness of the overlying tissue can modify the apparent depth of a feature. In the other type of distortion, features are displaced laterally because the surface shape and refraction affect the angle at which the OCT light beam propagates beneath the tissue surface.

Such artifacts have been observed in images in previously published papers, but have not been specifically studied.^{8, 10, 25, 26, 27, 28} Some researchers have used glycerol and ultrasound gel to reduce RI mismatch and to improve image quality. ^{2, 5, 9, 11, 29, 30, 31} The efficacy of these media in reducing image artifacts has not previously been assessed. In addition, the application of glycerol or ultrasound gel is known to affect the thickness of the skin but such effects on the epidermis have not previously been examined in detail.

In this study, we investigate and assess the efficacy of glycerol and ultrasound gel in reducing OCT image artifacts. Experiments were carried out on both novel skin mimicking-phantoms and on in vivo human skin. We also assessed the effect of these media on epidermal thickness, a common parameter in the assessment of skin. Based on our findings, recommendations are given for optimal OCT imaging of skin in vivo.

2.

Methods

2.1.

OCT System

A fiber-based, swept-source frequency-domain OCT system (OCS1300SS, Thorlabs, USA), comprising of a broadband, swept laser source and a Michelson interferometer with balanced detection, was used in this study. The laser source has a central wavelength of 1325 nm and a full-width at half-maximum bandwidth of 100 nm. The average optical power of the probing beam incident on the sample was 4.7 mW and the axial A-scan frequency was 16 kHz. The system has a manufacturer-specified axial resolution of 12 μm in free space corresponding to ∼8.4 μm in skin (assuming a skin RI of 1.43, Ref. 27). The B-scan (set of 800 A-scans) frequency was 15.4 frames/s and the three-dimensional (3D) acquisition time was 55 s. All images acquired were 4 × 4 × 3 mm (x × y × z) in size and consisted of 800×800×512 pixels.

The sample arm was configured as a handheld probe. An adjustable spacer was affixed to the probe (Fig. 2) to maintain a constant distance between the objective lens and the sample. A glass cover slip was attached to the bottom of the spacer to act as a window for imaging. The working distance and numerical aperture of the objective lens is 25.1 mm and 0.056, respectively, yielding a specified diffraction-limited lateral system resolution of 15 μm.

Fig. 2

OCT handheld probe assembly and imaging setup.

3.

Results

3.1.

Skin phantom study

Figure 3 illustrates the effects of glycerol and ultrasound gel on intensity artifacts in the intensity phantom. The dark lines visible below the skin surface in the en face images [Rows 1 to 3 of Fig. 3a] replicate the topography of the surface furrows. These lines arise from dark vertical streaks in the B-scans, which emerge under skin furrows, as shown in the last row. When glycerol or ultrasound gel is used, these streaks are reduced in magnitude and commence at a greater depth.

Fig. 3

OCT images of the intensity phantom designed to demonstrate intensity artifacts. Images obtained with: (a) air, (b) glycerol, and (c) ultrasound gel interfaces. The top three rows show en face images of the phantom at depths (uncorrected for RI) of 0.5, 1.0, and 1.5 mm from the surface, respectively; the bottom row shows typical vertical cross-sectional images of the phantom. The arrows highlight the same aggregation of TiO₂ scatterers in the phantom; vertical scale bar is not corrected for RI.

OCT images of the morphology phantom, acquired with the three different media, are shown in Fig. 4. For each medium, a set of 3D rendered images and two-dimensional cross-sections are shown. Morphological artifacts are evident under skin furrows, but are absent in regions under flat surfaces. Overall, the artifacts are much more pronounced for the case of an air interface. The bottom surface of the phantom was exposed to air during imaging. This resulted in full-depth vertical white streaks caused by strong bottom surface reflections at some locations evident in the B-scan images in Fig. 4, but these did not significantly affect the analysis.

Fig. 4

OCT images of the morphological phantom containing fibers designed to demonstrate morphological artifacts. For each of (a), (b), and (c): Upper left: 3D solid rendering of the phantom; Lower left: orthogonal cut-away views of the phantom; Middle: en face image of the phantom at the level through horizontal midline of the fibers’ cross-sections; Upper and lower right: vertical cross-sectional images CS1 and CS2, respectively, as indicated in en face image. Vertical scale bar is not corrected for RI.

Figure 4a demonstrates morphological distortion of the regularly spaced array of single-mode fibers embedded within the phantom for the case of an air interface. Visually, the circular cross sections of the optical fibers are broadened, thinned, or almost undetectable (labeled with solid arrows in CS1). These distortions are indicated by a lower Dice coefficient of 0.85 compared to that of the control 0.95 (p < 0.05) (Table 1, medium: air). In addition, the long axis (or transverse cross sections) of the fibers appeared furrowed (or bent) and truncated at some points (labeled with a solid arrow in CS2). This furrowed appearance of the fiber's long axis mirrors the shape of surface furrows, but on a reduced scale (due to the relative change in RI and ratios of air and phantom in the beam path). Similar distortions were also observed to affect the base surface of the phantom (labeled with dotted arrows in CS1 and CS2). The intensity artifacts seen in Fig. 3 could also be observed under the surface furrows in Fig. 4a.

Table 1

Quantification results of image artifacts and epidermal thickness.

Parameter	Quantification metric and statistical measure	Medium	Results
Morphological artifact	Dice coefficient: Median (IQR)1	Control	0.96 (0.95 to 0.96)
		Air	0.85 (0.75 to 0.90)
		Glycerol	0.95 (0.94 to 0.96)
		Ultrasound gel	0.95 (0.94 to 0.96)
Intensity artifact	Power ratio: Median (IQR)	Control	1.00 (0.92 to 1.09)
		Air	0.44 (0.33 to 0.59)
		Glycerol	1.02 (0.91 to 1.16)
		Ultrasound gel	1.01 (0.91 to 1.13)
Epidermal thickness	Percent changes,%: Mean (± standard error)	Air	5 min : 0
			20 min : 0
		Glycerol	5 min : 8 (± 0.5)
			20 min : 11 (± 0.9)
		Ultrasound gel	5 min : 9 (± 1.0)
			20 min : 14 (± 0.9)

IQR – interquartile range.

The application of RI-matching media is seen to effectively reduce these artifacts [Figs. 4b and 4c]. Under the surface furrows, both the cross sections and long axes of the fibers, as well as the bottom surface of the phantom, appear much less distorted [labeled by solid and dotted arrows in Figs. 4b and 4c]. The results in Table 1 show that use of either glycerol or ultrasound gel has decreased these morphological distortions by significantly increasing the Dice coefficient to 0.95 (p < 0.05), which is comparable to that of the control fibers (i.e., 0.96).

The mean group RI of the phantom, glycerol, and ultrasound gel was determined to be 1.41±0.02, 1.47±0.01, and 1.34±0.01, respectively. The RI of glycerol is higher than that of the phantom and the RI of ultrasound gel is lower, although both media have comparable mismatches (0.06 and 0.07, respectively) with the phantom. Glycerol and ultrasound gel showed subtle differences in reducing the intensity artifacts but statistically the same (p > 0.05) efficacy in minimizing the morphological artifacts. We observed that the intensity streaks were seen to commence at a greater depth by approximately 10% when using glycerol than when using ultrasound gel.

3.2.

Human Subject Study

For in vivo scans of human skin, a reduction was observed in the level of both intensity and morphological artifacts after the application of glycerol or ultrasound gel. Figure 5 shows the en face images of human ventral forearm skin at 0.5 and 1 mm below the skin's surface. As with the phantom data (Fig. 4), the onset of the intensity streaks was seen to occur at a greater depth after application of the media. These streaks, however, are less pronounced in the skin after RI matching. Note that air bubbles trapped within the RI-matching medium (labeled with arrows) can significantly attenuate the image intensity beneath the bubble.

Fig. 5

En face OCT images of human ventral forearm skin at 0.5 mm (top row) and 1 mm (bottom row) (uncorrected for RI) below the skin's surface, obtained with: (a) air; (b) glycerol; and (c) ultrasound gel interfaces. Arrows highlight shadow artifacts caused by trapped air bubbles within the RI-matching medium. H: Hair follicle.

Figure 6 shows two examples of artifact-containing B-scans of human ventral forearm skin containing blood vessels (labeled B) and hair follicles (labeled H) under the three interface conditions. In each example, images are approximately collocated. RI matching has reduced the intensity of surface reflections, and increased the useful penetration depth (from ∼1 mm with air, to ∼1.5 mm with glycerol or ultrasound gel). In the image obtained with an air interface, dark streaks (labeled with dotted arrows) partially obscure the blood vessels and hair follicles. The blood vessels are also distorted in shape, i.e., truncated and bent (labeled with solid arrows), corresponding to the morphological artifact described for the phantom. Both these effects are reduced after application of either medium.

Fig. 6

Two examples of OCT vertical cross-sectional images (B-scans) of human ventral forearm skin. Images were acquired using: (a) air; (b) glycerol; and (c) ultrasound gel interfaces. B: blood vessels; H: hair follicles; vertical scale bar is not corrected for RI.

Quantitative analysis (Table 1) shows that, with an air interface, the region under a skin furrow contained only 0.44 of the signal power present in the adjacent artifact-free region. Glycerol and ultrasound were found to minimize intensity artifacts, with power ratio of 1.02 and 1.01, respectively. This is comparable to the power ratios measured in the control regions of 1.0. The median power ratio is not expected to be exactly 1 as it is calculated from A-scans in different regions, albeit adjacent regions, of the skin. Both glycerol and ultrasound gel show comparable reduction in image artifacts and the difference between the two are not significant (p > 0.05).

Figure 7 shows the average percentage thickening of epidermis versus time for 10 subjects after the application of either medium. Both media cause an increase in the epidermal thickness by 5% to 6% within 30 s of application. The increase in epidermal thickness plateaued at ∼20 min and remained unchanged after 1 h. At 20 min, both glycerol and ultrasound gel had significantly thickened the epidermis, by 14% and 11%, respectively.

Fig. 7

Average percentage thickening of epidermis after the application of glycerol or ultrasound gel. Error bars represent standard error.

Table 1 summarizes the quantitative results on image artifacts and epidermal thickness.

4.

Discussion

The results presented here demonstrate the severity of intensity and morphological artifacts in OCT images of skin. Although we did not have access to gold standard histology, a priori knowledge of skin structures, combined with comparison of images recorded using different media and quantification has enabled us to draw a number of useful conclusions discussed below. These artifacts appear most prominently under skin furrows and are a function of the marked change in RI between tissue and air. The use of glycerol or ultrasound gel was shown to reduce these artifacts by reducing the RI-mismatch at the air-skin interface. The qualitative and quantitative assessment presented here has shown no significant difference, with regard to these artifacts, in the use of either medium for in vivo human skin imaging.

Both types of image artifacts in OCT have been shown to be detrimental to the correct interpretation and quantification of subsurface structures of the skin. The shape of blood vessels can be obscured by the dark streaks and can appear distorted. It seems likely that smaller features may be completely occluded. Support for this assertion is contained in Fig. 8, which shows a 3D projection image of a skin phantom containing a large collection of small scattering objects (aggregates of titanium dioxide particles). Many of these objects are undetectable in the air-interfaced phantom, but are visible after the application of either medium.

Fig. 8

Three-dimensional projection of a 3D clear render (i.e., background of the scattering structures was made transparent) of the intensity-distortion phantom images shown previously in Fig. 3 acquired with: (a) air; (b) glycerol; and (c) ultrasound gel interfaces. Arrows indicate examples of small scatterers (aggregates of titanium dioxide particles) occluded in (a); vertical scale bar is not corrected for RI.

The low-numerical aperture focusing optics in the sample arm limits the lateral extent of the illumination and collection of light, thus, discriminating in favor of the detection of backscattered/reflected light from the sample focal volume and suppressing signal outside this volume.^{1, 37} We hypothesize that the uneven surface of the skin, especially the skin furrows, causes rapid divergence or deflection of the incident beam. Since at the skin surface the furrow dimensions typically exceed those of the beam, the beam location in relation to the furrow determines the relative importance of divergence or deflection. The net effect, though, of either effect is a reduction in the signal intensity in the region beneath the skin furrows (i.e., an intensity artifact presenting as a low-intensity streak). In the in vivo human skin images shown in Fig. 5, RI matching is shown to reduce the signal loss and to delay the onset of these intensity artifacts to ∼1 mm in physical distance below the surface.

OCT produces an image based on the assumption that light has propagated along a straight beam path. However, when the light beam is deflected off-axis by the furrowed skin surface, the result is displacement of all or part of a feature laterally. The morphological distortions shown in Fig. 4a illustrate this. The broadening or thinning of the irregular and variable cross sections of the single-mode fibers embedded under the skin-textured portion of the morphology phantom is largely absent for the fibers under the flat portion.

OCT generates images in which the depth coordinate is based on the optical path length, not the physical path length. The morphological distortions caused by optical path length changes are illustrated in Fig. 9, which shows OCT vertical cross-sectional images of a phantom similar to the phantom that produced the images shown in Fig. 4, except that its surface is flat. The changes in optical path length cause features to appear stretched, as shown by the elongation (in height) of the circular cross sections of the fibers and the phantom (Fig. 9, top row). Changes in the optical path length occurring when the proportions of the sample and air vary can also cause displacement and truncation of a feature vertically relative to the adjacent sections. This is demonstrated in Fig. 9b (top row), in which the positions of adjacent fibers are shifted vertically, and the fiber's cross section below the edge is truncated (as labeled) due to a difference in the thickness of the layer above these fibers and the difference in RI between the phantom and air. Similar distortion can be observed for in vivo human skin in Fig. 6a, in which an abrupt change in the thickness of the epidermis has caused displacement in the image of portions of a blood vessel.

Fig. 9

Morphological distortion caused by optical path length. (a) Top: A silicone phantom embedded with an array of single-mode fibers; Bottom: As above, with vertical scaling by the RI of the phantom. (b) Top: The same phantom as in (a) overlaid with a thin additional layer of similar content (dotted box); Bottom: As above, with vertical scaling by the RI of the phantom. Vertical scale bar for the top row represents uncorrected distance, whereas the bottom row represents corrected distance.

Correction of such morphological distortions caused exclusively by changes in the optical path length can be performed by scaling or shifting the pixel location of the sample as a function of sample's RI, assuming it is known. In practice, this simply requires scaling the path length for each tissue type (or air) by the RI, as we have done in Fig. 9 (bottom row). This procedure accounts for differences between the optical and physical path length. However, it does not account for feature displacement in 3D arising from beam deflection caused by the irregular shape of skin-furrows.

Improvement in penetration depth/transmittance of skin by both glycerol and ultrasound gel were observed (Fig. 6) in this study. This is consistent with previous reports on the use of optical immersion agents such as glycerol, glycerol-water solutions, and ultrasound gel.^{38, 39, 40, 41} Several mechanisms have been proposed to explain this optical clearing effect. These include reduction in optical scattering due to matching of refractive indices of tissue scatterers (keratinocyte components in the epidermis, and collagen and elastin fibers in the dermis) with interstitial fluids;^{42, 43} dehydration of excised tissue samples;^{43, 44, 45} and dissociation⁴⁶ and structural modifications⁴⁰ of collagen fibers.

While this increased image penetration depth observed with glycerol and ultrasound gel is positive, their use was also observed to significantly thicken the epidermis (Fig. 7). This is undesirable in situations where measurement of the epidermal thickness is of clinical significance. The thickening (5% to 6%) of the skin we observed in the first 30 s after application of the medium is most likely to be due to optical changes rather than a physiological effect, as the diffusion of a medium through skin normally takes minutes³⁹ or hours.⁴² It is reasonable to expect that within the first minute, the stratum corneum, which is composed of dead and loosely connected keratinocytes, is likely to be permeated with the glycerol/ultrasound gel, replacing trapped air with the agent, and increasing the optical path length through the epidermis. We have confirmed this in a separate experiment (results not shown) using tape stripping to remove the stratum corneum prior to the application of the media. In this case, rapid increase in epidermal thickness is absent, although the same value is attained after 1 h. Tape stripping, however, is not feasible in many clinical measurement scenarios of interest.

As glycerol is a hydrophilic humectant (i.e., a water-soluble material with high water absorption capability),⁴⁷ subsequent swelling of the epidermis could be due to the attraction of water from the air or from deeper layers of the skin. It may also be due to the absorption of glycerol by the skin. Genina ⁴⁴ have shown that water loss and osmotic agent diffusion into the skin can take place simultaneously through intact stratum corneum albeit at a slower rate than in perforated skin. Ultrasound gel, on the other hand, contains a water-soluble polymer that may be absorbed by the skin.

Since RI-matching media reduce image artifacts, we recommend their use in an OCT skin-imaging protocol when imaging subsurface structures such as microvasculature and hair follicles. Epidermal thickness assessment should be performed without any RI-matching medium. As the efficacy of glycerol and ultrasound gel is comparable in reducing image artifacts, ultrasound gel is preferred for clinical applications as its higher viscosity makes its use more practical, although it is more susceptible to trapped air bubbles.

5.

Conclusion

We have shown that the RI-matching media glycerol and ultrasound gel are effective in reducing intensity and morphological artifacts in OCT images of skin-mimicking phantoms and in vivo human skin. Both media produced comparable reductions in image artifacts and both led to thickening of the epidermis. Thus, we recommend performing OCT scanning of skin in vivo with application of an RI-matching medium for subsurface structure assessment, and without for epidermal thickness assessment.