Spatially confined quantification of bilirubin concentrations by spectroscopic visible-light optical coherence tomography

: Spatially confined measurements of bilirubin in tissue can be of great value for noninvasive bilirubin estimations during neonatal jaundice, as well as our understanding of the physiology behind bilirubin extravasation. This work shows the potential of spectroscopic visible-light optical coherence tomography (sOCT) for this purpose. At the bilirubin absorption peak around 460 nm, sOCT suffers from a strong signal decay with depth, which we overcome by optimizing our system sensitivity through a combination of zero-delay acquisition and focus tracking. In a phantom study, we demonstrate the quantification of bilirubin concentrations between 0 and 650 µM with only a 10% difference to the expected value, thereby covering the entire clinical pathophysiological range. aspect of the technique, we also measured bilirubin concentrations behind an epidermis mimicking scattering layer with an accuracy of 10%.


Introduction
Bilirubin is the yellow, toxic breakdown product of hemoglobin, which is excreted by the body in bile and urine. Elevated levels of bilirubin lead to jaundice: a yellow discoloration of the eyes and skin. In case of severe jaundice (hyperbilirubinemia), bilirubin can accumulate in the brain and induce kernicterus, which results in irreversible brain damage [1]. Therefore, it is of vital importance to monitor bilirubin levels in high-risk population groups, such as newborns. Unfortunately, current clinically applied methods for bilirubin monitoring are either invasive (blood sampling), or are not accurate enough to fully replace invasive blood sampling (transcutaneous bilirubinometry by diffuse reflectance spectroscopy) [2,3].
Here, we investigate the use of visible-light spectroscopic optical coherence tomography (sOCT) for noninvasive bilirubin determinations. The unique advantage of sOCT is that it allows for both quantitative and spatially confined measurements of bilirubin concentrations. This potentially enables the noninvasive determination of bilirubin concentrations within blood vessels, facilitating a direct comparison with invasive blood sampling without any cross talk from surrounding tissue. Since the existing transcutaneous bilirubinometers are unable to spatially resolve detected photons, cross talk from surrounding tissue is their main accuracy limiting factor [2]. Another important potential application of sOCT is the noninvasive study of local bilirubin extravasation into skin and brain tissue. As such a study is currently impossible, this may lead to a more fundamental understanding on the development of pathologies like kernicterus [2] and processes like the cephalocaudal progression of jaundice [4].
Multiple studies have shown the feasibility of sOCT and the closely related technology low-coherence spectroscopy (LCS) for the ex vivo [5,6] and in vivo quantification of absorber concentrations in the visible wavelength range, including preclinical applications for highly localized tissue oximetry [7][8][9][10][11][12][13]. Using sOCT for bilirubin quantification in tissue introduces several sensitivity-related challenges, as bilirubin absorbs around the relatively short wavelengths of 440 nm (free bilirubin) and 470 nm (albumin-bound bilirubin). This wavelength region is not only associated with impaired penetration depth into tissue, but also comes with a sharper roll-off of the sensitivity with depth due to the finite pixel size of the detecting spectrograph [14]. Hence, bilirubin determinations require an sOCT system with superior sensitivity, which we realize here by combining spectral domain sOCT with focus tracking and zero-delay acquisition throughout the entire axial scanning range. Besides the optimization of system sensitivity, this also ensures that the measured attenuation is only affected by the sample itself, resulting in quantitative measurements of the optical properties of the sample without requiring any calibration procedure.
In this work, we validate our sOCT system in the wavelength range of 440-600 nm for quantitative bilirubin determinations in samples mimicking neonatal skin. We show that our sOCT method is able to estimate bilirubin concentrations up to 650 µM with an accuracy of 10% and a standard deviation in the order of 50 µM. The investigated bilirubin range covers the entire clinical pathophysiological range (50 -500 µM) [2,3]. To demonstrate the spatially confined aspect of the technique, we also measured bilirubin concentrations behind an epidermis mimicking scattering layer with an accuracy of 10%.  . To l axial range, i our previous w or image [15]. es a modulation city of the mov cquisition of on piezo-driven m res that for ev nt speed is cons akes 5 second n of I det (λ,t) w aining the freq ment ( Fig. 2

Short time Fourier transformation
For the purpose of quantitative and spatially confined spectroscopy, we are interested in I s as a function of both wavelength and geometrical depth d, with d = ΔOPL/(2n). Since I filt (λ,t) contains the spectral content of the OCT signal over the entire imaging range of the spectrograph (d max = 622 µm), the spectra are not yet spatially confined at each depth. Short time Fourier transformation (STFT) of I filt (λ,t) with respect to λ with a rectangular spectral window with a size of Δλ = 5 nm, results in a depth-resolved spectral data set I STFT (λ,d,t) with a spatial resolution Δd = λ 2 /(2nΔλ) ranging between 14 µm (at λ = 440 nm) and 27 µm (at λ = 600 nm). The STFT is applied directly in the wavelength domain, since an interpolation filter to convert data to the equidistant wave number domain has its own frequency characteristics which affect depth information and thus affects quantitative analysis. Averaging of |I STFT (λ,d,t)| with respect to time (i.e. for every line captured by the camera) yields I STFT (λ,d) (Fig. 2, step 2e). Since I r remains constant during the complete measurement, I STFT (λ,d) contains the depth profile of s I per wavelength, from which we obtain S(λ) = I STFT (λ,d = ∆d) (Fig. 2, step 2f). The backscattered spectrum at d = ∆d is used, since it has a better signalto-noise ratio compared to the backscattered spectrum at d = 0 (i.e. exactly at zero-delay), due to small remainders of the DC term after filtering.

Focus tracking and depth-resolved acquisition
In addition to zero-delay acquisition, we further optimize our system sensitivity by focus tracking. At the start of each measurement, the zero-delay and focus position are matched at the interface between the cuvette wall and the sample (Fig. 2, step 1). Subsequently, S(λ,d ZD ) is obtained by acquisition of S(λ) as a function of the zero-delay (ZD) position relative to the sample's surface d ZD , with step size dx and the method described in sections 2.2.1 and 2.2.2. Hereto, the reference arm is elongated with steps of dx n ⋅ and focus tracking is achieved by translating the sample lens with steps dx/n towards the sample (Fig. 2, step 3). This method for combined zero-delay acquisition and focus tracking results in a depth resolved and wavelength resolved OCT signal S(λ,d ZD ), from which the attenuation in depth only depends on the optical properties of the sample (Fig. 2, step 4).

Spectra of optical properties
Under the assumption of single scattering, the attenuation coefficient µ t (λ) of the sample is obtained by fitting a linear Lambert-Beer model to the natural logarithm of S(λ,d ZD ) (Fig. 2 with α(λ) and µ t (λ) free running fit parameters, and S bg a background term that is obtained at a depth of 1000 µm inside the non-scattering back wall of the sample containing cuvette. The individual contributions of the scattering coefficient µ s (λ) (modeled as a scatter power function aλ -b ) and the absorption coefficient µ a (λ) to the total sample attenuation µ t (λ) are obtained by fitting Eq. (4) to µ t (λ), according to our method in [12] (Fig. 2, step 6): with fit parameters: scaling factor a, scatter power b, and C i the concentration of the i th chromophore relative to the chromophore concentration at which the reference absorption spectrum µ a,i (λ) was measured. The lower boundaries of all fit parameters were set to 0. We obtained all µ a,i (λ) by transmission spectroscopy (UV-2401PC, Shimadzu, Japan), comprising the absorption spectra for free bilirubin and albumin-bound bilirubin.
To valida spectrum [  Fitting Eq bilirubin conc with, and wit within 10% w

Discussio
In this study, concentration skin mimicki bilirubin abso optical scatter confined mea At the bili the problem spectrograph' delay acquisit 11], our tech requires acqu reference mir measurement derive the att b) shows the b λ). Except for cted values. 4. a. OCT image uation spectra µ t (λ age of 3 measurem imate the theoretic a) shows the O O 2 -silicone lay er, after which depth acquisiti nuation spectru wn in Fig. 4 For potential future in vivo measurements, it is likely that motion artefacts occur within this measurement time, hampering both the quantitative, and localized analysis of the optical attenuation. Measurement time can be reduced by a) reducing the period of the reference mirror movement, b) decreasing the number of averages per depth position, and c) increasing axial step size dx or, depending on the application, reducing the depth range to a spatially even more confined region of interest. Options b) and c) potentially come at the cost of measurement accuracy. For the sample measurement behind the scattering silicone layer, the measured µ t (λ) is lower and has larger standard deviations than the µ t (λ) measured without the layer for wavelengths <460 nm (Fig. 4(b)). This can be caused by a) a lower signal to noise ratio inside the sample, due to the increased attenuation by the scattering silicone layer in this spectral region, and b) the influence of multiple scattering on the signal at increasing measurement depths. The presence of multiple scattering results in a lower scattering contribution when fitting Eq. (4) to the measured µ t (λ) [16]. Since absorption is not affected by the influence of multiple scattering [5], we still obtain approximately the same absorption contribution -and thus bilirubin concentration -compared to the measurement without the covering layer. Nevertheless, when translating this method to spatially confined measurements in tissue, a thorough investigation is required of the measurement accuracy as a function of depth and tissue optical density.
For all samples containing bilirubin, the averaged bilirubin concentration measurements agree within 10% of the expected values. The overestimation of the bilirubin concentration in the 0 µM bilirubin sample may be explained by the lower boundary of 0 for all fit parameters in Eq. (4), resulting in a positive bias for samples containing no or very little bilirubin. Since the standard deviation of the data is in the order of 50 µM, the precision of our method is lower than its accuracy. This precision is comparable to the precision of transcutaneous bilirubinometers for concentrations up to 200 µM [3]. Transcutaneous bilirubinometers however, systematically underestimate the clinically relevant bilirubin concentrations higher than 200 µM [3], whereas the accuracy of our method remains relatively constant for the full investigated range. Furthermore, sOCT has the ability of spatially confining the measurement volume to a small region of interest inside tissue. This can be of great value to I) study the physiological process of bilirubin extravasation into tissue, and II) noninvasively measure bilirubin concentrations inside blood vessels. The latter would overcome the intrinsic accuracy-limiting factor of current transcutaneous bilirubinometers, as these measure the bilirubin concentration in a relatively large skin volume, which does not correlate directly to the bilirubin concentration in blood [2]. Further research is required to investigate whether sOCT can accurately measure bilirubin concentrations in blood where new challenges such as tissue dynamics, especially perfusion, and high hemoglobin absorption (µ a ≈20 mm −1 ) around the same wavelengths as the bilirubin absorption peak arise.

Funding
Innovational Research Incentives Scheme of The Netherlands Organisation for Scientific Research (NWO), division Applied and Engineering Sciences (TTW) (personal grant NB: VENI-13615); Pioneers in Healthcare Innovation Fund (University of Twente); University of Twente.