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A novel non-imaging optics based Raman spectroscopy device for transdermal blood analyte measurement
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View: Figures


Image of FIG. 1.

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FIG. 1.

Comparison of CPC and CHC lengths with respect to maximum collimation conic half angles for a fixed input aperture radius of 2 mm, whose dimensions were determined as shown in Sec. II B.) Maximum half angle of zero degrees corresponds to perfect collimation.

Image of FIG. 2.

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FIG. 2.

(A) Comparison of CPC and CHC with the same input aperture size and degree of collimation. Both CPC and CHC have input aperture radius of 2 mm and maximum collimation half angle of 3.9°. (B) With the help of a matching focusing lens, the CHC achieves collimation while maintaining relatively short physical length. Due to the hyperbolic shape of the reflector, light emanating from point I (the actual focus of the lens) are collimated by the lens as if it were coming from point II.

Image of FIG. 3.

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FIG. 3.

(A) CHC formed by two hyperbolae and its associated design parameters. R i : input aperture radius, R o : output aperture radius, L: CHC length, θ max : maximum collimation conic half angle, ϕ: angle between the two hyperbolae, f 1A , f 2A , f 1B , f 2B : foci of hyperbolae. The diameter and focal length of the matching plano-convex focusing lens are clearly defined by R o and the distance between the output aperture and the hyperbolic foci (f 2A and f 2B ). (B) CHC design optimized for the clinical Raman instrument.

Image of FIG. 4.

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FIG. 4.

ZEMAX ray-tracing simulation for design verification. The optimized CHC design of Fig. 3(B) was imported into ZEMAX, and coupled with a matching plano-convex focusing lens of 150 mm focal length. Another focusing lens of f/2 was placed in front of the CHC output aperture, and a detector was placed near its focus. In order to simulate the limited numerical aperture (f/2) of the collection fiber bundle (Fig. 5), the detector only accepted light coming in at conic half angles of ±14° or less. With an isotropic light source (conic half angle ±90° or less) at the CHC input aperture, the detector received 78% of the original light intensity.

Image of FIG. 5.

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FIG. 5.

Transmission mode clinical Raman spectroscopy setup with CHC. (NDF: Neutral density filter, BPF: Bandpass filter, S: Shutter, FL : Focusing Lens, OFB: Optical fiber bundle.) With the Raman excitation light illuminated on one side of the thin tissue sample (such as the thenar skin fold), the Raman scattered light emerges from the other end of the tissue. The CHC collects Raman scattered light at all conic half angles (±90°) and collimates to within ±3.9°. The collimated light is then filtered through a holographic Rayleigh rejection notch filter and focused onto the collection fiber bundle. On the other end, the collection fiber bundle is arranged to form a column of light with its height matched to the size of the CCD detector.

Image of FIG. 6.

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FIG. 6.

CHC-based clinical Raman spectroscopy instrument used for human subject studies. (A) The entire instrument fits inside a wheeled cart to facilitate portability during clinical studies. (B) The manufactured CHC seen from the output aperture. The interior surface was evenly coated with pure gold and polished to optical grade fineness. (C) Side view of the transmission mode Raman spectroscopy setup.

Image of FIG. 7.

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FIG. 7.

Back-scattered mode Raman spectroscopy system with an optical fiber probe. (NDF: Neutral density filter, BPF: Bandpass filter, S: Shutter, FL : Focusing Lens, OF: Optical fiber, OFB: Optical fiber bundle.)

Image of FIG. 8.

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FIG. 8.

Clinical data obtained from one human subject (26 data points) using the Raman spectroscopy instrument during an oral glucose tolerance test. (A) Raw thenar skin fold tissue Raman spectra obtained at various blood glucose concentrations. (B) Time course of glucose concentrations measured during an oral glucose tolerance test, as measured by conventional finger-prick glucometer (○) and Raman spectroscopy via leave-one-out cross validation (×).

Image of FIG. 9.

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FIG. 9.

Leave-one-out cross validation results showing glucose concentrations measured by Raman spectroscopy and PLS with respect to reference glucose concentrations measured by a conventional finger-stick glucometer. (R2=0.81, Root-mean-squared-error-of-prediction=16.8 mg/dl, 18 human subjects, 28 OGTTs, 730 data points.)


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Due to its high chemical specificity, Raman spectroscopy has been considered to be a promising technique for non-invasive disease diagnosis. However, during Raman excitation, less than one out of a million photons undergo spontaneous Raman scattering and such weakness in Raman scattered light often require highly efficient collection of Raman scattered light for the analysis of biological tissues. We present a novel non-imaging optics based portable Raman spectroscopy instrument designed for enhanced light collection. While the instrument was demonstrated on transdermal blood glucose measurement, it can also be used for detection of other clinically relevant blood analytes such as creatinine, urea and cholesterol, as well as other tissue diagnosis applications. For enhanced light collection, a non-imaging optical element called compound hyperbolic concentrator (CHC) converts the wide angular range of scatteredphotons (numerical aperture (NA) of 1.0) from the tissue into a limited range of angles accommodated by the acceptance angles of the collection system (e.g., an optical fiber with NA of 0.22). A CHC enables collimation of scattered light directions to within extremely narrow range of angles while also maintaining practical physical dimensions. Such a design allows for the development of a very efficient and compact spectroscopy system for analyzing highly scattering biological tissues. Using the CHC-based portable Raman instrument in a clinical research setting, we demonstrate successful transdermal blood glucose predictions in human subjects undergoing oral glucose tolerance tests.


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752b84549af89a08dbdd7fdb8b9568b5 journal.articlezxybnytfddd
Scitation: A novel non-imaging optics based Raman spectroscopy device for transdermal blood analyte measurement