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Flexible silicon sensors for diffuse reflectance spectroscopy of tissue DAVID M. M ILLER
AND
N AN M. J OKERST *
Department of Electrical and Computer Engineering, Duke University, 101 Science Drive, Durham, North Carolina 27708, USA * [email protected]
Abstract: Diffuse reflectance spectroscopy (DRS) is being used in exploratory clinical applications such as cancer margin assessment on excised tissue. However, when interrogating nonplanar tissue anomalies can arise from non-uniform pressure. Herein is reported the design, fabrication, and test of flexible, thin film silicon photodetectors (PDs) bonded to a flexible substrate designed for use in conformal DRS. The PDs have dark currents and responsivities comparable to conventional Si PDs, and were characterized while flat and while flexed at multiple radii of curvature using liquid phantoms mimicking adipose and malignant breast tissue. The DRS and nearest neighbor crosstalk results were compared with Monte Carlo simulations, showing good agreement between simulation and experiment. c 2017 Optical Society of America
OCIS codes: (040.5160) Photodetectors; (170.6510) Spectroscopy, tissue diagnostics; (170.6935) Tissue characterization.
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#282709 Journal © 2017
https://doi.org/10.1364/BOE.8.001512 Received 13 Dec 2016; revised 28 Jan 2017; accepted 9 Feb 2017; published 14 Feb 2017
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14. R. A. Schwarz, W. Gao, D. Daye, M. D. Williams, R. Richards-Kortum, and A. M. Gillenwater, “Autofluorescence and diffuse reflectance spectroscopy of oral epithelial tissue using a depth-sensitive fiber-optic probe,” Appl. Opt. 47(6), 825 (2008). 15. G. Zonios, L. T. Perelman, V. Backman, R. Manoharan, M. Fitzmaurice, J. Van Dam, and M. S. Feld, “Diffuse reflectance spectroscopy of human adenomatous colon polyps in vivo,” Appl. Opt. 38(31), 6628 (1999). 16. H.-W. Wang, J.-K. Jiang, C.-H. Lin, J.-K. Lin, G.-J. Huang, and J.-S. Yu, “Diffuse reflectance spectroscopy detects increased hemoglobin concentration and decreased oxygenation during colon carcinogenesis from normal to malignant tumors,” Opt. Express 17(4), 2805 (2009). 17. Z. Volynskaya, A. S. Haka, K. L. Bechtel, M. Fitzmaurice, R. Shenk, N. Wang, J. Nazemi, R. R. Dasari, and M. S. Feld, “Diagnosing breast cancer using diffuse reflectance spectroscopy and intrinsic fluorescence spectroscopy,” J. Biomed. 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1.
Introduction
Cancer margin assessment performed in the surgical suite on excised breast tissue has the potential to significantly reduce costly, emotionally challenging re-excision surgery [1]. One approach to margin assessment is Diffuse Reflectance Spectroscopy (DRS), the analysis of reflected light from biological tissue, which is strongly affected by tissue structural and chemical components. Light incident on tissue undergoes multiple scattering and absorbing events while propagating, and a portion of this light returns to the surface as diffuse reflectance. DRS probes are designed to illuminate tissue in a controlled manner, detect the diffuse reflectance, and use the DRS signal to identify the wavelength-dependent coefficients of absorption (µa ) and scattering (µs ), which are directly related to the concentrations of sub-cellular organelles and biochemical markers. The process of extracting the optical properties can be performed by comparing the reflectance data to a previously calculated table [2–4], to a model based on the diffusion approximation [5, 6], or to Monte-Carlo simulations [7–10], yielding quantitative information that can determine chromophore concentrations in skin [11], diagnosis of skin cancers [12], detect oral [13, 14], colon [15, 16], and breast cancer [8, 17], and perform intra-operative breast cancer margin assessment [18, 19]. DRS systems include a light source, tissue illumination and collection probes, a spectrometer, and a detector. DRS probe designs include fiber-based light sources and detectors, where the collection fibers are either separated by a fixed distance [12, 16] or concentrically located around the source fiber [7, 11, 13–15, 17–19]. However, fiber-based probes limit the collection efficiency of the probe since the limited fiber numerical aperture (NA), fiber cladding, and circular crosssection reduce the total optical power that reaches the remote optical detectors, even in tightly packed fiber bundles. In contrast, the NA and packing density of photodetectors (PDs) can be significantly larger than optical fibers, particularly if the PDs are placed directly on or close to the tissue sample. One example of this is a custom 4 x 4 array of annular Si PDs implemented in a DRS probe for breast cancer margin assessment [20]. The collection efficiency of the Si PD probe was higher due to the larger NA of the Si PD (NA = 0.96), as compared to a typical DRS optical fiber (NA = 0.22-0.39). Apertures located the center of each annular PD in the 16 pixel array was used to illuminate the tissue through the center of each PD, which enabled control of the tissue illumination profile and optimized the collection efficiency of the PDs by surrounding each illuminated tissue region with detector material. The optimal pixel spacing was dependent upon the pixel to pixel crosstalk, defined as the reflectance signal measured at a PD due to aperture illumination of neighboring PDs. Higher resolution imaging for this system can be achieved through multiple probe placements on a single sample, which improved the resolution from the array pixel-to-pixel spacing (4.5mm) to sub-millimeter (0.75mm) [21]. Si PD arrays can also be scaled to multiple closely packed, concentric detectors, enabling Spatially Resolved DRS SRDRS [22], which enables depth sensitive measurements through multiple source-detector pairs at each illumination site on a tissue sample. Whether fiber or PD probes are used, flattening a nonplanar tissue sample can lead to inaccurate margin assessment. Measurements of breast tissue specimen dimensions can have a 46% loss in height between the time of surgery and pathology, on average decreasing from 2.57 to 1.36 cm [23]. When pressure is applied to excised tissue, there can be a loss in DRS signal accuracy, as pressure affects the measured optical properties of local tissue regions [24–26]. In flattened tissue samples, pathological examination found positive margins, however, upon re-excision to achieve a clear margin, 56% of the re-excised patients showed no residual tumor in resected tissue [23]. This poses a challenge for rigid, planar DRS probes used in margin assessment, since nonplanar tissue samples must be compressed to ensure contact with the probe surface across the full imaging area. Peaks and valleys in the tissue sample can create high pressure regions or gaps against the detector plane. An air gap between the detection plane and the tissue can also affect the measurement by decreasing the collection efficiency of reflected light by the fibers
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or PDs. The flexible PDs and substrate reported herein address this concern by enabling probe conformation to the uneven topography of tissue. To construct a flexible DRS probe, both the PDs and the substrate must be flexible. Bonding a flexible thin film single crystal semiconductor PD to a flexible substrate provides the highest responsivity thin film PD available, in contrast to polycrystalline and organic PDs. High responsivity, low dark current thin film single crystal PDs have been reported for over 30 years [27, 28] bonded to rigid substrates. Reports of flexible Si PDs bonded to flexible substrates are emerging, however, these PDs have not been designed or evaluated for DRS measurements [29]. For example, flexible optoelectronics constructed by integrating single crystal Si [30] or Ge [31] onto flexible plastic substrates have been applied to portable, low cost photovoltaic systems and imaging applications. Herein, the design, fabrication, and test of a probe comprised of flexible, thin film Si custom annular PDs bonded to a flexible substrate optimized for DRS imaging is reported. The flexibility of these 10 µm thick PDs enables conformal tissue sensing. The shape and size of these PDs was based upon previously reported conventional thick, rigid PDs that have been evaluated in the clinic [32]. The surface normal spectral response and dark current of the thin film PDs are comparable to the conventional Si PDs, validating the PD performance. The thin film PDs were tested on liquid phantoms with radii of curvature on the imaged face ranging from flat to 10 mm. Two liquid phantoms were prepared, representing benign and malignant breast tissue. Diffuse reflectance measurements were performed at five curvatures on both phantoms. The measured DRS photocurrents were compared to both the similarly shaped conventional PD measurements, and to Monte Carlo simulated reflectance values, with good agreement between simulation and experiment. 2.
Flexible DRS sensor design
The flexible DRS sensor reported herein is comprised of two custom annular thin film, single crystal Si pn junction PDs that are bonded to a flexible Kapton polymer substrate. Each PD annulus has an inner diameter of 0.75 mm and an outer diameter of 2.5 mm. The centerto-center distance between PDs is 4.5 mm, as illustrated and photographed in Fig. 1. The DRS photodetector design was optimized for high optical throughput and high signal to noise ratio [20] by varying the PD shape, size, and spacing to maximize the collected DRS signal while minimizing the back illumination and crosstalk signals.
Fig. 1. Illustration of thin film flexible PDs bonded to a flexible substrate. The DRS probe contains two flexible thin film Si PDs with center-to-center separation of 4.5 mm.
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2.1.
Materials and fabrication process
The thin film Si PDs were fabricated using standard Si manufacturing processes, leveraging low cost Si fabrication technology and low cost thin film photodiode manufacturing technology [33]. The PDs were fabricated on a commercial Si-on-Insulator (SOI) wafer: 400 µm thick Si handle layer (Addison Engineering) / 1µm SiO2 buried oxide layer / 10µm thick p-Si (30-60 Ω-cm, Boron) device layer. To fabricate the PDs, a n-type spin-on dopant (phosphorous-doped, 5×1020 cm −3 , Emulsitone Chemicals) was spin coated and annealed at 1000◦ C for 15 minutes to drive in an approximately 500 nm deep n-type region to form the p-n junction. The dopant oxide was then removed and ring shaped top contacts (100nm Ti / 200nm Au) were evaporated and patterned. The annular photodetector mesas were then defined using photolithography and the silicon device layer was etched in a deep reactive ion etcher (DRIE, SPTS Pegasus) using the buried oxide as a selective etch stop. The SOI wafer was then bonded to a Pyrex substrate using a Waferbond HT 10.10 (Brewer Sciences), and the 400 µm thick silicon handle layer of the SOI wafer was removed using DRIE. The buried oxide was then removed with buffered HF, leaving only the 10 µm thick device embedded in the Waferbond. Next, a broad area back contact (100nm Al / 50nm Ti / 50nm Ni / 200nm Au) was deposited on the device by evaporation, and was patterned. The thin film PDs were then bonded to a flexible polyimide substrate (DuPont Kapton HN, 127 µm thick) using thermocompression bonding. The Kapton surface was planarized by spin coating a 5 µm thick polyimide film (PI 2525, HD Microsystems) onto the Kapton, followed by a cure in a N2 atmosphere at 350◦ C for two hours. A patterned metal layer for thermocompression bonding of the PDs to the flexible substrate was then vacuum deposited and patterned (100 nm Ti / 200 nm Au) on the Kapton/polyimide substrate. Thermocompression bonding between the PD contact and the metal pad on the Kapton/polyimide substrate was performed in a vacuum oven at 250◦ C with a 1 kg weight applied. After bonding, the Pyrex wafer was removed by heating the assembly to 225◦ C, mechanically debonding the temporary adhesive and enabling removal of the Pyrex wafer. The residual Waferbond was removed using trichloroethylene. To complete the flexible probe, the Kapton/polyimide substrate with the bonded Si PDs had top leads and an anti-reflection coating deposited. To isolate the top contacts, 5 µm of polyimide PI 2525 was patterned and cured as an interlayer dielectric. Evaporated leads (100nm Ti / 200nm Au) were then deposited and patterned to electrically connect the PD top contacts to the device pads. These leads were designed with a serpentine pattern to enable flexing of the leads on the substrate without damage to the leads. Silicon nitride (52 nm thick) was deposited using plasma-enhanced chemical vapor deposition (PECVD, Advanced Vacuum Vision) as a surface passivation layer and anti-reflection coating. Finally, a transparent optical aperture was opened through the Kapton/polyimide substrate by CO2 laser ablation (H-Series 20x12, Full Spectrum Laser). To avoid thermal damage to the PD, the optical aperture had a diameter of approximately 200 µm, which is smaller than the annular Si PD 750 µm inner diameter. A photograph of the completed flexible probe is shown as Fig. 2. 2.2.
Photodetector characterization
Dark current and responsivity measurements were performed on the flexible PDs bonded to the Kapton substrate using a Keithley 4200 source measurement unit (SMU) and a Newport 300W Xenon lamp with a Newport CS130 monochromator. To measure the responsivity, the PDs were illuminated in a surface normal configuration. The Xenon lamp was coupled to a 1 mm diameter optical fiber with a NA of 0.39, and the output of the fiber was focused through a microscope objective lens to illuminate the front surface of the PD. The optical wavelengths measured were from 470 nm to 600 nm in increments of 10 nm, for a total of 14 measurements. The typical illumination power at the output of the microscope objective varied slightly with the lamp output wavelength, and was approximately 150 nW. The two thin film PDs were labeled PD1 and PD2.
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Fig. 2. Image of the flexible PD probe bent cylindrically outward, bending radius of 15 mm.
PD1 had higher responsivity and lower dark current than PD2. The responsivity for each PD was averaged over 100 measurements for each wavelength. The responsivity for PD1 at zero bias was 0.20-0.34 A/W for λ = 470 – 600 nm, and for PD2 the measured responsivity was 0.10 to 0.17 A/W for the same wavelength range. The zero bias dark current (averaged over 1000 measurements) was 10 pA for PD1, and 750 pA for PD2. With a total photodiode surface area of 0.0447cm2 , the dark current density for PD1 was 0.224 nA/cm2 , and 16.8 nA/cm2 for PD2. The poor performance of PD2 is is likely due to spiking of the contact into the shallowly diffused pn junction, which would result in the observed mixed photodiode/photoconductor behavior. This poor yield can be addressed with better process uniformity control across the SOI wafer. The rest of the paper will focus on measurements performed on the higher performance PD1. The zero-bias dark current of PD1 is similar to that of previously reported thin film Si devices of the same conformation and size bonded to Pyrex and Kapton, the PD1 responsivity exceeds previously reported values for flexible PDs on Kapton [29], and is comparable to the conventional devices on rigid Pyrex substrates [29]. Thus, thin film PD1 is a high performance device, with high responsivity and low dark current. Note that the higher responsivity of PD1 can be attributed to the anti-reflection coating, which was absent in the previously reported flexible Si PDs on Kapton. Surface normal responsivity measurements were also performed on PD1 under flexed conditions. The responsivity was measured using the same responsivity test apparatus as above, except that the PD/Kapton probe was held in a series of 3D printed plastic clamps with five different radii of curvature, which flexed the PD/Kapton probe. The responsivity for the wavelength range 470-600 nm for 5 curvatures is shown in Fig. 3(a). Each data point is the average of 100 measurements at each wavelength/curvature combination, with a standard deviation of 0.61%. The PD1 dark current and current-voltage characteristic for each curvature is shown in Fig. 3(b), averaging 100 non-illuminated measurements at zero bias voltage. The zero-bias dark current for PD1 ranged from 9-12 pA for all five curvatures. All PD1 performance metrics show negligible change as a function of flexure. 3. 3.1.
Phantom testing Liquid phantoms
Two liquid phantoms were prepared to represent human breast tissue (adipose and malignant), with scattering and absorbing coefficients within the range of reported values for wavelengths 470-600 nm [34]. Nigrosin absorber (Sigma-Aldrich) was mixed in deionized water with a concentration of 5.54×10 −5 weight/volume (benign) and 1.10×10 −4 w/v (malignant). Scattering polystyrene spheres, 0.99 µm diameter (Polysciences, Inc.), were added at a concentration of
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Fig. 3. (a) Surface normal responsivity of PD1 at multiple bend radii; (b) Dark current characterization of PD1 at multiple bend radii.
2.24×10 −3 w/v (benign) and 3.54×10 −3 w/v (malignant). The absorption coefficient was calculated from spectrophotometer transmission measurements (Shimadzu UV-3600) of the phantom before adding the scattering polystyrene spheres. The scattering coefficient and anisotropy were estimated using Mie theory calculations [35]. The measured absorption and calculated scattering properties for both phantoms are shown in Fig. 4.
Fig. 4. Absorption (µa , dotted lines) and scattering (µs , solid lines) coefficients of the liquid phantoms representing adipose (red lines) and malignant (blue lines) breast tissue.
For the wavelength range 470-600 nm, the absorption coefficient ranged from 2.1 to 6.0 cm −1 , and the scattering coefficient ranged from 87 to 182 cm −1 . The anisotropy (g) ranged from 0.921 to 0.927, as calculated from Mie theory. The reduced scattering coefficient, µs ’, where µs ’=µs (1-g), ranged from 6.9 to 13.3 cm −1 . 3.2.
Measurement apparatus
To perform diffuse reflectance measurement while flexing the PD/Kapton probe at various curvatures, pairs of plastic clamps were 3D printed with five discrete radii of curvature: flat (infinite radius), 50 mm, 25 mm, 15 mm, and 10 mm. The upper section of each pair of 3D printed
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parts included a hollow volume for containing the liquid phantom during testing. The inner walls of this hollow volume were painted black to minimize sidewall reflections. The open end of the phantom container was capped with 50 µm thick transparent PET film (McMaster-Carr). The PD/Kapton probe was placed below the PET film with the detectors facing upwards, towards the phantom. The two detectors were aligned equidistant from the center of the front face of the phantom, as illustrated and photographed in Fig. 5. Illumination was provided through a 1 mm diameter optical fiber (NA=0.39) aligned to the center aperture of PD1. To perform DRS measurements, the light source fiber was aligned to the back of the PD1 aperture. For crosstalk measurements, the test was repeated with the source fiber aligned to the PD2 aperture, and the PD1 photocurrent was measured. The optical power incident on the phantom through these apertures varied with wavelength, from 0.9-2.0 µW.
Fig. 5. Left: Illustration of liquid phantom testing at 10 mm radius of curvature: (a) Top clamp, with hollow region to contain liquid phantom and curved bottom edges; (b) Transparent PET film; (c) Flexible PDs on Kapton; (d) Bottom clamp, with curvature matching the top clamp. Right: Photograph of 10 mm bending radius test apparatus.
For each different curvature under test, the total throughput optical power was measured on a power meter (Thorlabs PM100) and the fiber alignment to the PD was optimized for maximum optical throughput. This measurement was taken with a dry container before the liquid phantom was added, and the fiber to PD alignment was unchanged during all of the phantom tests for that particular curvature. The diffuse reflectance for each curvature was then measured by placing each phantom into the container separately and illuminating the PD with the fiber. To normalize for light source fluctuations and fiber-to-aperture alignment mismatch, a reflectance standard photocurrent was recorded before each phantom and curvature measurement. The reflectance standard (Labsphere, Inc. SRS-99-010) was measured with the PD probe held in another, shorter 3D printed clamp that enabled close proximity between the PD and reflectance standard. For each wavelength, 50 reflectance standard measurements were averaged. The reflectance standard was then removed and the optical throughput was measured for this configuration on a power meter (Thorlabs PM100). The scaled flat case reflectance standard data shows that the thin film PD1 DRS signal is comparable to the previously reported DRS signals from a comparable rigid thick DRS detector [20, 32, 36]. Scaling the input power to compare the thin film and typical thickness (300 µm – 500 µm thick) PD measurements on a 99% reflectance standard yields an average percent difference of 3.33% over the common experimental wavelengths of 500 nm, 530 nm, 560 nm, and 600 nm. After each reflectance measurement, the liquid phantom was replaced with a highly absorbing nigrosin solution (µa > 50cm −1 ), and a background current measurement was taken. Herein, background current is defined as a PD output current that is not generated by the DRS signal.
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Since the PD is illuminated from the back side, and the light passes through the PD aperture, back illumination current can be generated by absorption at the inner edges of the PD. Also, reflection from the PET film can cause background current, and the photodiode dark current is considered a background current. To calculate the diffuse reflectance at each wavelength, the measured background current was subtracted from the corresponding reflectance measurement to assess only the photocurrent due to diffuse reflectance. This value was then divided by the optical throughput scaled by the ratio of the reflectance standard photocurrent to the reflectance standard throughput, since the puck and curved measurements had slightly different optical throughput values as the reflectance standard test was not on a curved surface. This expression for reflectance can be written as: R (λ, ρ, P) =
i phantom (λ, ρ, P) − i backillum (λ, ρ) i h d (λ) I0 (λ, ρ) Ii0,sst at anndar (λ) dar d
(1)
where R (λ, ρ, P) is the diffuse reflectance at wavelength λ, curvature ρ, and liquid phantom P; i phantom (λ, ρ, P) and i backillum (λ, ρ) are the photocurrent measurements of the liquid phantom reflectance and back illumination, respectively; I0 (λ, ρ) is the recorded optical throughput at wavelength λ, and curvature ρ; and i standard (λ) /I0,standard (λ) is the previously calculated ratio of reflectance standard photocurrent to optical throughput at wavelength λ. 3.3.
Measurement results
The reflectance data for fourteen wavelengths between 470 nm and 600 nm with 10 nm spacing for both phantoms at three of the five measured curvatures is shown in Fig. 6. The 50 mm and 25 mm curvatures are omitted for clarity; the reflectance values from these curvatures were similar to the flat test case. Each data point shows the average and standard deviation of 75 photocurrent measurements, in three successive spectral scans of 25 measurements each. Three successive scans were conducted to assess any bias due to phantom settling, which was negligible.
Fig. 6. DRS reflectance versus wavelength for both benign and malignant phantoms at fourteen wavelengths and three radii of curvature.
The data shows an increasing reflectance signal with increasing curvature (decreasing radius) for both phantoms. The increase in DRS signal due to curvature ranged from 5% to 8% for the benign phantom and 5% to 12% for the malignant phantom. This increase in reflectance is
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expected for increased curvature, since, in the diffusion approximation, the reflected intensity increases as the source-detector distance decreases [5]. In an inwardly curved state, the total path length within the bulk liquid phantom from optical source (PD aperture) to detector (PD surface) decreases. As a result, the collected diffuse reflectance optical power at the PD surface increases. Due to the higher absorption and higher scattering of the malignant phantom, it has a larger increase in reflected signal for the same reduction in path length compared to the benign phantom. The crosstalk was measured by illuminating the aperture of PD2 and measuring the PD1 photocurrent. The PD1 aperture was not illuminated. The measured crosstalk is shown in Fig. 7. The malignant phantom produced a significantly lower crosstalk signal, which is expected due to the higher absorption of the malignant phantom. Due to this high absorption, the crosstalk photocurrents were below the noise floor of the photodiode and could not be measured.
Fig. 7. Crosstalk reflectance for the benign phantom. All five radii of curvature were measured.
As a percentage of the DRS signal, the maximum crosstalk was 2.3% (benign phantom, λ=470 nm, 10 mm radius of curvature). The crosstalk signal shows a stronger dependence on curvature than the DRS signal, increasing up to 40% from a flat surface to a 10 mm radius of curvature. The decrease in path length due to increasing curvature is more pronounced in the case of the crosstalk detector than the DRS detector. Due to the larger reduction in path length, the percent change in crosstalk signal is higher when the phantom surface is curved. The crosstalk to DRS signal ratio remained below 3% at all curvatures and wavelengths. This crosstalk is comparable to the previously reported 6-9% maximum simulated crosstalk percent values for a rigid thick DRS system with similar probe geometry [32]. 4.
Zemax simulations
To validate the experimental results and toward a look-up table for coefficient extraction, Monte Carlo ray tracing simulations were performed using Radiant ZEMAX commercial software. The simulation geometry for each of the five curvatures were modeled. The side and back walls of the container were simulated as perfect absorbers, and the front wall was modeled as a 50 µm thick transparent PET film. The two PDs each had a 200 µm diameter aperture in the annulus. Rays were launched through the center aperture of the illuminated PD, incident on the phantom. The bulk scattering model in ZEMAX, with Henyey-Greenstein statistics for scattering angles,
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was used to simulate diffuse reflectance. All rays hitting each PD were summed to simulate the reflectance incident on that PD. Each curvature and wavelength were modeled with 50 million rays. The optical power recorded at the illuminated PD surface was defined as the DRS signal, and the power at the other, unilluminated PD was the crosstalk signal. The diffuse reflectance was calculated by dividing the optical power at each detector by the reflectance optical power from a 99% reflectance standard, where the reflectance standard was modeled separately using a Lambertian scattering surface. 4.1.
Comparison of theory to experiment
The simulation and experimental diffuse results for the two extreme curvature cases, flat and 10 mm, are compared in Fig. 8, and the crosstalk at five curvatures in Fig. 9, for all fourteen measured wavelengths.
Fig. 8. Experimental DRS signal (malignant phantom: red open circles, benign phantom: black open triangles) compared to Monte Carlo simulations (malignant phantom: red filled circles, benign phantom: black filled triangles) versus wavelength for the (a) flat test case and (b) 10 mm radius of curvature test case.
The mean percent error in the diffuse reflectance between simulation and experiment is 2.3% for the benign phantom and 1.23% for malignant phantom. The mean percent error in the crosstalk signal for the benign phantom was 4.85%. The main source of error was likely inaccuracy in the thickness of the PET film. For the malignant phantom, the low values of the crosstalk photocurrent were at or near the noise floor of the PD, which precluded measurement of the reflectance. The ability to differentiate between benign and malignant tissue in a DRS system can be evaluated by the signal contrast, which can be written as: S (λ) =
M (λ) − B (λ) B (λ)
(2)
where M (λ) and B (λ) are the malignant and benign diffuse reflectance signals, respectively, and S (λ) is the signal contrast. The signal contrast for the flat phantom measurements ranged from +4.0% at 570 nm to +13.6% at 470 nm, and the reflectance from the malignant phantom was higher than that of the benign phantom in all cases. The signal contrast at all fourteen wavelengths and five radii of curvature is shown in Fig. 10. The lowest contrast occurs at the wavelengths corresponding with the absorption peak of nigrosin, at approximately 560 nm. At all curvatures, the signal contrast remains positive. With
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Fig. 9. Experimental crosstalk reflectance (open triangles) compared to Monte Carlo simulations (filled triangles) for the benign phantom at five radii of curvature.
Fig. 10. Signal contrast between benign and malignant phantom reflectance at flat and flexed conditions.
increasing curvature, the average contrast across all wavelengths increases from 6.2% at the flat test case to 10.0% at a radius of 10 mm. The curvature has an approximately linear effect upon the diffuse reflectance. In Fig. 11, the diffuse reflectance experimental and simulated results are shown as a function of curvature (where curvature is the inverse of surface radius) for the benign and malignant phantoms for λ = 470, 500, 530, 580, and 600 nm for all five measured curvatures. Although the measured results are more scattered, the results from Zemax simulations suggest a linear dependence on curvature (average R2 = 0.992), shown as dashed lines. The phantom reflectance signal increased an average of 8.0% in simulation and 7.9% in measurement as the curvature increased from flat to a radius of 10 mm. In the experimental data, the increase in reflectance with increasing curvature was higher for the malignant phantom than
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Fig. 11. DRS signal (open circles) of (a) benign phantom; and (b) malignant phantom, compared to Monte Carlo simulations (filled circles) versus curvature (inverse of radius) at five wavelengths. Linear regression (dashed lines) shown for simulated data.
for the benign phantom, resulting in an increase in contrast with increasing curvature. 5.
Conclusions
We have designed, fabricated, and tested thin film, flexible single crystalline Si PDs on a flexible Kapton substrate. The PD1 responsivity and dark current performance was equal to that of conventional rigid PDs. The PD DRS performance was measured on two liquid phantoms, representing benign and malignant breast tissue. Measurements were taken over 5 curvatures, ranging from flat to 10 mm radius of curvature. The average error between the experimental results and Monte Carlo simulations for the DRS measurements was 1.8%. The flexible PDs presented herein demonstrate that flexible semiconductor probes can be used for DRS measurements, which may enable more accurate analysis of non-planar tissues. Increasing the phantom curvature from flat to a 10 mm radius of curvature increased the DRS signal by an average of 7.9%. The crosstalk to DRS signal ratio remained below 3% at all curvatures and wavelengths. Funding This research was supported by the National Institutes of Health through the Bioengineering Research Partnership award 1 RO1 EB011574-01A1. Additionally, this work was performed in part at the Duke University Shared Materials Instrumentation Facility (SMIF), a member of the North Carolina Research Triangle Nanotechnology Network (RTNN), which is supported by the National Science Foundation (Grant ECCS-1542015) as part of the National Nanotechnology Coordinated Infrastructure (NNCI). Acknowledgments The authors would like to thank the staff of Duke University Shared Materials Instrumentation Facility for their technical assistance, and Duke University and the Department of Electrical and Computer Engineering for their support.
Flexible silicon sensors for diffuse reflectance spectroscopy of tissue DAVID M. M ILLER
AND
N AN M. J OKERST *
Department of Electrical and Computer Engineering, Duke University, 101 Science Drive, Durham, North Carolina 27708, USA * [email protected]
Abstract: Diffuse reflectance spectroscopy (DRS) is being used in exploratory clinical applications such as cancer margin assessment on excised tissue. However, when interrogating nonplanar tissue anomalies can arise from non-uniform pressure. Herein is reported the design, fabrication, and test of flexible, thin film silicon photodetectors (PDs) bonded to a flexible substrate designed for use in conformal DRS. The PDs have dark currents and responsivities comparable to conventional Si PDs, and were characterized while flat and while flexed at multiple radii of curvature using liquid phantoms mimicking adipose and malignant breast tissue. The DRS and nearest neighbor crosstalk results were compared with Monte Carlo simulations, showing good agreement between simulation and experiment. c 2017 Optical Society of America
OCIS codes: (040.5160) Photodetectors; (170.6510) Spectroscopy, tissue diagnostics; (170.6935) Tissue characterization.
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1.
Introduction
Cancer margin assessment performed in the surgical suite on excised breast tissue has the potential to significantly reduce costly, emotionally challenging re-excision surgery [1]. One approach to margin assessment is Diffuse Reflectance Spectroscopy (DRS), the analysis of reflected light from biological tissue, which is strongly affected by tissue structural and chemical components. Light incident on tissue undergoes multiple scattering and absorbing events while propagating, and a portion of this light returns to the surface as diffuse reflectance. DRS probes are designed to illuminate tissue in a controlled manner, detect the diffuse reflectance, and use the DRS signal to identify the wavelength-dependent coefficients of absorption (µa ) and scattering (µs ), which are directly related to the concentrations of sub-cellular organelles and biochemical markers. The process of extracting the optical properties can be performed by comparing the reflectance data to a previously calculated table [2–4], to a model based on the diffusion approximation [5, 6], or to Monte-Carlo simulations [7–10], yielding quantitative information that can determine chromophore concentrations in skin [11], diagnosis of skin cancers [12], detect oral [13, 14], colon [15, 16], and breast cancer [8, 17], and perform intra-operative breast cancer margin assessment [18, 19]. DRS systems include a light source, tissue illumination and collection probes, a spectrometer, and a detector. DRS probe designs include fiber-based light sources and detectors, where the collection fibers are either separated by a fixed distance [12, 16] or concentrically located around the source fiber [7, 11, 13–15, 17–19]. However, fiber-based probes limit the collection efficiency of the probe since the limited fiber numerical aperture (NA), fiber cladding, and circular crosssection reduce the total optical power that reaches the remote optical detectors, even in tightly packed fiber bundles. In contrast, the NA and packing density of photodetectors (PDs) can be significantly larger than optical fibers, particularly if the PDs are placed directly on or close to the tissue sample. One example of this is a custom 4 x 4 array of annular Si PDs implemented in a DRS probe for breast cancer margin assessment [20]. The collection efficiency of the Si PD probe was higher due to the larger NA of the Si PD (NA = 0.96), as compared to a typical DRS optical fiber (NA = 0.22-0.39). Apertures located the center of each annular PD in the 16 pixel array was used to illuminate the tissue through the center of each PD, which enabled control of the tissue illumination profile and optimized the collection efficiency of the PDs by surrounding each illuminated tissue region with detector material. The optimal pixel spacing was dependent upon the pixel to pixel crosstalk, defined as the reflectance signal measured at a PD due to aperture illumination of neighboring PDs. Higher resolution imaging for this system can be achieved through multiple probe placements on a single sample, which improved the resolution from the array pixel-to-pixel spacing (4.5mm) to sub-millimeter (0.75mm) [21]. Si PD arrays can also be scaled to multiple closely packed, concentric detectors, enabling Spatially Resolved DRS SRDRS [22], which enables depth sensitive measurements through multiple source-detector pairs at each illumination site on a tissue sample. Whether fiber or PD probes are used, flattening a nonplanar tissue sample can lead to inaccurate margin assessment. Measurements of breast tissue specimen dimensions can have a 46% loss in height between the time of surgery and pathology, on average decreasing from 2.57 to 1.36 cm [23]. When pressure is applied to excised tissue, there can be a loss in DRS signal accuracy, as pressure affects the measured optical properties of local tissue regions [24–26]. In flattened tissue samples, pathological examination found positive margins, however, upon re-excision to achieve a clear margin, 56% of the re-excised patients showed no residual tumor in resected tissue [23]. This poses a challenge for rigid, planar DRS probes used in margin assessment, since nonplanar tissue samples must be compressed to ensure contact with the probe surface across the full imaging area. Peaks and valleys in the tissue sample can create high pressure regions or gaps against the detector plane. An air gap between the detection plane and the tissue can also affect the measurement by decreasing the collection efficiency of reflected light by the fibers
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or PDs. The flexible PDs and substrate reported herein address this concern by enabling probe conformation to the uneven topography of tissue. To construct a flexible DRS probe, both the PDs and the substrate must be flexible. Bonding a flexible thin film single crystal semiconductor PD to a flexible substrate provides the highest responsivity thin film PD available, in contrast to polycrystalline and organic PDs. High responsivity, low dark current thin film single crystal PDs have been reported for over 30 years [27, 28] bonded to rigid substrates. Reports of flexible Si PDs bonded to flexible substrates are emerging, however, these PDs have not been designed or evaluated for DRS measurements [29]. For example, flexible optoelectronics constructed by integrating single crystal Si [30] or Ge [31] onto flexible plastic substrates have been applied to portable, low cost photovoltaic systems and imaging applications. Herein, the design, fabrication, and test of a probe comprised of flexible, thin film Si custom annular PDs bonded to a flexible substrate optimized for DRS imaging is reported. The flexibility of these 10 µm thick PDs enables conformal tissue sensing. The shape and size of these PDs was based upon previously reported conventional thick, rigid PDs that have been evaluated in the clinic [32]. The surface normal spectral response and dark current of the thin film PDs are comparable to the conventional Si PDs, validating the PD performance. The thin film PDs were tested on liquid phantoms with radii of curvature on the imaged face ranging from flat to 10 mm. Two liquid phantoms were prepared, representing benign and malignant breast tissue. Diffuse reflectance measurements were performed at five curvatures on both phantoms. The measured DRS photocurrents were compared to both the similarly shaped conventional PD measurements, and to Monte Carlo simulated reflectance values, with good agreement between simulation and experiment. 2.
Flexible DRS sensor design
The flexible DRS sensor reported herein is comprised of two custom annular thin film, single crystal Si pn junction PDs that are bonded to a flexible Kapton polymer substrate. Each PD annulus has an inner diameter of 0.75 mm and an outer diameter of 2.5 mm. The centerto-center distance between PDs is 4.5 mm, as illustrated and photographed in Fig. 1. The DRS photodetector design was optimized for high optical throughput and high signal to noise ratio [20] by varying the PD shape, size, and spacing to maximize the collected DRS signal while minimizing the back illumination and crosstalk signals.
Fig. 1. Illustration of thin film flexible PDs bonded to a flexible substrate. The DRS probe contains two flexible thin film Si PDs with center-to-center separation of 4.5 mm.
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2.1.
Materials and fabrication process
The thin film Si PDs were fabricated using standard Si manufacturing processes, leveraging low cost Si fabrication technology and low cost thin film photodiode manufacturing technology [33]. The PDs were fabricated on a commercial Si-on-Insulator (SOI) wafer: 400 µm thick Si handle layer (Addison Engineering) / 1µm SiO2 buried oxide layer / 10µm thick p-Si (30-60 Ω-cm, Boron) device layer. To fabricate the PDs, a n-type spin-on dopant (phosphorous-doped, 5×1020 cm −3 , Emulsitone Chemicals) was spin coated and annealed at 1000◦ C for 15 minutes to drive in an approximately 500 nm deep n-type region to form the p-n junction. The dopant oxide was then removed and ring shaped top contacts (100nm Ti / 200nm Au) were evaporated and patterned. The annular photodetector mesas were then defined using photolithography and the silicon device layer was etched in a deep reactive ion etcher (DRIE, SPTS Pegasus) using the buried oxide as a selective etch stop. The SOI wafer was then bonded to a Pyrex substrate using a Waferbond HT 10.10 (Brewer Sciences), and the 400 µm thick silicon handle layer of the SOI wafer was removed using DRIE. The buried oxide was then removed with buffered HF, leaving only the 10 µm thick device embedded in the Waferbond. Next, a broad area back contact (100nm Al / 50nm Ti / 50nm Ni / 200nm Au) was deposited on the device by evaporation, and was patterned. The thin film PDs were then bonded to a flexible polyimide substrate (DuPont Kapton HN, 127 µm thick) using thermocompression bonding. The Kapton surface was planarized by spin coating a 5 µm thick polyimide film (PI 2525, HD Microsystems) onto the Kapton, followed by a cure in a N2 atmosphere at 350◦ C for two hours. A patterned metal layer for thermocompression bonding of the PDs to the flexible substrate was then vacuum deposited and patterned (100 nm Ti / 200 nm Au) on the Kapton/polyimide substrate. Thermocompression bonding between the PD contact and the metal pad on the Kapton/polyimide substrate was performed in a vacuum oven at 250◦ C with a 1 kg weight applied. After bonding, the Pyrex wafer was removed by heating the assembly to 225◦ C, mechanically debonding the temporary adhesive and enabling removal of the Pyrex wafer. The residual Waferbond was removed using trichloroethylene. To complete the flexible probe, the Kapton/polyimide substrate with the bonded Si PDs had top leads and an anti-reflection coating deposited. To isolate the top contacts, 5 µm of polyimide PI 2525 was patterned and cured as an interlayer dielectric. Evaporated leads (100nm Ti / 200nm Au) were then deposited and patterned to electrically connect the PD top contacts to the device pads. These leads were designed with a serpentine pattern to enable flexing of the leads on the substrate without damage to the leads. Silicon nitride (52 nm thick) was deposited using plasma-enhanced chemical vapor deposition (PECVD, Advanced Vacuum Vision) as a surface passivation layer and anti-reflection coating. Finally, a transparent optical aperture was opened through the Kapton/polyimide substrate by CO2 laser ablation (H-Series 20x12, Full Spectrum Laser). To avoid thermal damage to the PD, the optical aperture had a diameter of approximately 200 µm, which is smaller than the annular Si PD 750 µm inner diameter. A photograph of the completed flexible probe is shown as Fig. 2. 2.2.
Photodetector characterization
Dark current and responsivity measurements were performed on the flexible PDs bonded to the Kapton substrate using a Keithley 4200 source measurement unit (SMU) and a Newport 300W Xenon lamp with a Newport CS130 monochromator. To measure the responsivity, the PDs were illuminated in a surface normal configuration. The Xenon lamp was coupled to a 1 mm diameter optical fiber with a NA of 0.39, and the output of the fiber was focused through a microscope objective lens to illuminate the front surface of the PD. The optical wavelengths measured were from 470 nm to 600 nm in increments of 10 nm, for a total of 14 measurements. The typical illumination power at the output of the microscope objective varied slightly with the lamp output wavelength, and was approximately 150 nW. The two thin film PDs were labeled PD1 and PD2.
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Fig. 2. Image of the flexible PD probe bent cylindrically outward, bending radius of 15 mm.
PD1 had higher responsivity and lower dark current than PD2. The responsivity for each PD was averaged over 100 measurements for each wavelength. The responsivity for PD1 at zero bias was 0.20-0.34 A/W for λ = 470 – 600 nm, and for PD2 the measured responsivity was 0.10 to 0.17 A/W for the same wavelength range. The zero bias dark current (averaged over 1000 measurements) was 10 pA for PD1, and 750 pA for PD2. With a total photodiode surface area of 0.0447cm2 , the dark current density for PD1 was 0.224 nA/cm2 , and 16.8 nA/cm2 for PD2. The poor performance of PD2 is is likely due to spiking of the contact into the shallowly diffused pn junction, which would result in the observed mixed photodiode/photoconductor behavior. This poor yield can be addressed with better process uniformity control across the SOI wafer. The rest of the paper will focus on measurements performed on the higher performance PD1. The zero-bias dark current of PD1 is similar to that of previously reported thin film Si devices of the same conformation and size bonded to Pyrex and Kapton, the PD1 responsivity exceeds previously reported values for flexible PDs on Kapton [29], and is comparable to the conventional devices on rigid Pyrex substrates [29]. Thus, thin film PD1 is a high performance device, with high responsivity and low dark current. Note that the higher responsivity of PD1 can be attributed to the anti-reflection coating, which was absent in the previously reported flexible Si PDs on Kapton. Surface normal responsivity measurements were also performed on PD1 under flexed conditions. The responsivity was measured using the same responsivity test apparatus as above, except that the PD/Kapton probe was held in a series of 3D printed plastic clamps with five different radii of curvature, which flexed the PD/Kapton probe. The responsivity for the wavelength range 470-600 nm for 5 curvatures is shown in Fig. 3(a). Each data point is the average of 100 measurements at each wavelength/curvature combination, with a standard deviation of 0.61%. The PD1 dark current and current-voltage characteristic for each curvature is shown in Fig. 3(b), averaging 100 non-illuminated measurements at zero bias voltage. The zero-bias dark current for PD1 ranged from 9-12 pA for all five curvatures. All PD1 performance metrics show negligible change as a function of flexure. 3. 3.1.
Phantom testing Liquid phantoms
Two liquid phantoms were prepared to represent human breast tissue (adipose and malignant), with scattering and absorbing coefficients within the range of reported values for wavelengths 470-600 nm [34]. Nigrosin absorber (Sigma-Aldrich) was mixed in deionized water with a concentration of 5.54×10 −5 weight/volume (benign) and 1.10×10 −4 w/v (malignant). Scattering polystyrene spheres, 0.99 µm diameter (Polysciences, Inc.), were added at a concentration of
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Fig. 3. (a) Surface normal responsivity of PD1 at multiple bend radii; (b) Dark current characterization of PD1 at multiple bend radii.
2.24×10 −3 w/v (benign) and 3.54×10 −3 w/v (malignant). The absorption coefficient was calculated from spectrophotometer transmission measurements (Shimadzu UV-3600) of the phantom before adding the scattering polystyrene spheres. The scattering coefficient and anisotropy were estimated using Mie theory calculations [35]. The measured absorption and calculated scattering properties for both phantoms are shown in Fig. 4.
Fig. 4. Absorption (µa , dotted lines) and scattering (µs , solid lines) coefficients of the liquid phantoms representing adipose (red lines) and malignant (blue lines) breast tissue.
For the wavelength range 470-600 nm, the absorption coefficient ranged from 2.1 to 6.0 cm −1 , and the scattering coefficient ranged from 87 to 182 cm −1 . The anisotropy (g) ranged from 0.921 to 0.927, as calculated from Mie theory. The reduced scattering coefficient, µs ’, where µs ’=µs (1-g), ranged from 6.9 to 13.3 cm −1 . 3.2.
Measurement apparatus
To perform diffuse reflectance measurement while flexing the PD/Kapton probe at various curvatures, pairs of plastic clamps were 3D printed with five discrete radii of curvature: flat (infinite radius), 50 mm, 25 mm, 15 mm, and 10 mm. The upper section of each pair of 3D printed
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parts included a hollow volume for containing the liquid phantom during testing. The inner walls of this hollow volume were painted black to minimize sidewall reflections. The open end of the phantom container was capped with 50 µm thick transparent PET film (McMaster-Carr). The PD/Kapton probe was placed below the PET film with the detectors facing upwards, towards the phantom. The two detectors were aligned equidistant from the center of the front face of the phantom, as illustrated and photographed in Fig. 5. Illumination was provided through a 1 mm diameter optical fiber (NA=0.39) aligned to the center aperture of PD1. To perform DRS measurements, the light source fiber was aligned to the back of the PD1 aperture. For crosstalk measurements, the test was repeated with the source fiber aligned to the PD2 aperture, and the PD1 photocurrent was measured. The optical power incident on the phantom through these apertures varied with wavelength, from 0.9-2.0 µW.
Fig. 5. Left: Illustration of liquid phantom testing at 10 mm radius of curvature: (a) Top clamp, with hollow region to contain liquid phantom and curved bottom edges; (b) Transparent PET film; (c) Flexible PDs on Kapton; (d) Bottom clamp, with curvature matching the top clamp. Right: Photograph of 10 mm bending radius test apparatus.
For each different curvature under test, the total throughput optical power was measured on a power meter (Thorlabs PM100) and the fiber alignment to the PD was optimized for maximum optical throughput. This measurement was taken with a dry container before the liquid phantom was added, and the fiber to PD alignment was unchanged during all of the phantom tests for that particular curvature. The diffuse reflectance for each curvature was then measured by placing each phantom into the container separately and illuminating the PD with the fiber. To normalize for light source fluctuations and fiber-to-aperture alignment mismatch, a reflectance standard photocurrent was recorded before each phantom and curvature measurement. The reflectance standard (Labsphere, Inc. SRS-99-010) was measured with the PD probe held in another, shorter 3D printed clamp that enabled close proximity between the PD and reflectance standard. For each wavelength, 50 reflectance standard measurements were averaged. The reflectance standard was then removed and the optical throughput was measured for this configuration on a power meter (Thorlabs PM100). The scaled flat case reflectance standard data shows that the thin film PD1 DRS signal is comparable to the previously reported DRS signals from a comparable rigid thick DRS detector [20, 32, 36]. Scaling the input power to compare the thin film and typical thickness (300 µm – 500 µm thick) PD measurements on a 99% reflectance standard yields an average percent difference of 3.33% over the common experimental wavelengths of 500 nm, 530 nm, 560 nm, and 600 nm. After each reflectance measurement, the liquid phantom was replaced with a highly absorbing nigrosin solution (µa > 50cm −1 ), and a background current measurement was taken. Herein, background current is defined as a PD output current that is not generated by the DRS signal.
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Since the PD is illuminated from the back side, and the light passes through the PD aperture, back illumination current can be generated by absorption at the inner edges of the PD. Also, reflection from the PET film can cause background current, and the photodiode dark current is considered a background current. To calculate the diffuse reflectance at each wavelength, the measured background current was subtracted from the corresponding reflectance measurement to assess only the photocurrent due to diffuse reflectance. This value was then divided by the optical throughput scaled by the ratio of the reflectance standard photocurrent to the reflectance standard throughput, since the puck and curved measurements had slightly different optical throughput values as the reflectance standard test was not on a curved surface. This expression for reflectance can be written as: R (λ, ρ, P) =
i phantom (λ, ρ, P) − i backillum (λ, ρ) i h d (λ) I0 (λ, ρ) Ii0,sst at anndar (λ) dar d
(1)
where R (λ, ρ, P) is the diffuse reflectance at wavelength λ, curvature ρ, and liquid phantom P; i phantom (λ, ρ, P) and i backillum (λ, ρ) are the photocurrent measurements of the liquid phantom reflectance and back illumination, respectively; I0 (λ, ρ) is the recorded optical throughput at wavelength λ, and curvature ρ; and i standard (λ) /I0,standard (λ) is the previously calculated ratio of reflectance standard photocurrent to optical throughput at wavelength λ. 3.3.
Measurement results
The reflectance data for fourteen wavelengths between 470 nm and 600 nm with 10 nm spacing for both phantoms at three of the five measured curvatures is shown in Fig. 6. The 50 mm and 25 mm curvatures are omitted for clarity; the reflectance values from these curvatures were similar to the flat test case. Each data point shows the average and standard deviation of 75 photocurrent measurements, in three successive spectral scans of 25 measurements each. Three successive scans were conducted to assess any bias due to phantom settling, which was negligible.
Fig. 6. DRS reflectance versus wavelength for both benign and malignant phantoms at fourteen wavelengths and three radii of curvature.
The data shows an increasing reflectance signal with increasing curvature (decreasing radius) for both phantoms. The increase in DRS signal due to curvature ranged from 5% to 8% for the benign phantom and 5% to 12% for the malignant phantom. This increase in reflectance is
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expected for increased curvature, since, in the diffusion approximation, the reflected intensity increases as the source-detector distance decreases [5]. In an inwardly curved state, the total path length within the bulk liquid phantom from optical source (PD aperture) to detector (PD surface) decreases. As a result, the collected diffuse reflectance optical power at the PD surface increases. Due to the higher absorption and higher scattering of the malignant phantom, it has a larger increase in reflected signal for the same reduction in path length compared to the benign phantom. The crosstalk was measured by illuminating the aperture of PD2 and measuring the PD1 photocurrent. The PD1 aperture was not illuminated. The measured crosstalk is shown in Fig. 7. The malignant phantom produced a significantly lower crosstalk signal, which is expected due to the higher absorption of the malignant phantom. Due to this high absorption, the crosstalk photocurrents were below the noise floor of the photodiode and could not be measured.
Fig. 7. Crosstalk reflectance for the benign phantom. All five radii of curvature were measured.
As a percentage of the DRS signal, the maximum crosstalk was 2.3% (benign phantom, λ=470 nm, 10 mm radius of curvature). The crosstalk signal shows a stronger dependence on curvature than the DRS signal, increasing up to 40% from a flat surface to a 10 mm radius of curvature. The decrease in path length due to increasing curvature is more pronounced in the case of the crosstalk detector than the DRS detector. Due to the larger reduction in path length, the percent change in crosstalk signal is higher when the phantom surface is curved. The crosstalk to DRS signal ratio remained below 3% at all curvatures and wavelengths. This crosstalk is comparable to the previously reported 6-9% maximum simulated crosstalk percent values for a rigid thick DRS system with similar probe geometry [32]. 4.
Zemax simulations
To validate the experimental results and toward a look-up table for coefficient extraction, Monte Carlo ray tracing simulations were performed using Radiant ZEMAX commercial software. The simulation geometry for each of the five curvatures were modeled. The side and back walls of the container were simulated as perfect absorbers, and the front wall was modeled as a 50 µm thick transparent PET film. The two PDs each had a 200 µm diameter aperture in the annulus. Rays were launched through the center aperture of the illuminated PD, incident on the phantom. The bulk scattering model in ZEMAX, with Henyey-Greenstein statistics for scattering angles,
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was used to simulate diffuse reflectance. All rays hitting each PD were summed to simulate the reflectance incident on that PD. Each curvature and wavelength were modeled with 50 million rays. The optical power recorded at the illuminated PD surface was defined as the DRS signal, and the power at the other, unilluminated PD was the crosstalk signal. The diffuse reflectance was calculated by dividing the optical power at each detector by the reflectance optical power from a 99% reflectance standard, where the reflectance standard was modeled separately using a Lambertian scattering surface. 4.1.
Comparison of theory to experiment
The simulation and experimental diffuse results for the two extreme curvature cases, flat and 10 mm, are compared in Fig. 8, and the crosstalk at five curvatures in Fig. 9, for all fourteen measured wavelengths.
Fig. 8. Experimental DRS signal (malignant phantom: red open circles, benign phantom: black open triangles) compared to Monte Carlo simulations (malignant phantom: red filled circles, benign phantom: black filled triangles) versus wavelength for the (a) flat test case and (b) 10 mm radius of curvature test case.
The mean percent error in the diffuse reflectance between simulation and experiment is 2.3% for the benign phantom and 1.23% for malignant phantom. The mean percent error in the crosstalk signal for the benign phantom was 4.85%. The main source of error was likely inaccuracy in the thickness of the PET film. For the malignant phantom, the low values of the crosstalk photocurrent were at or near the noise floor of the PD, which precluded measurement of the reflectance. The ability to differentiate between benign and malignant tissue in a DRS system can be evaluated by the signal contrast, which can be written as: S (λ) =
M (λ) − B (λ) B (λ)
(2)
where M (λ) and B (λ) are the malignant and benign diffuse reflectance signals, respectively, and S (λ) is the signal contrast. The signal contrast for the flat phantom measurements ranged from +4.0% at 570 nm to +13.6% at 470 nm, and the reflectance from the malignant phantom was higher than that of the benign phantom in all cases. The signal contrast at all fourteen wavelengths and five radii of curvature is shown in Fig. 10. The lowest contrast occurs at the wavelengths corresponding with the absorption peak of nigrosin, at approximately 560 nm. At all curvatures, the signal contrast remains positive. With
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Fig. 9. Experimental crosstalk reflectance (open triangles) compared to Monte Carlo simulations (filled triangles) for the benign phantom at five radii of curvature.
Fig. 10. Signal contrast between benign and malignant phantom reflectance at flat and flexed conditions.
increasing curvature, the average contrast across all wavelengths increases from 6.2% at the flat test case to 10.0% at a radius of 10 mm. The curvature has an approximately linear effect upon the diffuse reflectance. In Fig. 11, the diffuse reflectance experimental and simulated results are shown as a function of curvature (where curvature is the inverse of surface radius) for the benign and malignant phantoms for λ = 470, 500, 530, 580, and 600 nm for all five measured curvatures. Although the measured results are more scattered, the results from Zemax simulations suggest a linear dependence on curvature (average R2 = 0.992), shown as dashed lines. The phantom reflectance signal increased an average of 8.0% in simulation and 7.9% in measurement as the curvature increased from flat to a radius of 10 mm. In the experimental data, the increase in reflectance with increasing curvature was higher for the malignant phantom than
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Fig. 11. DRS signal (open circles) of (a) benign phantom; and (b) malignant phantom, compared to Monte Carlo simulations (filled circles) versus curvature (inverse of radius) at five wavelengths. Linear regression (dashed lines) shown for simulated data.
for the benign phantom, resulting in an increase in contrast with increasing curvature. 5.
Conclusions
We have designed, fabricated, and tested thin film, flexible single crystalline Si PDs on a flexible Kapton substrate. The PD1 responsivity and dark current performance was equal to that of conventional rigid PDs. The PD DRS performance was measured on two liquid phantoms, representing benign and malignant breast tissue. Measurements were taken over 5 curvatures, ranging from flat to 10 mm radius of curvature. The average error between the experimental results and Monte Carlo simulations for the DRS measurements was 1.8%. The flexible PDs presented herein demonstrate that flexible semiconductor probes can be used for DRS measurements, which may enable more accurate analysis of non-planar tissues. Increasing the phantom curvature from flat to a 10 mm radius of curvature increased the DRS signal by an average of 7.9%. The crosstalk to DRS signal ratio remained below 3% at all curvatures and wavelengths. Funding This research was supported by the National Institutes of Health through the Bioengineering Research Partnership award 1 RO1 EB011574-01A1. Additionally, this work was performed in part at the Duke University Shared Materials Instrumentation Facility (SMIF), a member of the North Carolina Research Triangle Nanotechnology Network (RTNN), which is supported by the National Science Foundation (Grant ECCS-1542015) as part of the National Nanotechnology Coordinated Infrastructure (NNCI). Acknowledgments The authors would like to thank the staff of Duke University Shared Materials Instrumentation Facility for their technical assistance, and Duke University and the Department of Electrical and Computer Engineering for their support.
