2.1.

Instrumentation

A schematic of the benchtop SFDI instrument is shown in Figure 1 and is described here briefly for completeness. This realization is similar to those reported by Sibai et al.¹⁷ The custom-built SFDI system comprises a xenon arc lamp that provides broadband light via a liquid light guide (Sutter Instruments) to a digital micromirror array projector module (CEL5500, DLI, Texas) to modulate the light before it is focused onto the sample at the object plane.

Fig. 1

(a) Schematic of the benchtop SFDI system and (b) photograph of its physical implementation.

On the detection side, light is collected and passed through a visible-range liquid-crystal tunable filter (LCTF: varispec-07-02, Perkin Elmer. Inc., Waltham, Massachusetts) before it is registered on a 14-bit sCMOS camera (Pixelfly, PCO AG, Kelheim, Germany). Two-by-two binning is chosen for all image acquisitions to enhance signal to noise under acquisition times of interest (50 to 1000 ms). Seven patterns of light are projected at spatial frequencies of 0, 0.1, 0.3, 0.5, 0.6, 0.8, and $1{\mathrm{mm}}^{-1}$. Control of the projected light as well as the image acquisition are done in LabVIEW (National Instruments), while a MATLAB script (MathWorks) quantifies the optical properties across the acquired FOV.

To modify the benchtop SFDI approach for clinical neurosurgery, a commercial surgical microscope stand (S1, Zeiss) was modified to accept light from the projector module and focus it on the image plane (Fig. 2). The detection subassembly, consisting of the same LCTF and sCMOS camera components, was attached to the side port of the microscope head via a custom optical adapter (TrueTex) that ensured that the magnification and image FOV were consistent with the corresponding ocular view.

Fig. 2

(a) Schematic of the clinical SFDI system and (b) photograph of the system. The Xe Arc lamp (not shown) is mounted to the base of the microscope stand.

2.2.

Tissue-Simulating Phantom Studies

Diffuse reflectance measurements were acquired with both the benchtop and clinical SFDI setups on tissue-simulating phantoms covering a wide range of absorption and transport scattering coefficients, produced by varying concentrations of India Ink (VWR) and Intralipid (Intralipid 20%, Patterson Veterinary Supply), respectively, spanning the range of optical properties found in brain tissue.¹⁸ The absorption coefficient was varied from 0.001 to $0.3{\mathrm{mm}}^{-1}$ in $3\times$ intervals for a total of six values, while scattering properties were varied from 0.4 to $2.0{\mathrm{mm}}^{-1}$ in intervals of $0.2{\mathrm{mm}}^{-1}$ for a total of nine values, creating a total of 54 combinations of absorption and scattering properties calculated at 670 nm.

For the modified SFDI system to measure optical properties in vivo on complex surfaces accurately, variations in surface topology that translate into differences in surface-to-detector distances across the FOV were corrected. Two experiments were conducted, accordingly: (1) to determine accuracy of optical property estimation and establish a profilometry calibration for the clinical system and (2) to evaluate and compare spatial resolution across the FOV of the benchtop and clinical SFDI systems. For the first test, a $12\mathrm{cm}\times 12\mathrm{cm}\times 3\mathrm{cm}$ deep dish was filled with each phantom material [Fig. 3(a)]. Under white light illumination, diffuse reflectance images from 450 to 710 nm were acquired at 5-nm intervals, with exposure times of 250 ms. The method for phase profilometry is described elsewhere^19–21 and briefly summarized here. By projecting an intensity sine wave with three different phase shifts and recovering the fringe phase at each pixel with the phantom positioned in multiple orientations and at different heights, we estimated the linear relationship between the inverse of sample height variation and the inverse of phase variation. For the clinical system, height was varied over a 3-cm range using a $z$-translation stage, while tilt angle was varied from $-20\mathrm{deg}$ to $+20\mathrm{deg}$ by rotating the head of the clinical microscope to determine the height-dependent and angle-dependent intensity calibrations. In total, 155 images were acquired for each spatial frequency for a total integration time of 46.8 s. Images were analyzed using a rapid two-frequency lookup table (LUT) (see Sec. 2.3 below).

Fig. 3

Diagrams of phantom experiments evaluated. (a) Experiment 1 involved homogeneous phantoms in a single well and (b) experiment 2 involved a heterogeneous phantom consisting of five separate wells filled with solutions of different optical properties.

For the second test, five different solutions were selected to represent a range of absorption and scattering parameters. Solutions were placed in a container that had five sections, each being $3\mathrm{cm}\times 1\mathrm{cm}\times 2\mathrm{cm}$ separated by 0.17-mm glass cover slips [Fig. 3(b)]. Acquisition timing remained at 250 ms. In this test, the heterogeneous phantom was imaged in multiple orientations to assess the system’s ability to return optical properties under different sample-detector distances similar to more complex clinical settings. For the benchtop system, the phantom was translated in the $z$-dimension and for the clinical system, the phantom was translated in the $z$-dimension and the head of the microscope was tilted. Using a two-frequency LUT, edge-response functions were averaged for the different phantom orientations and compared between benchtop and clinical setups.

2.3.

Lookup Table Method

To build the LUT, a reference set of phantoms was created to measure diffuse reflectance at two spatial frequencies of 0 and $1{\mathrm{mm}}^{-1}$. Absorption values were varied from 0.001 to $0.5{\mathrm{mm}}^{-1}$ in $2\times$ intervals for a total of 10 absorption values, while the reduced scattering coefficients were varied from 0.3 to $2.1{\mathrm{mm}}^{-1}$ in intervals of $0.2{\mathrm{mm}}^{-1}$, for a total of 10 scattering values, to create a set of 100 phantoms with varying optical property combinations. The measured diffuse reflectance was used to construct a two-frequency LUT.^14,18 A MATLAB script based on cubic spline interpolation was formed to infer optical properties from newly acquired diffuse reflectance images based on this reference phantom dataset.

2.4.

In Vivo Human Experiments

To test performance in an operative setting, the clinical SFDI system was used under an Internal Review Board-approved protocol at Dartmouth Hitchcock Medical Center to estimate absorption and scattering parameters of exposed cortex, intraoperatively. In the surgical setting, we limited acquisitions to reflectance images at 670 and 710 nm, as the optical properties at these wavelengths are needed to estimate fluorescence at depth.^11,12

Three images with the spatial frequency of $0.10{\mathrm{mm}}^{-1}$ were acquired for profilometry at each wavelength. For acquisition of diffuse reflectance values, $R_d(f=1{\mathrm{mm}}^{-1})$ and $R_d(f=0{\mathrm{mm}}^{-1})$, three phases of $f=1{\mathrm{mm}}^{-1}$, one uniform illumination ($f=0{\mathrm{mm}}^{-1}$), and one dark image were acquired at each wavelength (totaling 16 image acquisitions, requiring $\sim 4\mathrm{s}$ total). Images were processed for height and optical property estimation with profilometry correction.

3.1.

Optical Property Estimation Accuracy

As the phantom solutions were homogeneous over the FOV, the estimated optical properties were averaged over a region of interest that was chosen as the area over which light was projected within the acquired image. Table 1 summarizes the errors in estimating the absorption and transport reduced scattering coefficients for benchtop and clinical systems averaging over the set of 54 phantoms. Even with reduced light intensity, the errors were comparable between the benchtop and clinical systems.

Table 1

Summary of errors in absorption and reduced scattering coefficients between benchtop and the clinical SFDI systems.

3.2.

Spatial Resolution

Spatial resolution was evaluated with the phantom filled with five solutions varying from low scattering (${\mu }_s^{\prime }=0.4{\mathrm{mm}}^{-1}$) and absorption (${\mu }_a=0.001{\mathrm{mm}}^{-1}$) to high scattering (${\mu }_s^{\prime }=2.4{\mathrm{mm}}^{-1}$) and absorption (${\mu }_a=0.3{\mathrm{mm}}^{-1}$). Figure 4 (a) absorption and (b) reduced scattering maps acquired with the clinical system, as well as (c) a photograph of the heterogeneous phantom, itself. From left to right, wells contained solutions with the following properties; ${\mu }_s^{\prime }=2.4{\mathrm{mm}}^{-1}$ and ${\mu }_a=0.3{\mathrm{mm}}^{-1}$, ${\mu }_s^{\prime }=0.4{\mathrm{mm}}^{-1}$ and ${\mu }_a=0.001{\mathrm{mm}}^{-1}$, ${\mu }_s^{\prime }=1.2{\mathrm{mm}}^{-1}$ and ${\mu }_a=0.03{\mathrm{mm}}^{-1}$, ${\mu }_s^{\prime }=2.4{\mathrm{mm}}^{-1}$ and ${\mu }_a=0.001{\mathrm{mm}}^{-1}$, and ${\mu }_s^{\prime }=0.4{\mathrm{mm}}^{-1}$ and ${\mu }_a=0.3{\mathrm{mm}}^{-1}$. For the benchtop system, a total of four heights were tested over a range of 3 cm, and for the clinical system, four front and side tilt angles from $-20\mathrm{deg}$ to $+20\mathrm{deg}$ were evaluated including a horizontal orientation at each of four heights over a range of 3 cm for a total of 45 orientations. Edge response functions were measured and averaged for the orientations of the benchtop and clinical setups and are shown in Figs. 5(a) and 5(b) for absorption and scattering, respectively. Here, spatial resolution was defined as the distance at which the edge-response contrast was reduced by 90%, as described elsewhere.¹³ The edge resolutions from the benchtop and clinical systems were similar for both absorption and scattering maps. The spatial resolution of scattering estimates for the transition between high- and low absorption wells was 0.51 and 0.53 mm for the benchtop and clinical systems, respectively, and similarly, 0.51 and 0.54 mm for low and medium absorption wells, respectively. The corresponding spatial resolutions in absorption were 0.63 and 0.65 mm for transitions between high and low, and 0.65 and 0.70 mm between low and medium absorption properties. These results are shown in Table 2.

Fig. 4

(a) Absorption and (b) reduced scattering maps with corresponding scale bars of the heterogeneous phantom acquired at 670 nm using spatial frequency of $1{\mathrm{mm}}^{-1}$ with the clinical SFDI system. (c) Photograph of the heterogeneous phantom.

Fig. 5

Edge response functions for the benchtop and clinical SFDI systems for (a) absorption and (b) reduced scattering maps of the heterogeneous phantom.

Table 2

Summary of spatial resolution estimates for absorption and reduced scattering coefficients from the benchtop and clinical SFDI systems.

3.3.

In Vivo Imaging

Processing the data into 2-D maps of profilometry corrected absorption and transport scattering values involved $\sim 1\min$ of computation time. Figure 6 shows an example of the absorption and scattering property maps of the cortical surface. Blood vessels in the field correspond to higher relative absorption values and lower relative values in scattering, as expected. Dashed lines in the reduced scattering map correspond to areas of specular reflection. Further discussion of the limitations with this current setup will be explored in Sec. 4.

Fig. 6

(a) Absorption and (b) reduced scattering maps measured at 670 nm of the cortical surface during a brain tumor resection surgery. $\text{Scale bar}=1\mathrm{cm}$. Dashed lines in reduced scattering map highlight areas of specular reflection.