2.1.

System Design and Overview

The key design parameters of the system developed were: 1. the capability of simultaneous video-rate imaging in NIR emission (fluorescence signals), NIR excitation (intrinsic signal), and visible wavelength bands; 2. the capability of providing image normalization; 3. operation at long working distances, large field of view, high detection sensitivity, and high resolution; and 4. a highly optimized detection sensitivity in all spectral bands considered.

To reach these functional specifications, a three-channel detection system was developed as shown in Fig. 1 . The selection of three cameras is based on the need to use different camera specifications for color and near-infared (NIR) imaging while providing simultaneous readings of all the necessary measurements. The system specifications, as described in Sec. 3, are shown in Table 1 . Illumination was provided by a $750\text{-}\mathrm{nm}$ cw laser diode (BWF2-750-0, B and W Tek, Newark, Delaware; maximum power $300\mathrm{mW}$ ) for fluorochrome excitation, and a $250\text{-}\mathrm{W}$ halogen lamp (KL-2500 LCD, Edmund Optics, Barrington, New Jersey) for white-light color imaging. The system has the capacity to operate in epi-illumination and transillumination modes. Epi-illumination imaging was performed using a front illuminating expanded beam by guiding laser light through a multimode fiber ( $200\text{-}\mu \mathrm{m}$ core/0.22 NA) to a collimator and a diffuser (F260SMA-B, ED1-S20, Thorlabs, Newton, New Jersey) for beam expansion and uniform illumination. White light was also coupled with a fiber optic bundle and homogenized by a beam expander (EG8546, Edmund Optics). Transillumination was accomplished by guiding laser light through a multimode fiber that was scanned on the bottom surface of the target, through an antireflection-coated glass window, using the $x\text{-}y$ translation stage (MAXY4009Q1-S4, Velmex, Bloomfield, New York). Transillumination measurements were achieved using a diffuser in front of the multimode fiber to achieve more homogeneous back-illumination. $x$ and $y$ axes were scanned with four steps of $35\mathrm{mm}$ and three steps of $33.75\mathrm{mm}$ , respectively.

Fig. 1

(a) Schematic of the imaging system developed to operate in real-time (video) mode and implementing concurrent fluorescence, attenuation (intrinsic), and color measurements. All optical path lengths from the lens to the cameras are equal (not plotted in scale on the drawing). Three CCDs and the linear stage are under computer control. (b) Imaging windows display and user interface for operational guidance. Measurements at excitation (intrinsic) and emission wavelength, and corresponding normalized fluorescence image are shown. (c) Spectra of light transmittance in each imaging channel. Clear separation of wavelength for visible and near-infrared is originated from the combination of dichroic mirror and bandpass filters.

Table 1

Imaging system specifications.

Light from the surgical field was collected using a zoom lens (NT58-240, Edmund Optics). Linear polarizers were employed for minimizing the effects of glare and specular reflection in the attenuation and color images by enabling cross-polarized measurements for white and excitation (intrinsic) light measurements. The additional advantage of polarization is that it samples deeper into tissue compared to nonpolarized imaging, which is an important parameter in getting a more accurate reading on tissue attenuation in epi-illumination mode. However, when switching to imaging through the cross-polarizers, lower resolution is generally achieved. This feature is used selectively when specular reflection indeed occurs.

The primary image of the object formed by the zoom lens falls at the focal plane of each relay lens group (MAP10100100-A1, Thorlabs) of three imaging channels. The setup employs two dichroic mirrors (700 DCXXR, z 750 RDC, Chroma, Rockingham, Vermont) that splits light by excitation, emission, and visible wavelengths into three imaging channels. The final images are formed on each chip surface of a $12\text{-bit}$ grayscale charge-coupled device (CCD) camera with 20% quantum efficiency in the NIR area and a $24\text{-bit}$ color CCD camera (PCO AG, Donaupark, Kelheim, Germany) by each relay lens group. This relay lens group has the same object and image distance, so it does not change the magnification. Its purpose is the delivery of the primary image from the lens to each CCD chip, while allowing the intersection of dichroic mirrors, filters, and potentially other imaging components. All optical pathlengths from the lens to the cameras are equal (not plotted in scale in Fig. 1). For accurate wavelength-band selection and cross talk elimination, bandpass filters (z $750∕10$ , HQ $795∕50$ , Chroma) were used in front of CCD cameras in excitation (intrinsic) and emission imaging channels. The transmission spectrum of each imaging channel, which is induced by the combination of dichroic mirrors and the bandpass filter, can be seen in Fig. 1c. For implementing a portable self-contained system, all optical components were tightly mounted in custom coupler/holders via C-mount optics after exact optical alignment.

The control software and user interface for surgical guidance was developed under the LabVIEW (National Instruments, Austin, Texas) and Matlab (The MathWorks, Natick, Massachusetts) environments. In continuous acquisition mode, NIR excitation (intrinsic), NIR emission (fluorescence), and anatomical color image can be displayed simultaneously in real time and refreshed continuously at a rate set by the camera exposure time using two computers. In a single snap acquisition mode, when a set of images is captured, then instantly the normalized fluorescence image is calculated and displayed, as shown in Fig. 1b. The user interface provides complete control on the imaging system including camera gain, exposure time, number of binning, image display, and image archiving.

2.2.

Measurement of Optical Resolution

To measure the spatial resolution of the system, the image of a standard three-bar resolving power test target (USAF-1951, NT53-714, Edmund Optics) was taken using the visible wavelength channel and white-light illumination. This target contains numbered bar groups, each consisting of three vertical and three horizontal bar patterns. The ratio of the spatial frequency of an element to the one of the next higher element is $1:2^{1∕6}$ . After extraction of cross-sectional intensity profiles from each horizontal and vertical bar element, the points of the contrast transfer function (CTF) are obtained with the contrast equation, i.e.,

2.3.

Phantoms

To examine the metrics of fluorescence quantification correction as it relates to spatial variations of excitation light in tissue, we constructed a phantom consisting of five tubes containing the same fluorochrome concentration but different concentrations of absorbing dye, resulting in variation in the underlying optical properties. The five tubes were made of translucent $\sim 3\text{-}\mathrm{mm}$ -diam plastic and contained $400\mathrm{nM}$ of Alexa Fluor 750 dye (Molecular Probes, Eugene, Oregon) and 1% Intralipid solution. In addition, the tubes contained India ink of varying concentration from $50\text{to}250\mathrm{ppm}$ with $50\text{-}\mathrm{ppm}$ increments to impart differences in attenuation between tubes. The imaging chamber had an open upper surface, whereas its bottom surface was a glass window, which was antireflection coated for the NIR to facilitate both epi-illumination and transillumination measurements. The tubes were placed in an imaging chamber and positioned parallel to each other, allowing a $6\text{-}\mathrm{mm}$ gap between consecutive tubes. The chamber was filled with 1% Intralipid solution (Sigma, Saint, Louis, Missouri) and $50\text{-}\mathrm{ppm}$ India ink (yielding optical properties of absorption coefficient ${\mu }_a=0.4{\mathrm{cm}}^{-1}$ and reduced scattering coefficient ${\mu }_s^{\prime }=8{\mathrm{cm}}^{-1}$ , characteristic of small animals). The thickness of the fluid phantom was $1\mathrm{cm}$ and the tops of all the tubes were positioned $2\mathrm{mm}$ below the top surface of the fluid solution. A similar phantom containing a single tube was also constructed and was employed to examine the effects of background absorption variation. A single tube was immersed in the Intralipid solution, and the absorption of the Intralipid solution was altered by titrating India ink $\left(50\text{to}250\mathrm{ppm}\right)$ in the solution that surrounded the tissue.

2.4.

Tissue Imaging

Euthanized mice (C57BL/6J, $\mathrm{n}=4$ ) were employed to showcase imaging feasibility simulating intraoperative imaging. After euthanasia, the abdominal sides of the mouse were surgically exposed, and the lumbar nodes beside the inferior vena cava were visualized by removing the surrounding peritonea and fat. Imaging of the nodes was selected because it allows for visualizing well defined volumes. $8\mu \mathrm{M}$ of Alexa Fluor 750 with $300\mathrm{ppm}$ of India ink was injected into one lymph node and the same fluorescent solution, and without ink into another lymph node ( $10\mu \mathrm{l}$ for each node), using a 30-gauge needle.

2.5.

Measurements and Normalization

The correction scheme considered assumes a raw fluorescence image $I$ in epi-illumination or transillumination modes, i.e.,

3.1.

Resolution

The spatial resolution of the implemented system was determined utilizing the USAF-1951 pattern target image and contrast transfer function (CTF) as shown in Fig. 2 . The test pattern image consisting of repeated gratings is obtained at visible wavelength channels with the highest zoom setting [Fig. 2a]. For every different grating size in the field of view, we calculated the CTF parameter using Eq. 1. Figures 2b and 2c illustrate the CTF values (vertical axis) versus the different grating frequencies (horizontal axis).

Fig. 2

Spatial resolution of the implemented imaging system. (a) Color video image of a U.S. Air Force 1951 standard resolution target was taken. (b) and (c) Horizontal and vertical contrast transfer function are obtained by determining the specific contrast produced by bar group patterns of various spatial frequencies [see Eq. 1]. By using a CTF fitting curve [solid black lines in (b) and (c)] the transverse resolution of the system was measured as 68.21 and $36\mu \mathrm{m}$ by the Rayleigh criterion (26.4% contrast) and the Sparrow criterion (0% contrast), respectively. Similarly, the lateral resolution of the system was determined to be 66.58 and $34\mu \mathrm{m}$ by the Rayleigh criterion and the Sparrow criterion, respectively. Use of a different front lens can increase the resolution but limit the field of view.

Figures 2b and 2c show the contrast transfer function in the image plane as a function of the spatial frequency $f$ along the horizontal and vertical directions, respectively. To quantify the transverse or lateral resolution of the system, we applied the Rayleigh criterion and the Sparrow criterion to the profiles measured from the phantom by the system developed. The Rayleigh criterion defines resolution as the distance between two point objects, where the Airy disk minimum of the image of one object coincides with the maximum of the Airy disk of the other object. Typically at this limit the contrast between the maximum and minimum intensity observed between the two objects becomes about 26.4%. The Sparrow criterion defines the resolution of an optical system as being equal to the cut-off distance, i.e., the distance below which two objects cannot be distinguished by means of contrast. Since the line pair corresponding to a contrast value of 26.4% was 7.33 and $7.51\mathrm{lp}∕\mathrm{mm}$ in the horizontal and vertical CTF curves, respectively, by taking the half-width between two line pairs we determined the transverse resolution of the system to be $68.21\mu \mathrm{m}$ and the lateral resolution to be $66.58\mu \mathrm{m}$ by the Rayleigh criterion. Correspondingly, the transverse and the lateral resolutions of the system by the Sparrow criterion were found to be 36 and $34\mu \mathrm{m}$ , respectively. Since three imaging channels are constructed with identical optical components and pathways, the resolution characteristics were similar in all channels with minor variations $(<5\mu \mathrm{m})$ , largely due to wavelength-dependent characteristics of the optical components utilized.

3.2.

Epi-Illumination Imaging

Figure 3 shows the results obtained in epi-illumination from the fluid phantom containing the five tubes of identical fluorochrome concentration but of varying attenuation, simulating tissue optical heterogeneity. Figure 3a depicts the image obtained at the excitation wavelength (intrinsic image). A characteristic pattern of varying light absorption by the tubes is visible, in an increasing fashion from tube 1 to tube 5, due to the increasing concentration of the absorber. The uncorrected fluorescence image in Fig. 3b shows the dependency of the fluorescence signal on the varying absorption in the tubes, and demonstrates correspondingly a decreasing fluorescence signal with increasing tube attenuation. In this case, while all the tubes have the same fluorescence concentration, the image reports varying fluorescence intensity that could be erroneously interpreted as changes in the amount of fluorescence present in the tubes. In contrast, the normalized fluorescence image shown in Fig. 3c correctly depicts a similar fluorescence strength from all tubes by correcting the effects of background attenuation. In addition, it shows a better delineation of the tubes compared to the uncorrected images. Finally, the images demonstrate very low specula reflection, as it is successfully eliminated by means of the cross-polarized linear polarizers employed in the system.

Fig. 3

Epi-illumination imaging performance from a phantom of five fluorescence tubes immersed in a diffusive fluid as a function of absorption variation in the tubes shown. (a) Attenuation (intrinsic) image at the excitation wavelength, (b) the fluorescence image, and (c) the corresponding normalized fluorescence image. All images were acquired simultaneously. (d) Intensity strength ratio for the five tubes imaged, for normalized and conventional fluorescence measurements.

To quantitatively determine the differences between non-normalized and normalized images, the average photon counts in all tubes were calculated over identical regions of interest during image post-processing. The region of interest was chosen as an area spanning the length of the tube and a width of $30\text{pixels}$ , with the center being the center of the tube. The ratio graph in Fig. 3d indicates that the quantification performance of the normalized fluorescence image is significantly better than in the standard epi-illumination image, the latter showing signal variations as high as 25%.

Figure 4 correspondingly depicts the quantification results between corrected and uncorrected fluorescence imaging when the absorption of the background was changed. As can be seen, changes in the background also affect the signal recorded, as expected, and in this case again the corrected approach yields better insensitivity to the variation of light attenuation.

Fig. 4

Epi-illumination imaging performance from a single tube phantom as a function of absorption variation in the diffusive medium that surrounds the tube. Corrected images show again better insensitivity to the effects of increasing background attenuation compared to uncorrected images.

3.3.

Transillumination Imaging

Figure 5 depicts the transillumination imaging results obtained from the same phantom as in Fig. 3. In transillumination, the light propagates through the entire volume of the tubes. This generally leads to more prominent variations of signal intensity compared to epi-illumination, since a larger optically heterogeneous volume is sampled. Figures 5a and 5b depict the raw data obtained in the excitation and emission wavelengths, respectively. The fluorescence signal intensity varies significantly in the uncorrected image. In contrast, Fig. 5c depicts the normalized image showing a significantly more homogenous response from the fluorescent tubes. Finally, the chart shown in Fig. 5d summarizes the imaging results. It can be observed that the quantification performance of normalized transillumination imaging is significantly better than the uncorrected transillumination image. Especially for absorption values up to $3\times$ of the reference sample, the error of normalized imaging is less than 5% while the corresponding error for uncorrected transillumination images reaches 24%.

Fig. 5

Transillumination imaging with varying background optical properties. (a) Attenuation (intrinsic) image at the excitation wavelength, (b) the fluorescence image, and (c) the corresponding normalized fluorescence image. (d) Intensity strength ratio for the five tubes imaged, for normalized and conventional fluorescence measurements. The normalized transillumination image reveals the robustness on optical property variation, while the standard transillumination image reports significant error as the concentration of India ink is increased from tube 1 to tube 5.

3.4.

Tissue Imaging

While Figs. 3, 4, 5 demonstrate general performance metrics obtained from measurements on phantoms, Fig. 6 depicts epi-illumination images obtained from a mouse after euthanasia and surgical exposure of the abdominal area. In this case, two lymph nodes were injected with solutions of identical fluorochrome concentration but different absorption to simulate varying tissue optical properties, as described in Sec. 2. This model could be relevant to real-time pathology studies, where nodes are sampled for achieving real-time intraoperative pathology, but it is more generally used here to obtain performance metrics using tissues.

Fig. 6

Postmortem imaging of the surgically exposed mouse abdominal area in epi-illumination mode. (a) Color image showing the lumbar lymph nodes (single-line arrows) around the inferior vena cava (double arrow). (b) Image obtained at the excitation wavelength (intrinsic image) depicts absorption differences between the two lymph nodes due to the differential injection of an absorber. (c) Conventional fluorescence image showing low signal intensity from the lymph nodes compared to bright background signals. (d) Normalized image showing markedly improved fluorescence quantification, correctly resolving the underlying fluorescence activity in the nodes. (Color online only.)

Figure 6a shows a color image obtained from the field of view using the color video camera. This view closely relates to the surgeon’s field of view and it is important for reference purposes. The inferior vena cava (double arrow) and two lumbar nodes (single-line arrows) are seen on the color figure. Figures 6c and 6d depict the attenuation (intrinsic), fluorescence, and normalized fluorescence images, respectively. Although the same concentration of fluorochrome was injected in both lymph nodes, the right lymph node (dotted line arrow) was darker (i.e., more attenuating) compared to the left node (solid line arrow), as identified in Fig. 6b. Figure 6c is the uncorrected fluorescence image that shows fluorescence activity from areas surrounding the node. This behavior has been previously documented²⁰ and is likely due to escaping photons from the node toward the lower attenuating surrounding tissue. From an imaging standpoint this is an erroneous result, as it indicates that the fluorescence dye was injected in the tissue surrounding the node. Conversely, the normalized image of Fig. 6d depicts a significantly more accurate picture of fluorescence biodistribution and correctly resolves the nodes as the areas of fluorescence origin.

These images further confirm the observations of Figs. 3, 4, 5, where the measured fluorescence intensity did not depend only on the fluorochrome concentration but also on the heterogenous attenuation of light in the underlying tissue. It also further showcases that the spatial appearance and contrast of the image also improves when the fluorescence image is corrected for the underlying optical properties. In that respect, correction can lead to a more accurate and robust performance.