In Vivo Fluorescence Hyperspectral Imaging of Oral Neoplasia

In Vivo Fluorescence Hyperspectral Imaging of Oral Neoplasia Darren Roblyer a, Cristina Kurachi b, Ann M. Gillenwater c, Rebecca Richards-Kortum a a Department of Bioengineering, Rice University, 6100 Main St., Houston, TX 77005; b Institute of Physics of São Carlos, University of São Paulo, São Paulo, Brazil; c Department of Head and Neck Surgery, University of Texas M.D. Anderson Cancer Center, 1515 Holcombe Boulevard, Unit 441 Houston, Texas 77030 ABSTRACT A hyperspectral imaging system using a liquid-crystal tunable filter (LCTF) was constructed for the purpose of in vivo optical imaging of oral neoplasia. The system operates in fluorescence mode and has the dual capability of capturing high quality widefield images and detecting fluorescence emission spectra from arbitrary locations within the captured field of view (FOV). The system was calibrated and evaluated for spectral resolution and accuracy. In vivo hyperspectral images were obtained from two normal volunteers and two patients with confirmed oral malignancy. Normal volunteer measurements revealed differences in intensity and lineshape of spectra between different anatomic locations, but intensity and lineshape were similar between different measurement sites from the same anatomic location. Measurements from normal and neoplastic areas of two patients with previously confirmed oral neoplasia showed differences in intensity, lineshape, and location of peak intensity. We have demonstrated that this system can provide both high quality widefield images, and spectral information at chosen locations within the field of view. Keywords: Hyperspectral, Cancer Detection, Fluorescence, Oral Cancer, Imaging 1. INTRODUCTION Cancer of the oral cavity accounts for over 30,000 deaths per year in the United States 1 and 127,000 deaths per year worldwide 2. Early detection has the potential to greatly improve survival rates. The most common screening method is visual inspection with white light followed by biopsy and histopathology. This approach is limited by the inspector s experience and by low visual contrast between normal and abnormal tissue, and may be confounded by factors such as benign lesions and inflammation. Surgical resection involves removing diseased tissue while minimizing removal of surrounding normal tissue. Frozen section pathology diagnosis may be used during surgery to help ensure removal of all diseased tissue. This method has the drawback of requiring expensive pathology facilities and it does not provide realtime tissue diagnostics in the operating room. Optical techniques are being explored as an aid to both current screening techniques and surgical resection. Widefield multispectral imaging techniques (several centimeters field-of-view) have been shown to be useful for detecting oral malignancies and pre-malignancies by increasing the optical contrast between normal and abnormal regions 3,4. Most notably, a decrease in autofluorescence signal under blue light excitation has been observed in dysplastic and cancerous regions. Red fluorescence, attributed to porphyrins, has also been commonly observed on neoplastic lesions 3,5-7. Additionally, imaging modalities such as narrowband reflectance imaging and orthogonal polarization imaging are being explored for their ability to reveal vascular density in abnormal regions 3,8. Optical point spectroscopy is also being explored as a diagnostic aid in the oral cavity 9-12. Under fluorescence mode, the intensity and spectral shape may be used to help distinguish abnormal tissue from surrounding normal tissue. Gillenwater et al. found a sensitivity of 88% and specificity of 100% for detecting lesions compared to normal tissue using fluorescence spectroscopy 13. Other groups have found comparable results using fluorescence spectroscopy alone or in combination with reflectance spectroscopy 12,14,15. Several groups have developed spectroscopy probes to target specific tissue depths where early biochemical changes occur during malignant progression 11,16,17. A limitation of Advanced Biomedical and Clinical Diagnostic Systems VII, edited by Anita Mahadevan-Jansen, Tuan Vo-Dinh, Warren S. Grundfest, Proc. of SPIE Vol. 7169, 71690J 2009 SPIE CCC code: 1605-7422/09/$18 doi: 10.1117/12.807226 Proc. of SPIE Vol. 7169 71690J-1

optical spectroscopy is that only a small area of tissue is interrogated at one time making it impractical for screening of the entire oral mucosa 5. A combination of widefield imaging, which allows for inspection of the entire oral mucosa, with point spectroscopy methods, which provides more complete spectral information, may provide a powerful tool for in vivo optical detection of oral neoplasia. Hyperspectral imaging offers a method to combine these techniques. This method allows spectra to be obtained from arbitrary locations after images are aquired. Several hyperspectral technologies exist including acoutooptical tunable filters (AOTF), computer tomographic imaging spectrometer (CTIS), and liquid crystal tunable filters (LCTF). Each provides the ability to divide collected light from an image into discrete spectral components which can be used to create emission spectra. Hyperspectral systems have been designed for a variety of bio-applications including measuring tissue perfusion in humans 18, retinal imaging 19, hemoglobin saturation in mouse models 20, and cancer detection. Martin et al. developed a fluorescence endoscopic based LCTF hyperspectral imaging system with laser excitation at 410 nm used for mouse skin 21. They observed decreased fluorescence intensity in malignant regions as well as a peak shift to the red in nude mice with induced tracheal carcinoma. A peak at 600 nm was also observed in spectra obtained from normal and malignant tissue. Gebhart et al. used a LCTF based hyperspectral imaging system for reflectance and fluorescence measurements for brain tumor detection 22. Fluorescence measurements at 340 nm excitation were taken of cortex and glioma. A decreased signal from the glioma tissue was observed. We present a fluorescence hyperspectral imaging system which uses LCTF technology for spectral imaging in the oral cavity using visible excitation and emission wavelengths. We characterize the system by using fluorescence standards to determine the system s spectral resolution and accuracy. We demonstrate, through in vivo measurements of two normal volunteers and two patients with oral mucosal lesions, that the system can provide both high quality widefield images, and spectral information at chosen locations within the field of view. 2.1 Instrumentation 2. METHODS The hyperspectral imaging system was constructed by integrating a LCTF with a multispectral digital microscope (MDM), which is described elsewhere 3. The system is designed around a Zeiss (Thornwood, NY) surgical microscope fitted with 100 watt mercury lamp illumination, excitation and emission filter wheels, scientific grade color CCD cameras (Retiga Exi, QImaging, Burnaby, BC Canada ) and computer control (see figure 1). The system is capable of fluorescence excitation in the UV and near-uv range. The system collects images from a field-of-view (FOV) between 2 and 7 cm. Collected light is split into two optical paths. One path is routed through a collection filter wheel and onto a CCD camera. The second path is routed through the LCTF and to a second camera. This setup allows collection of high spatial, low spectral resolution images through the first path while simultaneously collecting hyperspectral image sets at lower spatial resolution in the second path. Hyperspectral images are collected at a lower spatial resolution because the CCD sensor operates in an 8 x 8 pixel binning mode to decrease exposure times. Proc. of SPIE Vol. 7169 71690J-2

CCD Cameras Mercury Arc Lamp In Illumination Filter Wheel L LCTF - C DC TJ I Collection Filter Wheel Li / Light Guide Microscope Objective Tissue Fig. 1. Optical diagram of the hyperspectral system. The LCTF (CRI, Woburn, MA) is a computer controlled optical filter with a 20 nm full width half maximum (FWHM). The center wavelength of the LCTF passband can be varied through the visible spectrum (400 nm to 720 nm). A hyperspectral image sequence is obtained by setting the LCTF to a desired center wavelength, waiting approximately 150 ms (typical tuning response time), capturing a digital image, and repeating these steps for all desired wavelengths. During post processing a single pixel or region of interest can be chosen from the obtained image set and an emission spectrum plot can be made using the pixel intensity from that region. We have used measurements taken in 2 nm increments for system characterization and measurements taken in 5 nm increments for in vivo tissue measurements. Typical collection times for a complete data block ranged between 30 seconds and 2 minutes depending on the exposure time used for the camera. The accuracy of the resulting spectrum is affected by the tuning accuracy of the LCTF, which is defined as ±3 nm for the LCTF, and by the spectral resolution of the LCTF, which is limited by the intrinsic properties of the device. The spectra experience a low pass smoothing effect equivalent to a convolution of the gaussian shaped pass band of the LCTF with the raw spectral data. The LCTF transmits light of a single linear polarization reducing total transmission of unpolarized light by 50%. Additionally, the LCTF has a transmission ranging between approximately 2% at 400 nm to 53% at 700 nm of the transmitted polarized light. The low transmission required binning of the CCD camera to increase sensitivity and decrease exposure times. Resulting hyperspectral images were captured with 174 x 130 spatial resolution with 8 x 8 pixel binning. Digital images were also collected at full spatial resolution (1392 x 1040 pixels) from the opposing collection path. The spectra obtained using the hyperspectral imaging path could then be spatially correlated with the higher resolution image. 2.2 System Characterization and Calibration The raw spectra produced by the system were corrected to account for system transmission, spectral sensitivity of the CCD detector, and the transmission characteristics of the bayer color filter on the CCD chip. To accomplish this, a calibration lamp with a known emission spectrum (LS-1-CAL, Ocean Optics, Dunedin, Fl) was measured using the system. A hyperspectral image cube was obtained by collecting images at 2 nm intervals from 400 nm to 700 nm. The Proc. of SPIE Vol. 7169 71690J-3

exposure time and gain were kept constant for all images. The measured spectrum was obtained by averaging the pixel values over the portion of FOV showing the calibration lamp for each image. The known spectrum was divided by the measured spectrum to generate a correction factor. This correction factor was used for all subsequent measurements. All spectra shown were smoothed using a moving average filter. Several standards were measured in order to confirm the system correction and spectral accuracy of the system. A mercury lamp emission spectrum was obtained in reflectance mode by placing a 99 % spectralon reflectance standard at the focal plane of the system and illuminating with unfiltered light. Measured emission peak wavelengths were compared to known peaks. Five quantum dot solutions with different emission peaks were measured by the system in fluorescence mode. The location of emission peaks computed with the hyperspectral system were compared with peaks measured with a laboratory spectrophotometer (Spex Fluorolog). A green diode laser (532 nm) was measured with the system in order to confirm spectral accuracy and the full width half maximum (FWHM) of the passband of the LCTF. The measured FWHM was compared with the manufacturer s specifications. 2.3 In Vivo Human Measurements In vivo measurements of human subjects were performed at the University of Texas M.D. Anderson Cancer Center. The study protocol was approved by Institutional Review Boards at the University of Texas M.D. Anderson Cancer Center and at Rice University. Normal volunteers had no history or suspicion of oral carcinoma and were measured in a darkened exam room. Patients were measured under general anesthesia in an operating room prior to surgical resection. Resected tissue was reviewed using standard histopathology. Clinical margins for the lesions were determined by the surgeon (A.G.) and recorded; all regions of abnormal pathology referred to in this study were located inside the clinically determined margins. Proc. of SPIE Vol. 7169 71690J-4

450 500 550 500 Waderth (nm) known spetrum measured spectrum correction factor 650 700 450 500 550 600 650 700 Wa,erngtb (nm) 532 550 564 602 634-7VY\ / 4 Il 7 / /7\ A / / / / / I / / \ / / / / \//\ - / / / / \ / N / / I /\ "/ N //\ / \ \ / / / / \ \ \ N N- ---- 450 500 550 655 650 755 510 515 520 25 530 535 540 545 550 555 592 Wa'elength (nfl) Waunlength (nm) Fig. 2. A. The known spectrum of the tungsten calibration lamp, the spectrum of the calibration lamp measured using the hyperspectral system, and the correction factor obtained by dividing these measurements. B. Mercury arc lamp emission spectrum measured by the hyperspectral system. C. The normalized emission spectra of several quantum dot solutions excited at 365 nm. D. The spectrum of a green diode laser centered at 532 nm. The full width half max of the measurement is 19.1 nm. 3.1 Standards Measurements 3. RESULTS Figure 2A shows the known spectrum of the calibration lamp, the spectrum measured with the hyperspectral system, and the correction factor used for all subsequent measurements. Note that the measured signal in the blue part of the spectrum is relatively low and the correction factor is large. This results from a combination of the lower quantum efficiency of the CCD detector in the blue, the relatively low transmission of the blue portion of the Bayer color mask, and the low transmission of the LCTF in the blue compared to longer wavelengths. Figure 2B shows the spectrum of the mercury lamp used for illumination in the hyperspectral system. Maxima were detected in the spectrum at 550 nm and 578 nm. These correlate with the known mercury lines at 546 nm and 578 nm. Note that the spectrum shown appears smoothed compared to a mercury lamp spectrum obtained using a high spectralresolution instrument. Figure 2C shows the emission spectra of five quantum dot solutions excited at 365 nm. Maxima were detected at 532 nm, 550 nm, 564 nm, 602 nm, and 634 nm. These values had an average deviation of 11 nm compared to the maxima found using a laboratory spectrophotometer (Spex Fluorolog). Figure 2D shows the measured spectra from a 532 nm green laser. The detected maximum was located at 532 nm and the measured FWHM was 19.1 nm. This corresponds closely with the 20 nm FWHM expected from the manufacturer s specifications. Proc. of SPIE Vol. 7169 71690J-5

3.2 In Vivo Human Measurements Spectra of hard palate, buccal mucosa, lip, and tongue were obtained from two normal volunteers. Figure 3 shows representative examples of measurements from indicated locations. The tissue was excited at 405 nm. This example shows variation in spectral intensity and shape at different anatomic locations. Different measurements at the same anatomical site were similar in intensity and line shape. 450 500 550 600 Wavelength (nm) 650 700 Fig. 3. The spectra of several locations measured in a normal volunteer at 405 nm excitation. Measurements from two patients were taken in the operating room before surgical resection of previously confirmed oral lesions. Figure 4A shows a white light image of the left lower lip of the first patient. The pathologic diagnosis for this area is superficially invasive squamous cell carcinoma with areas of carcinoma in situ. The dashed line indicates the clinically determined margin of the abnormal lesion. Figure 4B shows the same field of view in fluorescence mode at 365 nm excitation. The region labeled area a. shows a region inside the clinical margin with a decrease in autofluorescence signal compared to surrounding tissue. Area b. shows an area outside the clinical margin with a comparatively higher level of autofluorescence. Figure 4C. shows the line spectra obtained from these two regions. Each spectrum was obtained by collecting the average pixel value within a circular region with a 8 pixel diameter which corresponds to a circular tissue area of diameter of approximately 2 mm. The maximum intensity of the spectrum from area b. is 2 times the maximum for area a. Both spectra show a broad blue-green signal. Area b. has a maximum located at 475 nm and area a. has a maximum located at 460 nm. Proc. of SPIE Vol. 7169 71690J-6

C. a. 400 450 500 550 600 650 700 WaeIength (nn,) Fig. 4. Figure 4A shows a white light image of the left upper lip from a patient with carcinoma. The dashed oval indicates the clinical margin of the lesion. Figure 4B shows the same field of view under 365 nm excitation. Figure 4C shows the line spectra from regions a. and b. Region a. is located on an area that appears dark inside the clinically abnormal area. Region b. is located outside the clinically abnormal margin. Figure 5A shows a white light image of the left dorsal tongue from the second patient with histopathologically determined multifocal moderately differentiated squamous cell carcinoma. The area also contains white leukoplakia. The dashed line indicates the clinically determined margin of carcinoma. Figure 5B shows a fluorescence image of the same FOV under 405 nm excitation. Red fluorescence is clearly visible in a portion of the image. The region labeled a. includes this red fluorescence and is outside the clinical margin of carcinoma. Region b. is an area with a normal clinical impression without red fluorescence. Region c. is an area of carcinoma with a decreased autofluorescence. The spectra obtained from regions a., b., and c. are shown in Figure 5C. A distinct peak located at 650 nm is present in the spectrum obtained from region a. There is also a significant red shift in the local maximum in the green region from the spectrum from region b. at 495 nm, to region a. at 520 nm. Proc. of SPIE Vol. 7169 71690J-7

450 500 550 600 650 7W Wa..elength (nm) Fig. 5. Figure 5A shows a white light image of the left dorsal tongue of a patient with luekoplakia and carcinoma. The dashed line indicates the clinical margin of carcinoma. Figure 5B shows the same FOV under 405 nm excitation. Figure 5C shows the line spectra from sites a., b. and c. 4. Discussion We have demonstrated the implementation of a liquid-crystal based in vivo hyperspectral imaging system for hyperspectral interrogation of oral neoplasia. The system is capable of capturing multispectral images and line spectra simultaneously. The system was calibrated and characterized to correct for system transmission and to determine spectral accuracy. Measurements of a mercury lamp and green laser emission revealed the smoothing effects on spectra resulting from the bandwidth of the LCTF. The measurement of the mercury lamp emission demonstrated some spectral inaccuracy showing one mercury lamp emission line 4 nm from the known peak. This is not surprising considering that the manufacturer specifies a tuning accuracy of ±3 nm and that additional smoothing in post processing was applied. The measurement of five different quantum dot fluorescence emission also revealed some minor spectral inaccuracies. Differences in excitation and collection geometry between the hyperspectral system and the laboratory spectrophotometer may also contribute to the discrepancies. Imaging measurements of the first patient comparing normal and abnormal tissue regions with 365 nm excitation fluorescence revealed areas of decreased autofluorescence associated with neoplasia. Spectra showed both differences in intensity and subtle differences in spectral shape including shifts in the peak wavelength. Imaging of the second patient at 405 nm showed decreased intensity in the blue-green fluorescence in some areas inside the clinically determined carcinoma region and red fluorescence in other areas. Red fluorescence was also observed outside the clinically determined carcinoma region. The line spectra from the areas with red fluorescence revealed high intensity peaks between 645 and 650 nm consistent with porphyrin emission. Line shape differences and the location of maxima were Proc. of SPIE Vol. 7169 71690J-8

also different between the clinically normal site with no red fluorescence, the site with red fluorescence, and the site with pathologically confirmed carcinoma. Decreased autofluorescence and a decrease in spectral intensity in the blue-green portion of the visible spectrum in neoplastic areas has been previously reported in the oral cavity 3,4,13,23. This decrease in intensity is likely due to increased density of microvasculature which contains light absorbing hemoglobin and a decrease in fluorescent collagen crosslinks in the stroma region of the epithelium 24,25. The red fluorescence observed at 405 nm excitation has been well documented by several groups both by imaging and spectroscopy 6,10,13,14,26. Most groups using point spectroscopy methods observe the major peak in the red at approximately 635 nm 13,14. Our measurements show a peak at somewhat longer wavelength around 645 nm to 650 nm. Several factors could contribute to this discrepancy including differences in biological factors, collection geometry of the hyperspectral system compared to point spectroscopy, and spectral accuracy of the system. Gebhart et al. present an analysis on the effects of collection geometry and variability between point spectroscopic systems and hyperspectral imaging systems 22. In vivo hyperspectral imaging, in principle, allows for screening of relatively large areas of tissue using traditional imaging methods with the added ability to more closely interrogate suspicious areas by creating a line spectrum of emitted light from these areas. This method has the potential of increasing sensitivity and specificity of screening and margin detection during surgery above that available with either method alone. The utility of hyperspectral systems for cancer detection and disease classification depends on their ability to obtain spectral data with density and accuracy similar to those available with current spectroscopy systems. A limitation of the hyperspectral imaging system described in this article is the spectral resolution and accuracy. These limitations are due to the broad (~20 nm) FWHM of the passband of the LCTF and the large correction factor needed to correct measured raw spectra. This limits the ability to precisely measure the location of spectra maximums, often an important classifier for traditional point-measurement spectroscopy systems. LCTFs are available with smaller FWHM passbands but transmission is further reduced, which in turn requires either longer exposure times or more sensitive and expensive detection equipment. Although some differences in intensity, line shape, and maxima location were observed when comparing normal and abnormal tissue regions, it is unclear whether this type of hyperspectral imaging will supply more information than that available with autofluorecence imaging alone. Studies with more patients are needed to determine whether this hyperspectral system is able to increase the specificity of disease detection beyond that of autofluorescence imaging alone. REFERENCES [1] [2] [3] [4] [5] [6] [7] [8] [9] [10] [11] American Cancer Society, "Cancer Facts and Figures: 2005". (2005). Parkin, D.M., et al., "Global cancer statistics, 2002". CA Cancer J Clin. 55(2): p. 74-108, (2005). Roblyer, D., et al., "Multispectral optical imaging device for in vivo detection of oral neoplasia". J Biomed Opt. 13(2): p. 024019, (2008). Lane, P.M., et al., "Simple device for the direct visualization of oral-cavity tissue fluorescence". J Biomed Opt. 11(2): p. 024006, (2006). De Veld, D.C., et al., "The status of in vivo autofluorescence spectroscopy and imaging for oral oncology". Oral Oncol. 41(2): p. 117-31, (2005). Onizawa, K., et al., "Usefulness of fluorescence photography for diagnosis of oral cancer". Int J Oral Maxillofac Surg. 28(3): p. 206-10, (1999). Ingrams, D.R., et al., "Autofluorescence characteristics of oral mucosa". Head Neck. 19(1): p. 27-32, (1997). Subhash, N., et al., "Oral cancer detection using diffuse reflectance spectral ratio R540/R575 of oxygenated hemoglobin bands". J Biomed Opt. 11(1): p. 014018, (2006). de Veld, D.C., et al., "Clinical study for classification of benign, dysplastic, and malignant oral lesions using autofluorescence spectroscopy". J Biomed Opt. 9(5): p. 940-50, (2004). Betz, C.S., et al., "Autofluorescence imaging and spectroscopy of normal and malignant mucosa in patients with head and neck cancer". Lasers Surg Med. 25(4): p. 323-34, (1999). Schwarz, R.A., et al., "Autofluorescence and diffuse reflectance spectroscopy of oral epithelial tissue using a depth-sensitive fiber-optic probe". Appl Opt. 47(6): p. 825-34, (2008). Proc. of SPIE Vol. 7169 71690J-9

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