Optical diagnosis and characterization of dental caries with polarization-resolved hyperspectral stimulated Raman scattering microscopy Zi Wang,1 Wei Zheng,1,2 Stephen Chin-Ying Hsu,3 and Zhiwei Huang1,* 1
Optical Bioimaging Laboratory, Department of Biomedical Engineering, Faculty of Engineering, National University of Singapore, 117576 Singapore Department of Medicine, Yong Loo Lin School of Medicine, National University of Singapore and National University Health System, 119260 Singapore 3 Department of Dentistry, Faculty of Dentistry, National University Health System and National University of Singapore, 119083 Singapore * [email protected] 2
Abstract: We report the utility of a rapid polarization-resolved hyperspectral stimulated Raman scattering (SRS) imaging technique developed for optical diagnosis and characterization of dental caries in the tooth. Hyperspectral SRS images (512 × 512 pixels) of the tooth covering both the fingerprint (800-1800 cm−1) and high-wavenumber (2800-3600 cm−1) regions can be acquired within 15 minutes, which is at least 103 faster in imaging speed than confocal Raman mapping. Hyperspectral SRS imaging uncovers the biochemical distributions and variations across the carious enamel in the tooth. SRS imaging shows that compared to the sound enamel, the mineral content in the body of lesion decreases by 55%; while increasing up to 110% in the surface zone, indicating the formation of a hyper-mineralized layer due to the remineralization process. Further polarized SRS imaging shows that the depolarization ratios of hydroxyapatite crystals (ν1-PO43- of SRS at 959 cm−1) of the tooth in the sound enamel, translucent zone, body of lesion and the surface zone are 0.035 ± 0.01, 0.052 ± 0.02, 0.314 ± 0.1, 0.038 ± 0.02, respectively, providing a new diagnostic criterion for discriminating carious lesions from sound enamel in the teeth. This work demonstrates for the first time that the polarization-resolved hyperspectral SRS imaging technique can be used for quantitatively determining tooth mineralization levels and discriminating carious lesions from sound enamel in a rapid fashion, proving its promising potential of early detection and diagnosis of dental caries without labeling. ©2016 Optical Society of America OCIS codes: (180.4315) Nonlinear microscopy; (180.5655) Raman microscopy; (190.4180) Multiphoton processes; (170.1850) Dentistry.
References and links 1. 2. 3. 4. 5. 6.
R. H. Selwitz, A. I. Ismail, and N. B. Pitts, “Dental caries,” Lancet 369(9555), 51–59 (2007). J. D. Featherstone, “Prevention and reversal of dental caries: role of low level fluoride,” Community Dent. Oral Epidemiol. 27(1), 31–40 (1999). K. Ekstrand, V. Qvist, and A. Thylstrup, “Light microscope study of the effect of probing in occlusal surfaces,” Caries Res. 21(4), 368–374 (1987). J. Vaarkamp, J. J. Ten Bosch, E. H. Verdonschot, and S. Tranaeus, “Quantitative diagnosis of small approximal caries lesions utilizing wavelength-dependent fiber-optic transillumination,” J. Dent. Res. 76(4), 875–882 (1997). B. T. Amaechi, A. Podoleanu, S. M. Higham, and D. A. Jackson, “Correlation of quantitative light-induced fluorescence and optical coherence tomography applied for detection and quantification of early dental caries,” J. Biomed. Opt. 8(4), 642–647 (2003). A. C. Ko, L. P. Choo-Smith, M. Hewko, L. Leonardi, M. G. Sowa, C. C. S. Dong, P. Williams, and B. Cleghorn, “Ex vivo detection and characterization of early dental caries by optical coherence tomography and Raman spectroscopy,” J. Biomed. Opt. 10(3), 031118 (2005).
#255753 © 2016 OSA
Received 17 Dec 2015; revised 1 Mar 2016; accepted 8 Mar 2016; published 15 Mar 2016 1 Apr 2016 | Vol. 7, No. 4 | DOI:10.1364/BOE.7.001284 | BIOMEDICAL OPTICS EXPRESS 1284
7. 8. 9. 10. 11. 12. 13. 14. 15. 16. 17. 18. 19. 20. 21. 22. 23. 24. 25. 26. 27. 28. 29.
A. C. Ko, L. P. i. Choo-Smith, M. Hewko, M. G. Sowa, C. C. S. Dong, and B. Cleghorn, “Detection of early dental caries using polarized Raman spectroscopy,” Opt. Express 14(1), 203–215 (2006). H. Tsuda and J. Arends, “Orientational micro-Raman spectroscopy on hydroxyapatite single crystals and human enamel crystallites,” J. Dent. Res. 73(11), 1703–1710 (1994). H. Salehi, E. Terrer, I. Panayotov, B. Levallois, B. Jacquot, H. Tassery, and F. Cuisinier, “Functional mapping of human sound and carious enamel and dentin with Raman spectroscopy,” J. Biophotonics 6(10), 765–774 (2013). A. Zumbusch, G. R. Holtom, and X. S. Xie, “Three-dimensional vibrational imaging by coherent anti-Stokes Raman scattering,” Phys. Rev. Lett. 82(20), 4142–4145 (1999). C. W. Freudiger, W. Min, B. G. Saar, S. Lu, G. R. Holtom, C. He, J. C. Tsai, J. X. Kang, and X. S. Xie, “Labelfree biomedical imaging with high sensitivity by stimulated Raman scattering microscopy,” Science 322(5909), 1857–1861 (2008). F. Lu, W. Zheng, C. Sheppard, and Z. Huang, “Interferometric polarization coherent anti-Stokes Raman scattering (IP-CARS) microscopy,” Opt. Lett. 33(6), 602–604 (2008). F. Lu, W. Zheng, J. Lin, and Z. Huang, “Integrated coherent anti-Stokes Raman scattering and multiphoton microscopy for biomolecular imaging using spectral filtering of a femtosecond laser,” Appl. Phys. Lett. 96(13), 133701 (2010). P. K. Upputuri, J. Lin, L. Gong, X. Y. Liu, H. Wang, and Z. Huang, “Circularly polarized coherent anti-Stokes Raman scattering microscopy,” Opt. Lett. 38(8), 1262–1264 (2013). C. L. Evans, E. O. Potma, M. Puoris’haag, D. Côté, C. P. Lin, and X. S. Xie, “Chemical imaging of tissue in vivo with video-rate coherent anti-Stokes Raman scattering microscopy,” Proc. Natl. Acad. Sci. U.S.A. 102(46), 16807–16812 (2005). B. G. Saar, C. W. Freudiger, J. Reichman, C. M. Stanley, G. R. Holtom, and X. S. Xie, “Video-Rate Molecular Imaging in Vivo with Stimulated Raman Scattering,” Science 330(6009), 1368–1370 (2010). Y. Liu, C. Y. S. Hsu, C. M. J. Teo, and S. H. Teoh, “Potential Mechanism for the Laser-Fluoride Effect on Enamel Demineralization,” J. Dent. Res. 92(1), 71–75 (2013). Z. Wang, W. Zheng, C. Y. S. Hsu, and Z. Huang, “Polarization-resolved hyperspectral stimulated Raman scattering microscopy for label-free biomolecular imaging of the tooth,” Appl. Phys. Lett. 108(3), 033701 (2016). H. Tsuda and J. Arends, “Raman spectroscopy in dental research: a short review of recent studies,” Adv. Dent. Res. 11(4), 539–547 (1997). J. Lin, W. Zheng, Z. Wang, and Z. Huang, “Label-free three-dimensional imaging of cell nucleus using thirdharmonic generation microscopy,” Appl. Phys. Lett. 105(10), 103705 (2014). J. Lin, S. Teh, W. Zheng, Z. Wang, and Z. Huang, “Multimodal nonlinear optical microscopic imaging provides new insights into acetowhitening mechanisms in live mammalian cells without labeling,” Biomed. Opt. Express 5(9), 3116–3122 (2014). Z. Wang, W. Zheng, C. Y. S. Hsu, and Z. Huang, “Epi-detected quadruple-modal nonlinear optical microscopy for label-free imaging of the tooth,” Appl. Phys. Lett. 106(3), 033701 (2015). J. Hicks, F. Garcia-Godoy, and C. Flaitz, “Biological factors in dental caries enamel structure and the caries process in the dynamic process of demineralization and remineralization (part 2),” J. Clin. Pediatr. Dent. 28(2), 119–124 (2005). C. Robinson, R. C. Shore, S. J. Brookes, S. Strafford, S. R. Wood, and J. Kirkham, “The chemistry of enamel caries,” Crit. Rev. Oral Biol. Med. 11(4), 481–495 (2000). J. A. Weatherell, C. Robinson, and A. S. Hallsworth, “Variations in the chemical composition of human enamel,” J. Dent. Res. 53(2), 180–192 (1974). B. Richards and E. Wolf, “Electromagnetic diffraction in optical systems. II. Structure of the image field in an aplanatic system,” Proc. R. Soc. Lond. A Math. Phys. Sci. 253(1274), 358–379 (1959). J. Lin, W. Zheng, H. Wang, and Z. Huang, “Effects of scatterers’ sizes on near-field coherent anti-Stokes Raman scattering under tightly focused radially and linearly polarized light excitation,” Opt. Express 18(10), 10888– 10895 (2010). F. Lu, W. Zheng, and Z. Huang, “Coherent anti-Stokes Raman scattering microscopy using tightly focused radially polarized light,” Opt. Lett. 34(12), 1870–1872 (2009). S. Y. Chen, C. Y. S. Hsu, and C. K. Sun, “Epi-third and second harmonic generation microscopic imaging of abnormal enamel,” Opt. Express 16(15), 11670–11679 (2008).
1. Introduction Dental caries is caused by an imbalance between the demineralization and re-mineralization processes of the tooth related to the acid generated from the bacterial activity [1]. Early detection and diagnosis of early dental caries in the tooth is essential to timely intervene the carious process through preventative treatments [2]. The visual and tactile inspection method that is commonly used to detect dental caries lacks sufficient sensitivity for detecting early onset of caries. Moreover, the use of sharp dental explorer during the visual and tactile examination may result in irreversible traumatic defects on incipient lesion [3]. Other
#255753 © 2016 OSA
Received 17 Dec 2015; revised 1 Mar 2016; accepted 8 Mar 2016; published 15 Mar 2016 1 Apr 2016 | Vol. 7, No. 4 | DOI:10.1364/BOE.7.001284 | BIOMEDICAL OPTICS EXPRESS 1285
diagnostic techniques, such as fiber optic trans-illumination [4], quantitative light-induced fluorescence [5], have problem with high false positive or false negative rates for detecting early stage of dental caries. With a high biochemical/biomolecular specificity and biomolecular orientation sensitivity, polarized Raman spectroscopy and microscopy has emerged as an appealing tool for biochemical characterization and diagnosis in the tooth [6– 8]. However, due to the extremely weak tissue Raman scattering process, the measurements of tissue micro-Raman images may take tens of hours, hampering its wide applications in rapid diagnosis and characterization of biomedical tissues. The concomitant strong fluorescence background from tooth samples may also overwhelm weak tissue Raman signals [9], making micro-Raman imaging on dental caries even more challenging. To address the above problems encountered in confocal Raman imaging, coherent Raman scattering (CRS) (e.g., coherent anti-Stokes Raman scattering (CARS) and stimulated Raman scattering (SRS)) has been recently developed for biomedical imaging [10–14]. CRS enhances the weak Raman signal by 6 orders of magnitude through resonant enhancement nonlinear processes, enabling CRS imaging speed up to video rate [15, 16]. Unlike CARS, SRS is free from nonresonant background interference and has a linear dependence on the biochemical concentration, making SRS to be an attractive tool for quantitative imaging of biochemical compositions and distributions in unstained live cells and tissue [11]. In this work, we report the utility of a rapid polarization-resolved hyperspectral SRS imaging technique developed for optical diagnosis and characterization of carious lesions in the tooth without labeling. We demonstrate that the hyperspectral SRS imaging can acquire Raman images of the tooth covering both the fingerprint (FP) (800-1800 cm−1) and high-wavenumber (HW) (2800-3600 cm−1) regions within 15 mins, offering 103 faster in imaging speed than conventional confocal Raman imaging. The complementary information acquired in both the FP and HW Raman regions can further improve the understanding of biochemical/biomolecular distributions and orientations associated with carious lesions in the tooth. 2. Material and methods 2.1 Tooth samples preparation This research protocol was approved by the Institutional Review Board (IRB) of National University of Singapore. Without identifiers, the teeth samples donated from various individuals/clinics were washed in distilled water and stored in saline. The teeth were then cleaned using a soft toothbrush and transferred to a 0.1% thymol storage medium before being sectioned longitudinally from cusp tip to cemento enamel junction (CEJ). Sections were prepared in the labial/buccal-lingual/palatal direction and centred through the unworn cuspal tips and the underlying dentine horns, using a Buehler IsoMet 1000 with a cutting diamondwafering blade. With approximately 150-180 μm in thickness, the sections were hand ground using a graded series of Buehler Met-II grinding pads (P800, P1000, P1200, P2500, P4000) with silicon carbide abrasive on a Buehler Phoenix Beta Grinding/Polishing Machine, until a thickness of 80-100 μm was attained and confirmed with vernier calipers. Sections were then washed using distilled water, and air dried for 24 hours to remove smear layer and contaminants from the surface, before preliminary characterization using Olympus BX51 polarized light microscope with a digital microscope camera (Olympus DP25) and imaging software (Olympus Cell D) before polarization-resolved hyperspectral SRS imaging.. 2.2 Polarization-resolved hyperspectral SRS imaging system Figure 1 shows the schematic of the polarization-resolved hyperspectral SRS imaging system developed for label-free molecular imaging [18]. The excitation laser sources consist of an Nd:YVO4 laser (High-Q Laser, Austria) and an optical parametric oscillator (OPO) (Levante Emerald, APE-Berlin). A portion (20%) of the fundamental output of Nd:YVO4 laser (80 MHz, 7.5 ps pulses at 1064 nm) serves as the Stokes beam for SRS imaging. The remaining
#255753 © 2016 OSA
Received 17 Dec 2015; revised 1 Mar 2016; accepted 8 Mar 2016; published 15 Mar 2016 1 Apr 2016 | Vol. 7, No. 4 | DOI:10.1364/BOE.7.001284 | BIOMEDICAL OPTICS EXPRESS 1286
output (80%) is frequency doubled (532 nm) to pump the OPO, whereby the tunable signal generation ranging from 670 to 980 nm is used as the pump beam for SRS imaging. The wide tunable range of OPO enables the SRS imaging in both fingerprint (800–1800 cm−1) and highwavenumber (2800–3600 cm−1) spectral regions. The 1064 nm Stokes beam is modulated at 20 MHz by an electro-optic modulator (EOM). The spatially and temporally overlapped pump and Stokes beams are sent to a multiphoton scanning microscope and focused onto the sample through a water-immersion objective (XLUMPLFLN 20 × , NA = 1.0, Olympus Inc.). The transmitted pump beam is collected in the forward direction with a condenser (NA = 1.4, Nikon Inc.) and spectrally isolated from the Stokes beam with bandpass filter sets (a filter set centered at 780 nm with a 160 nm bandwidth for SRS imaging in HW region; another filter set centered at 900 nm with a 100 nm bandwidth is used for SRS imaging in FP region) for hyperspectral SRS imaging. The modulation of the pump beam intensity due to the stimulated Raman loss (SRL) process is detected by a photodiode (FDS1010, Thorlabs Inc.). A lock-in amplifier with the time constant as short as 100 ns is used to demodulate the pump beam to acquire the SRS signal. SRS images (e.g., 959 cm−1 (ν1-PO43- of hydroxyapatite (HA) crystals of the tooth) [19]) of 512 × 512 pixels can be acquired within 0.3 s/frame with a 1μs pixel dwell time. The SRS imaging speed can be further boosted to video rate by using resonant scanners. One notes that other nonlinear microscopy imaging modalities (e.g., secondharmonic generation (SHG), third-harmonic generation (THG), two-photon excitation fluorescence (TPEF), coherent anti-Stokes Raman scattering (CARS)) can also be incorporated into the hyperspectral SRS imaging system (Fig. 1) to better understand the morphological architectures, biochemical structures and compositions of tissue in a comprehensive fashion [20, 21]. The choices of photomultiplier tubes (PMTs), filters and dichroic mirrors for epi-detected SHG/THG/TPEF/CARS imaging modalities have been reported in our previous work [22]. For comparison purpose, a confocal Raman microscope (Renishaw inVia, UK) equipped with a 20 × objective (N.A. = 1.0) and a 785 nm laser with an excitation power of 50 mW is utilized for micro-Raman spectroscopy and imaging of the tooth. The spectral resolution of the confocal Raman microscope is ~4 cm−1.The pixel dwell time is typically >5 s to ensure the good SNRs of Raman spectra for tooth Raman imaging. Hyperspectral SRS imaging is accomplished by scanning the pump beam wavelengths of OPO through computer control of the crystal temperature, tilt angles of Lyot filter and the cavity length of OPO, generating a three-dimensional data stack (x, y, Ω), where Ω = ωp-ωs is Raman shift. The tuning of pump wavelengths is synchronized to the frame trigger of microscope to achieve automatically spectral scanning of SRS images at different Raman shifts. The hyperspectral SRS stack is then normalized by the intensity of the pump beam at each wavelength recorded by a photodiode which detects a small fraction of the OPO output after a beam splitter. Automated wavelength tuning, SRS images acquisition, intensity normalization, and three-dimensional hyperspectral SRS data generation are controlled by using a house-built software with LabVIEW programming. For polarization-resolved SRS imaging, the polarizations of the pump and Stokes beams are rotated independently by using achromatic half-wave plates mounted in step motors. An analyzer is placed before the photodiode with the analyzer polarization parallel to the pump beam polarization. Polarization-resolved SRS images as a function of the polarization angle (θ) between the polarization directions of the pump and Stokes beams are acquired from 0° to 180° with 20° intervals. The total acquisition time of the polarized SRS image (512 × 512 pixels) is typically
Optical Bioimaging Laboratory, Department of Biomedical Engineering, Faculty of Engineering, National University of Singapore, 117576 Singapore Department of Medicine, Yong Loo Lin School of Medicine, National University of Singapore and National University Health System, 119260 Singapore 3 Department of Dentistry, Faculty of Dentistry, National University Health System and National University of Singapore, 119083 Singapore * [email protected] 2
Abstract: We report the utility of a rapid polarization-resolved hyperspectral stimulated Raman scattering (SRS) imaging technique developed for optical diagnosis and characterization of dental caries in the tooth. Hyperspectral SRS images (512 × 512 pixels) of the tooth covering both the fingerprint (800-1800 cm−1) and high-wavenumber (2800-3600 cm−1) regions can be acquired within 15 minutes, which is at least 103 faster in imaging speed than confocal Raman mapping. Hyperspectral SRS imaging uncovers the biochemical distributions and variations across the carious enamel in the tooth. SRS imaging shows that compared to the sound enamel, the mineral content in the body of lesion decreases by 55%; while increasing up to 110% in the surface zone, indicating the formation of a hyper-mineralized layer due to the remineralization process. Further polarized SRS imaging shows that the depolarization ratios of hydroxyapatite crystals (ν1-PO43- of SRS at 959 cm−1) of the tooth in the sound enamel, translucent zone, body of lesion and the surface zone are 0.035 ± 0.01, 0.052 ± 0.02, 0.314 ± 0.1, 0.038 ± 0.02, respectively, providing a new diagnostic criterion for discriminating carious lesions from sound enamel in the teeth. This work demonstrates for the first time that the polarization-resolved hyperspectral SRS imaging technique can be used for quantitatively determining tooth mineralization levels and discriminating carious lesions from sound enamel in a rapid fashion, proving its promising potential of early detection and diagnosis of dental caries without labeling. ©2016 Optical Society of America OCIS codes: (180.4315) Nonlinear microscopy; (180.5655) Raman microscopy; (190.4180) Multiphoton processes; (170.1850) Dentistry.
References and links 1. 2. 3. 4. 5. 6.
R. H. Selwitz, A. I. Ismail, and N. B. Pitts, “Dental caries,” Lancet 369(9555), 51–59 (2007). J. D. Featherstone, “Prevention and reversal of dental caries: role of low level fluoride,” Community Dent. Oral Epidemiol. 27(1), 31–40 (1999). K. Ekstrand, V. Qvist, and A. Thylstrup, “Light microscope study of the effect of probing in occlusal surfaces,” Caries Res. 21(4), 368–374 (1987). J. Vaarkamp, J. J. Ten Bosch, E. H. Verdonschot, and S. Tranaeus, “Quantitative diagnosis of small approximal caries lesions utilizing wavelength-dependent fiber-optic transillumination,” J. Dent. Res. 76(4), 875–882 (1997). B. T. Amaechi, A. Podoleanu, S. M. Higham, and D. A. Jackson, “Correlation of quantitative light-induced fluorescence and optical coherence tomography applied for detection and quantification of early dental caries,” J. Biomed. Opt. 8(4), 642–647 (2003). A. C. Ko, L. P. Choo-Smith, M. Hewko, L. Leonardi, M. G. Sowa, C. C. S. Dong, P. Williams, and B. Cleghorn, “Ex vivo detection and characterization of early dental caries by optical coherence tomography and Raman spectroscopy,” J. Biomed. Opt. 10(3), 031118 (2005).
#255753 © 2016 OSA
Received 17 Dec 2015; revised 1 Mar 2016; accepted 8 Mar 2016; published 15 Mar 2016 1 Apr 2016 | Vol. 7, No. 4 | DOI:10.1364/BOE.7.001284 | BIOMEDICAL OPTICS EXPRESS 1284
7. 8. 9. 10. 11. 12. 13. 14. 15. 16. 17. 18. 19. 20. 21. 22. 23. 24. 25. 26. 27. 28. 29.
A. C. Ko, L. P. i. Choo-Smith, M. Hewko, M. G. Sowa, C. C. S. Dong, and B. Cleghorn, “Detection of early dental caries using polarized Raman spectroscopy,” Opt. Express 14(1), 203–215 (2006). H. Tsuda and J. Arends, “Orientational micro-Raman spectroscopy on hydroxyapatite single crystals and human enamel crystallites,” J. Dent. Res. 73(11), 1703–1710 (1994). H. Salehi, E. Terrer, I. Panayotov, B. Levallois, B. Jacquot, H. Tassery, and F. Cuisinier, “Functional mapping of human sound and carious enamel and dentin with Raman spectroscopy,” J. Biophotonics 6(10), 765–774 (2013). A. Zumbusch, G. R. Holtom, and X. S. Xie, “Three-dimensional vibrational imaging by coherent anti-Stokes Raman scattering,” Phys. Rev. Lett. 82(20), 4142–4145 (1999). C. W. Freudiger, W. Min, B. G. Saar, S. Lu, G. R. Holtom, C. He, J. C. Tsai, J. X. Kang, and X. S. Xie, “Labelfree biomedical imaging with high sensitivity by stimulated Raman scattering microscopy,” Science 322(5909), 1857–1861 (2008). F. Lu, W. Zheng, C. Sheppard, and Z. Huang, “Interferometric polarization coherent anti-Stokes Raman scattering (IP-CARS) microscopy,” Opt. Lett. 33(6), 602–604 (2008). F. Lu, W. Zheng, J. Lin, and Z. Huang, “Integrated coherent anti-Stokes Raman scattering and multiphoton microscopy for biomolecular imaging using spectral filtering of a femtosecond laser,” Appl. Phys. Lett. 96(13), 133701 (2010). P. K. Upputuri, J. Lin, L. Gong, X. Y. Liu, H. Wang, and Z. Huang, “Circularly polarized coherent anti-Stokes Raman scattering microscopy,” Opt. Lett. 38(8), 1262–1264 (2013). C. L. Evans, E. O. Potma, M. Puoris’haag, D. Côté, C. P. Lin, and X. S. Xie, “Chemical imaging of tissue in vivo with video-rate coherent anti-Stokes Raman scattering microscopy,” Proc. Natl. Acad. Sci. U.S.A. 102(46), 16807–16812 (2005). B. G. Saar, C. W. Freudiger, J. Reichman, C. M. Stanley, G. R. Holtom, and X. S. Xie, “Video-Rate Molecular Imaging in Vivo with Stimulated Raman Scattering,” Science 330(6009), 1368–1370 (2010). Y. Liu, C. Y. S. Hsu, C. M. J. Teo, and S. H. Teoh, “Potential Mechanism for the Laser-Fluoride Effect on Enamel Demineralization,” J. Dent. Res. 92(1), 71–75 (2013). Z. Wang, W. Zheng, C. Y. S. Hsu, and Z. Huang, “Polarization-resolved hyperspectral stimulated Raman scattering microscopy for label-free biomolecular imaging of the tooth,” Appl. Phys. Lett. 108(3), 033701 (2016). H. Tsuda and J. Arends, “Raman spectroscopy in dental research: a short review of recent studies,” Adv. Dent. Res. 11(4), 539–547 (1997). J. Lin, W. Zheng, Z. Wang, and Z. Huang, “Label-free three-dimensional imaging of cell nucleus using thirdharmonic generation microscopy,” Appl. Phys. Lett. 105(10), 103705 (2014). J. Lin, S. Teh, W. Zheng, Z. Wang, and Z. Huang, “Multimodal nonlinear optical microscopic imaging provides new insights into acetowhitening mechanisms in live mammalian cells without labeling,” Biomed. Opt. Express 5(9), 3116–3122 (2014). Z. Wang, W. Zheng, C. Y. S. Hsu, and Z. Huang, “Epi-detected quadruple-modal nonlinear optical microscopy for label-free imaging of the tooth,” Appl. Phys. Lett. 106(3), 033701 (2015). J. Hicks, F. Garcia-Godoy, and C. Flaitz, “Biological factors in dental caries enamel structure and the caries process in the dynamic process of demineralization and remineralization (part 2),” J. Clin. Pediatr. Dent. 28(2), 119–124 (2005). C. Robinson, R. C. Shore, S. J. Brookes, S. Strafford, S. R. Wood, and J. Kirkham, “The chemistry of enamel caries,” Crit. Rev. Oral Biol. Med. 11(4), 481–495 (2000). J. A. Weatherell, C. Robinson, and A. S. Hallsworth, “Variations in the chemical composition of human enamel,” J. Dent. Res. 53(2), 180–192 (1974). B. Richards and E. Wolf, “Electromagnetic diffraction in optical systems. II. Structure of the image field in an aplanatic system,” Proc. R. Soc. Lond. A Math. Phys. Sci. 253(1274), 358–379 (1959). J. Lin, W. Zheng, H. Wang, and Z. Huang, “Effects of scatterers’ sizes on near-field coherent anti-Stokes Raman scattering under tightly focused radially and linearly polarized light excitation,” Opt. Express 18(10), 10888– 10895 (2010). F. Lu, W. Zheng, and Z. Huang, “Coherent anti-Stokes Raman scattering microscopy using tightly focused radially polarized light,” Opt. Lett. 34(12), 1870–1872 (2009). S. Y. Chen, C. Y. S. Hsu, and C. K. Sun, “Epi-third and second harmonic generation microscopic imaging of abnormal enamel,” Opt. Express 16(15), 11670–11679 (2008).
1. Introduction Dental caries is caused by an imbalance between the demineralization and re-mineralization processes of the tooth related to the acid generated from the bacterial activity [1]. Early detection and diagnosis of early dental caries in the tooth is essential to timely intervene the carious process through preventative treatments [2]. The visual and tactile inspection method that is commonly used to detect dental caries lacks sufficient sensitivity for detecting early onset of caries. Moreover, the use of sharp dental explorer during the visual and tactile examination may result in irreversible traumatic defects on incipient lesion [3]. Other
#255753 © 2016 OSA
Received 17 Dec 2015; revised 1 Mar 2016; accepted 8 Mar 2016; published 15 Mar 2016 1 Apr 2016 | Vol. 7, No. 4 | DOI:10.1364/BOE.7.001284 | BIOMEDICAL OPTICS EXPRESS 1285
diagnostic techniques, such as fiber optic trans-illumination [4], quantitative light-induced fluorescence [5], have problem with high false positive or false negative rates for detecting early stage of dental caries. With a high biochemical/biomolecular specificity and biomolecular orientation sensitivity, polarized Raman spectroscopy and microscopy has emerged as an appealing tool for biochemical characterization and diagnosis in the tooth [6– 8]. However, due to the extremely weak tissue Raman scattering process, the measurements of tissue micro-Raman images may take tens of hours, hampering its wide applications in rapid diagnosis and characterization of biomedical tissues. The concomitant strong fluorescence background from tooth samples may also overwhelm weak tissue Raman signals [9], making micro-Raman imaging on dental caries even more challenging. To address the above problems encountered in confocal Raman imaging, coherent Raman scattering (CRS) (e.g., coherent anti-Stokes Raman scattering (CARS) and stimulated Raman scattering (SRS)) has been recently developed for biomedical imaging [10–14]. CRS enhances the weak Raman signal by 6 orders of magnitude through resonant enhancement nonlinear processes, enabling CRS imaging speed up to video rate [15, 16]. Unlike CARS, SRS is free from nonresonant background interference and has a linear dependence on the biochemical concentration, making SRS to be an attractive tool for quantitative imaging of biochemical compositions and distributions in unstained live cells and tissue [11]. In this work, we report the utility of a rapid polarization-resolved hyperspectral SRS imaging technique developed for optical diagnosis and characterization of carious lesions in the tooth without labeling. We demonstrate that the hyperspectral SRS imaging can acquire Raman images of the tooth covering both the fingerprint (FP) (800-1800 cm−1) and high-wavenumber (HW) (2800-3600 cm−1) regions within 15 mins, offering 103 faster in imaging speed than conventional confocal Raman imaging. The complementary information acquired in both the FP and HW Raman regions can further improve the understanding of biochemical/biomolecular distributions and orientations associated with carious lesions in the tooth. 2. Material and methods 2.1 Tooth samples preparation This research protocol was approved by the Institutional Review Board (IRB) of National University of Singapore. Without identifiers, the teeth samples donated from various individuals/clinics were washed in distilled water and stored in saline. The teeth were then cleaned using a soft toothbrush and transferred to a 0.1% thymol storage medium before being sectioned longitudinally from cusp tip to cemento enamel junction (CEJ). Sections were prepared in the labial/buccal-lingual/palatal direction and centred through the unworn cuspal tips and the underlying dentine horns, using a Buehler IsoMet 1000 with a cutting diamondwafering blade. With approximately 150-180 μm in thickness, the sections were hand ground using a graded series of Buehler Met-II grinding pads (P800, P1000, P1200, P2500, P4000) with silicon carbide abrasive on a Buehler Phoenix Beta Grinding/Polishing Machine, until a thickness of 80-100 μm was attained and confirmed with vernier calipers. Sections were then washed using distilled water, and air dried for 24 hours to remove smear layer and contaminants from the surface, before preliminary characterization using Olympus BX51 polarized light microscope with a digital microscope camera (Olympus DP25) and imaging software (Olympus Cell D) before polarization-resolved hyperspectral SRS imaging.. 2.2 Polarization-resolved hyperspectral SRS imaging system Figure 1 shows the schematic of the polarization-resolved hyperspectral SRS imaging system developed for label-free molecular imaging [18]. The excitation laser sources consist of an Nd:YVO4 laser (High-Q Laser, Austria) and an optical parametric oscillator (OPO) (Levante Emerald, APE-Berlin). A portion (20%) of the fundamental output of Nd:YVO4 laser (80 MHz, 7.5 ps pulses at 1064 nm) serves as the Stokes beam for SRS imaging. The remaining
#255753 © 2016 OSA
Received 17 Dec 2015; revised 1 Mar 2016; accepted 8 Mar 2016; published 15 Mar 2016 1 Apr 2016 | Vol. 7, No. 4 | DOI:10.1364/BOE.7.001284 | BIOMEDICAL OPTICS EXPRESS 1286
output (80%) is frequency doubled (532 nm) to pump the OPO, whereby the tunable signal generation ranging from 670 to 980 nm is used as the pump beam for SRS imaging. The wide tunable range of OPO enables the SRS imaging in both fingerprint (800–1800 cm−1) and highwavenumber (2800–3600 cm−1) spectral regions. The 1064 nm Stokes beam is modulated at 20 MHz by an electro-optic modulator (EOM). The spatially and temporally overlapped pump and Stokes beams are sent to a multiphoton scanning microscope and focused onto the sample through a water-immersion objective (XLUMPLFLN 20 × , NA = 1.0, Olympus Inc.). The transmitted pump beam is collected in the forward direction with a condenser (NA = 1.4, Nikon Inc.) and spectrally isolated from the Stokes beam with bandpass filter sets (a filter set centered at 780 nm with a 160 nm bandwidth for SRS imaging in HW region; another filter set centered at 900 nm with a 100 nm bandwidth is used for SRS imaging in FP region) for hyperspectral SRS imaging. The modulation of the pump beam intensity due to the stimulated Raman loss (SRL) process is detected by a photodiode (FDS1010, Thorlabs Inc.). A lock-in amplifier with the time constant as short as 100 ns is used to demodulate the pump beam to acquire the SRS signal. SRS images (e.g., 959 cm−1 (ν1-PO43- of hydroxyapatite (HA) crystals of the tooth) [19]) of 512 × 512 pixels can be acquired within 0.3 s/frame with a 1μs pixel dwell time. The SRS imaging speed can be further boosted to video rate by using resonant scanners. One notes that other nonlinear microscopy imaging modalities (e.g., secondharmonic generation (SHG), third-harmonic generation (THG), two-photon excitation fluorescence (TPEF), coherent anti-Stokes Raman scattering (CARS)) can also be incorporated into the hyperspectral SRS imaging system (Fig. 1) to better understand the morphological architectures, biochemical structures and compositions of tissue in a comprehensive fashion [20, 21]. The choices of photomultiplier tubes (PMTs), filters and dichroic mirrors for epi-detected SHG/THG/TPEF/CARS imaging modalities have been reported in our previous work [22]. For comparison purpose, a confocal Raman microscope (Renishaw inVia, UK) equipped with a 20 × objective (N.A. = 1.0) and a 785 nm laser with an excitation power of 50 mW is utilized for micro-Raman spectroscopy and imaging of the tooth. The spectral resolution of the confocal Raman microscope is ~4 cm−1.The pixel dwell time is typically >5 s to ensure the good SNRs of Raman spectra for tooth Raman imaging. Hyperspectral SRS imaging is accomplished by scanning the pump beam wavelengths of OPO through computer control of the crystal temperature, tilt angles of Lyot filter and the cavity length of OPO, generating a three-dimensional data stack (x, y, Ω), where Ω = ωp-ωs is Raman shift. The tuning of pump wavelengths is synchronized to the frame trigger of microscope to achieve automatically spectral scanning of SRS images at different Raman shifts. The hyperspectral SRS stack is then normalized by the intensity of the pump beam at each wavelength recorded by a photodiode which detects a small fraction of the OPO output after a beam splitter. Automated wavelength tuning, SRS images acquisition, intensity normalization, and three-dimensional hyperspectral SRS data generation are controlled by using a house-built software with LabVIEW programming. For polarization-resolved SRS imaging, the polarizations of the pump and Stokes beams are rotated independently by using achromatic half-wave plates mounted in step motors. An analyzer is placed before the photodiode with the analyzer polarization parallel to the pump beam polarization. Polarization-resolved SRS images as a function of the polarization angle (θ) between the polarization directions of the pump and Stokes beams are acquired from 0° to 180° with 20° intervals. The total acquisition time of the polarized SRS image (512 × 512 pixels) is typically
