Journal of Microscopy, Vol. 254, Issue 1 2014, pp. 13–18
doi: 10.1111/jmi.12119
Received 26 July 2013; accepted 8 February 2014
Fibre-optical microendoscopy M. GU, H. BAO & H. KANG Centre for Micro-Photonics, Faculty of Science, Engineering and Technology, Swinburne University of Technology, Hawthorn, Victoria, Australia
Key words. Fibre-optical fluorescence imaging, microendoscopy, microsurgery.
Summary Microendoscopy has been an essential tool in exploring micro/nano mechanisms in vivo due to high-quality imaging performance, compact size and flexible movement. The investigations into optical fibres, micro-scanners and miniature lens have boosted efficiencies of remote light delivery to sample site and signal collection. Given the light interaction with materials in the fluorescence imaging regime, this paper reviews two classes of compact microendoscopy based on a single fibre: linear optical microendoscopy and nonlinear optical microendoscopy. Due to the fact that fluorescence occurs only in the focal volume, nonlinear optical microendoscopy can provide stronger optical sectioning ability than linear optical microendoscopy, and is a good candidate for deep tissue imaging. Moreover, one-photon excited fluorescence microendoscopy as the linear optical microendoscopy suffers from severe photobleaching owing to the linear dependence of photobleaching rate on excitation laser power. On the contrary, nonlinear optical microendoscopy, including two-photon excited fluorescence microendoscopy and second harmonic generation microendoscopy, has the capability to minimize or avoid the photobleaching effect at a high excitation power and generate high image contrast. The combination of various nonlinear signals gained by the nonlinear optical microendoscopy provides a comprehensive insight into biophenomena in internal organs. Fibre-optical microendoscopy overcomes physical limitations of traditional microscopy and opens up a new path to achieve early cancer diagnosis and microsurgery in a minimally invasive and localized manner.
Introduction The development of optical microscopy has advanced investigations into diagnosis of microstructures, especially in biological samples. However, a biopsy procedure is normally required to remove tissue for its examination owing to the Correspondence to: Min Gu, Centre for Micro-Photonics, Faculty of Science, En-
low movement flexibility and large size of optical microscopy, which introduces intrusion to animal or human, and affects original cellular structures. To address this issue, a miniature instrument is highly desired to substitute bulky microscopy instruments for in vivo diagnostics. Microendoscopy has received a great deal of interest in turning a cumbersome imaging head of microscopy into a small probe at the millimetre scale (Dickensheets et al., 1996; Seibel et al., 2002; Yelin et al., 2004; Myaing et al., 2006; Bao et al., 2008; Hoy et al., 2008; Kiesslich et al., 2008) because of the improvement of the following aspects. First, optical fibres including single-mode fibres and double-clad fibres (DCFs) (Yelin et al., 2004; Bao et al., 2008, 2009) are used to deliver light into tight space, including cells residing within hollow tissue tracts or within solid organs, and allow for handheld imaging in freely moving animals. Second, laser-scanning mechanisms employ microfabricated scanners with fast scanning rates. Third, microlenses are adopted for the imaging purpose which can greatly reduce the system size. With the integration of these three components, compact microendoscopy can be formed and be applied to in vivo imaging and even microsurgery. Today, fibre-optical microendoscopy has been well developed with good performance. Three-dimensional (3D) imaging capability with diffraction-limited resolution and a large field of view up to 475 μm × 475 μm has been demonstrated (Bao et al., 2008). The advent of new fabrication technology reduces weight and size of probe heads to subgram and millimetre (Engelbrecht et al., 2008; Zhang et al., 2012). Together with the above properties, fast scanning frame rates make real-time in vivo detection possible (Engelbrecht et al., 2008). In this review, two categories of single fibre-based optical microendoscopy (linear optical microendoscopy and nonlinear optical microendoscopy) are illustrated, which are defined with respect to the relationship between photon absorption and fluorescence generation. Moreover, their applications in biomedical areas are exploited.
Linear optical microendoscopy
gineering and Technology, Swinburne University of Technology, P. O. Box 218, Hawthorn, 3122 VIC, Australia. Tel: +61 3 9214 8776; fax: +61 3 9214 5435; e-mail: [email protected]
C 2014 The Authors C 2014 Royal Microscopical Society Journal of Microscopy
In linear optical microendoscopy, one-photon excited fluorescence (OPEF) is usually utilized as a detection signal. A
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Fig. 1. OPEF images of goblet cells at (A) surface, (B) 44 μm, (C) 88 μm, (D) 132 μm and (E) 176 μm under the mouse intestine epithelium. Size of the images: 200 μm × 200 μm (Bao et al., 2009).
488-nm continuous-wave laser beam is coupled into the core of a single-mode fibre (SMF) and is delivered to a microendoscopy probe. Inside the probe, multiple-element microlenses focus a laser beam from the output of the SMF into a specimen. OPEF generated at the laser focal spot is collected backward through the microlenses and the core of the SMF. The core of the SMF acts as a small aperture for illumination, as well as the spatial pin hole for confocal imaging. Only fluorescence returning from the focus can be efficiently recaptured into the SMF for detection, eliminating complicated optical alignments required in bulky optical confocal microscopes. Based on two-dimensional scanning with horizontally and vertically driven coils, the z-axis movement can be performed by a shape memory alloy to achieve 3D imaging built up with a scanning area of 500 μm × 500 μm and a penetration depth of 250 μm (Kiesslich et al., 2008). Images are acquired point by point at selected depth and reconstructed with a computer, allowing for 3D reconstruction of topologically complex objects (Goetz et al., 2007). Intestinal metaplasia of the stomach is a premalignant lesion of gastric cancer and is considered to be the earliest indicator of potential malignant progression to gastric cancer. Goblet cells appearing in the epithelial layer of the colon and intestine are an indicator of pathognomonic for intestinal metaplasia. As a consequence, 3D imaging of goblet cells is of significance for early cancer detection. The capability of this compact OPEF microendoscopy has been demonstrated in 3D viewing of human and mouse gastrointestinal system in vivo and differentiating epithelial gaps and goblet cells (Bao et al., 2009). A mouse intestine with many goblet cells is used as a model of intestinal metaplasia. Between 0.3 and 0.4 mL of the fluorescein (1% solution diluted in saline) was intravenously injected into the mouse intestine to enable the visualization of most cells. The appearance of goblet cells can be shown as dark void structures because of the mucin contained in their cytoplasm. Figure 1 displays the 3D confocal microendoscopy images of goblet cells under the mouse intestine epithelium. Although the image at a depth down to 176 μm is obtained, only the surface image of mouse intestine is clear. Moreover, severe photobleaching of the fluorescein occurs in the presence of high excitation energy fluence because the photobleaching rate for one-photon excitation has a linear
Fig. 2. OPEF image of the mouse intestine with 440 μW power on the sample. (A) First scan, (B) scanned for 5 min and (C) scanned for 10 min. Size of the images: 200 μm × 200 μm (Bao et al., 2009).
dependence on the excitation laser power. Figure 2 presents OPEF images of the mouse intestine with 440 μW power on the sample after different amounts of the scanning time. It can be seen that the OPEF intensity is decreased dramatically with the rising of the scanning time, confirming that high laser power induces high photobleaching in OPEF imaging. Nonlinear optical microendoscopy Nonlinear optical microendoscopy employs signals generated from nonlinear optical processes including two-photon excitation (TPE), second-harmonic generation (SHG) or other third-harmonic generation to visualize cellular layers and intracellular organs with deep penetration depth. These signals provide valuable information for diagnosing diseases at an early stage and understanding complex mechanisms of biophenomena in living animals (Denk et al., 1990; Brown et al., 2003; Campagnola et al., 2003). Similar to linear optical microendoscopy, nonlinear optical microendoscopy can use an optical fibre to offer a comfortable distance between a target and optical and electronic hardware for imaging internal organs. Although an SMF was first tried for fibre-optic two-photon fluorescence imaging system (Bird et al., 2002), its main drawbacks are (1) severe temporal and spectral broadening of ultrashort pulses due to the chromatic dispersion and the nonlinearity and (2) the insufficient collection of nonlinear signals associated with the low numerical aperture. Photonic crystal fibres utilizing a unique manner of light guidance have demonstrated capabilities of improving both the efficiencies of the excitation beam delivery to samples and nonlinear signal collection. Apart from the research into optical fibres, a C 2014 The Authors C 2014 Royal Microscopical Society, 254, 13–18 Journal of Microscopy
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Fig. 3. Schematic diagram of nonlinear microendoscopy (Bao et al., 2008).
miniature scanning mechanism including microelectromechanical systems mirrors and a piezoelectric or an electromagnetic actuators have found their applications in scanning the light coupled from fibres in the compact microendoscopy head to generate a high-resolution two-dimensional image (Fu et al., 2007). To date, another kind of fibres, a DCF has been incorporated with a tubular piezoelectric actuator to form a microendoscope, which has been proven to be an excellent candidate for nonlinear optical imaging (Bao et al., 2008, 2009). Figure 3 shows the schematic setup of a typical nonlinear optical microendoscope based on a DCF (Bao et al., 2008). A laser beam with ultrashort pulses is used as an excitation source. Near-infrared wavelengths can be selected to excite nonlinear optical signals for a deep penetration depth because of its low absorption and scattering coefficients in tissue. In order to compensate for chromatic dispersion of the optical fibre, a prechirp unit such as a grating pair is adopted. Then the laser beam is coupled to the core of the DCF, which supports a singlemode light propagation, and delivered to a handheld probe for nonlinear excited fluorescence generation. The nonlinear signal from a sample is collected by the DCF in the backward direction and separated from the excitation light by a dichroic mirror before coupled by an objective via a multimode fibre to a photomultiplier tube. In a handheld probe, a custom design lens (diameter 3 mm, numerical aperture 0.35) which consists of multiple elements to accurately correct image aberration is employed for imaging, and a 3D micro-scanner is designed for scanning the DCF line by line and moving the DCF in the axial direction to realize images at different penetration depth. The nonlinear signal is synchronized with the 3D micro-scanner R by a modified image processor FIVE1 (Optiscan Pty Ltd, Victoria, Australia) to build up an image dimension frame C 2014 The Authors C 2014 Royal Microscopical Society, 254, 13–18 Journal of Microscopy
by frame up to a field of view of 475 μm × 475 μm × 250 μm (Bao et al., 2008). In the two-photon excited fluorescence (TPEF) microendoscopy system, two photons are absorbed simultaneously by fluorophores to excite an electron into a higher energy state, from which it decays and emits a fluorescence signal. As the intensity of the TPEF has a quadratic dependence on the peak intensity of the excitation light, the TPE can be highly localized at the focus. In this way, out-of-focus fluorescence is blocked and optical sectioning is realized. These features provide the ability to reduce photobleaching and photodamage of specimen and realize 3D imaging and prediagnosis of various kinds of diseases (Denk et al., 1990; Flusberg, 2005). To make a comparison between OPEF and TPEF techniques, the same specimen used in OPEF imaging is employed for TPEF imaging. The wavelength of 790 nm is employed for TPE, which coincides with the peak of TPE spectrum of fluorescein. Figure 4 shows the TPEF images of mouse intestine with the nonlinear microendoscopy (Bao et al., 2009). Although both OPEF microendoscopy images and TPEF microendoscopy images show clear surface images of mouse intestine, more goblet cells can be observed from TPEF imaging. Moreover, the gland structure of OPEF imaging starts to be blurred at 132 μm deep (Fig. 1D) while the gland structure can still be identified at 176 μm deep from TPEF imaging, demonstrating that TPEF has a higher sectioning ability. Less photobleaching is another advantage of TPEF imaging over OPEF imaging. From Figure 5, it can be seen that no obvious change of TPEF intensity is presented after scanning for 5 and 10 min at the excitation laser power of 38 mW, which is approximately two orders of magnitude larger than the excitation power used in OPEF imaging.
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Fig. 4. TPEF images of goblet cells at (A) surface, (B) 44 μm, (C) 88 μm, (D) 132 μm and (E) 176 μm under the mouse intestine epithelium. Size of the images: 200 μm × 200 μm (Bao et al., 2009).
Fig. 5. TPEF images of the mouse intestine with 38 mW power on the sample. (A) First scan, (B) scanned for 5 min and (C) scanned for 10 min. Size of the images: 200 μm × 200 μm (Bao et al., 2009).
SHG, another stream of nonlinear optical processes, involves simultaneous interaction of two photons with noncentrosymmetrical structures without absorption, producing radiation at exactly half of the excitation wavelength. Unlike the TPEF imaging, neither excitation of endogenous florescent molecules nor phototoxicity effects or photobleaching is observed in the SHG process. In particular, SHG polarization anisotropy utilizing polarization-dependent SHG signals is favourable in determining orientations and organization degrees of proteins in tissues (Brown et al., 2003; Campagnola et al., 2003). The establishment of SHG as a nondestructive imaging modality holds a promise for both basic research and relevance to clinical pathology. The SHG imaging with fibreoptical microendoscopy has a substantial impact on studies in biology and medicine, including tissue organization, wound healing, myofilament assembly, muscle development and disease, aging and the division cycles of normal and cancerous cells. By combining TPEF imaging and SHG imaging, the morphological and functional information of biological structures are obtained to get a comprehensive understanding of samples in vivo. The separate investigations into these two kinds of information can be realized by switching two filters suitable for detecting TPEF and SHG signals only. Figure 6 reveals twochannel images of a mouse-tail tendon with high resolution (Bao et al., 2010). It is obvious that the difference between the SHG images and the TPEF images can be identified. The TPEF images show the distribution of fluorescein in the mousetail tendon absorbed from the bloodstream, while SHG images display noncentrosymmetrical orientations and structures of proteins in the mouse-tail tendon.
Fig. 6. Two-channel images of a mouse-tail tendon by a nonlinear microendoscope. (A), (D) SHG images. (B), (E) TPEF images. (C), (F) Combination of two-channel images. Red: TPEF image. Green: SHG image. The size of the images: 150 μm × 150 μm. Arrows are the polarization directions of the excitation laser beam (Bao et al., 2010).
Microendoscopy for microsurgery A lot of efforts have been made to ease pains of people suffering from cancer. Chemotherapy, radiation therapy and surgery are the most commonly used methods as cancer management options in hospitals. However, shortcomings associated with them, including causing toxicity to adjacent healthy tissue and removing a wide surgical margin or a free margin, always produce side effects to patients. As a consequence, an approach to curing cancer cells in a minimally invasive and observable manner without damage to the surrounding area is essential. One of the promising treatment methods is microsurgery of cancer cells using femtosecond pulsed laser beams and optical microscopy, which has allowed for ablation of cancer cells and 3D imaging through thick tissue media. However, the utilization of bulky optics and multilaser systems has been a main hurdle for femtosecond-laser-based microscopes to be potentially useful in in vivo and noninvasive environments with a desirable 3D selectivity and specificity. Nonlinear optical microendoscopy is a powerful tool for 3D in vivo imaging and can be utilized to deliver a femtosecond pulsed beam for microsurgery (Hoy et al., 2008; Gu et al., 2010). Transferrin-conjugated gold nanorods can be C 2014 The Authors C 2014 Royal Microscopical Society, 254, 13–18 Journal of Microscopy
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Fig. 7. Three-dimensional imaging and localized necrosis of cancer cells in the collagen matrix tissue phantom mixed with propidium Iodide. Two-photon photoluminescence image of HeLa cells labelled with biofunctional gold nanorods at (A1) the collagen surface, (B1) 20 μm and (C1) 40 μm below the surface. HeLa cells were shown to be selectively killed by the focused excitation laser beam at (A2) the collagen surface, (B2) 20 μm and (C2) 40 μm below the surface (Gu et al., 2010).
attached specifically to molecular abnormalities in certain tumors. Under this circumstance, they are employed not only as a targeted therapeutic ‘drug’ for specific cancer treatment, but also as targeted ‘light’ for identifying cancer cells. The cancer treatment can be undertaken by inducing instant damage through the necrosis process, or by stopping cell proliferation through the apoptosis process. Typically, necrosis occurs when strong external factors are applied. As a result, cell membrane cannot be maintained. On the contrary, apoptosis is a much less severe cell destruction process without tissue inflammation and it can be controlled by both intra- and extracellular signals (Majno et al., 1995; Oleinick et al., 2002). By controlling the power level of a single femtosecond pulsed laser beam, minimally invasive single cancer-cell necrosis and apoptosis are realized, respectively (Fig. 7) (Li et al., 2008; Gu et al., 2010). Figures 7(A1)–(C1) show TPE photoluminescence images of cancer cells (HeLa). The images clearly reveal individual cancer cells at different depths. The HeLa cells without gold nanorod labelling cannot be detected by the microendoscope. With the help of this visualization method of cancer cells, the gold-nanorod enhanced cancer-cell necrosis process can be investigated. To this end, live HeLa cells labelled by transferrinconjugated gold nanorods were placed in the collagen matrix tissue phantom mixed with propidium iodide. Propidium iodide can be used to examine the death of cells after being irradiated by an excitation laser beam because it only stains the nucleuses of dead cells. The whole process of the single cancer-cell microsurgery was completed by imaging the cancer cells first, and focusing the laser beam to a selected single cell for a period of time (Fig. 7). The two-photon-excited photoluminescence image (Fig. 7A1) of the live cancer cells at the surface of the collagen was obtained using the laser beam at the wavelength of 790 nm, which coincides with the longitudinal surface plas C 2014 The Authors C 2014 Royal Microscopical Society, 254, 13–18 Journal of Microscopy
mon resonance wavelength of the transferrin-conjugated gold nanorods. These cells forming an ‘L’ shape were selected. The excitation laser beam at 5 mW was focused on each of those chosen cells with a dwelling time of 0.5 s. Six minutes after the entire selected cells were illuminated, the nonlinear fluorescence microendoscope was used to examine the cell death. The excitation wavelength was changed to 740 nm (Fig. 7A2), which is the peak two-photon absorption wavelength of propidium iodide. Figure 7(A2) shows a clear ‘L’ shape in the field of view and confirms that the selected cells irradiated with a dwelling time of 0.5 s were dead without destructing the rest of the cells. The localized necrosis effect below the surface of the collagen can be observed at different penetration depths. Figures 7(B1), (B2), (C1)and (C2) present two-photon-excited photoluminescence images at 20 and 40 μm below the surface, respectively. It can be seen that the cells outside the selected laser irradiated area were not affected at all, indicating that the microendoscope can be used for the highly localized and selective 3D necrosis treatment of cancer cells, especially at early stages before it metastasizes. The power of 35 mW, which is far below the medical safety level of 100 mJ cm–2 (American Laser Institute, 2000), is required to damage nongold-nanorod-labelling cancer cells (Gu et al., 2010). It should be noted that the power and energy threshold for cancer-cell necrosis are significantly decreased by using the biofunctional gold nanorods. A 100% death rate of a cancer cell can be confirmed through necrosis or apoptosis processes when the laser power is higher than 3 mW, corresponding to an energy fluence of 0.53 mJ cm–2 (Gu et al., 2010). Conclusion The research work reviewed in this paper has demonstrated the applications of fibre-optical microendoscopy in imaging
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fields. 3D optical sectioning abilities through cells have been confirmed. Compared with OPEF microendoscopy, the advent of TPEF and SHG microendoscopy technique paves a way to improve endoscopic image contrast and image resolution and observe functional and morphological changes of microstructures in vivo. With compact microendoscopy, endoscopists can be provided with clear and robust visualization of biological samples and the detection of small and subtle dysplastic lesions to enhance the early detection, diagnosis and staging of gastric cancers. This, in turn, will facilitate the early detection of cancers at the stage of dysplasia and the enhanced cure rate of cancer patients. In this regard, two specific functionalities of fibre-optical microendoscopy, superresolution imaging (Gu et al., 2014a) and cellular manipulation or tweezing (Gu et al., 2014b), are highly desirable.
References American Laser Institute (2000) American National Standard for Safe Use of Lasers ANSI Z136.1. Orlando, Florida. Bao, H., Allen, J., Pattie, R., Vance, R. & Gu, M. (2008) Fast handhold two-photon fluorescence microendoscope with a 475μm × 475μm field of view for in vivo imaging. Opt. Lett. 33, 1333–1335. Bao, H., Boussioutas, A., Reynolds, J., Russell, S. & Gu, M. (2009) Imaging of goblet cells as a marker for intestinal metaplasia of the stomach by one-photon and two-photon fluorescence endomicroscopy. J. Biomed. Opt. 14, 064031–1–4. Bao, H., Boussioutas, A., Reynolds, J., Russell, S. & Gu, M. (2010) Second harmonic generation imaging via nonlinear endomicroscopy. Opt. Express 18, 1255–1260. Bird, D. & Gu, M. (2002) Fibre-optic two-photon scanning fluorescence microscopy. J. Microsc. 208, 35–48. Brown, E., McKee, T., diTomaso, E., Pluen, A., Seed, B., Boucher, Y. & Jain, R.K. (2003) Dynamic imaging of collagen and its modulation in tumors in vivo using second-harmonic generation. Nat. Med. 9, 796–800. Campagnola, P.J. & Loew, L.M. (2003) Second-harmonic imaging microscopy for visualizing biomolecular arrays in cells, tissues and organisms. Nat. Biotechnol. 21, 1356–1360. Denk, W., Strickler, J.H. & Webb, W.W. (1990) Two-photon laser scanning fluorescence microscopy. Science 248, 73–76. Dickensheets, D.L. & Kino, G.S. (1996) Micromachined scanning confocal optical microscope. Opt. Lett. 21, 764–766.
Engelbrecht, C.J., Johnston, R.S., Seibel, E.J. & Helmchen, F. (2008) Ultracompact fiber-optic two-photon microscope for functional fluorescence imaging in vivo. Opt. Express 16, 5556–5564. Flusberg, B.A., Cocker, E.D., Piyawattanametha, W., Jung, J.C., Cheung, E.L.M. & Schnitzer, M.J. (2005) Fiber-optic fluorescence imaging. Nat. Methods 2, 941–950. Fu, L. & Gu, M. (2007) Fibre-optic nonlinear optical microscopy and endoscopy. J. Microsc. 226, 195–206. Goetz, M., Fottner, C., Schirrmacher, E. et al. (2007) In-vivo confocal realtime mini-microscopy in animal models of human inflammatory and neoplastic diseases. Endoscopy 39, 350–356. Gu, M., Bao, H. & Li, J. (2010) Cancer-cell microsurgery using nonlinear optical endomicroscopy. J. Biomed. Opt. 15, 050502–1–3. Gu, M., Bao, H., Gan, X., Stokes, N. & Wu, J. (2014a) Tweezing and manipulating micro- and nanoparticles by optical nonlinear endoscopy. Light: Sci. Appl. 3, doi:10.1038/lsa.2014.7. Gu, M., Kang, H. & Li, X. (2014b) Breaking the diffraction-limited resolution barrier in fiber-optical two-photon fluorescence endoscopy by an azimuthally-polarized beam. Sci. Rep. 4, 3627. Hoy, C.L., Durr, N.J., Chen, P., Piyawattanametha, W., Ra, H., Solgaard, O. & Ben-Yakar, A. (2008) Miniaturized probe for femtosecond laser microsurgery and two-photon imaging. Opt. Express 16, 9996– 10005. Kiesslich, R., Galle, P.R. & Neurath, M. (2008) Atlas of Endomicroscopy. Springer, Heidelberg. Li, J., Day, D. & Gu, M. (2008) Ultra-low energy threshold for cancer photothermal therapy of cancer using transferrin-conjugated gold nanorods. Adv. Mater. 20, 3866–3871. Majno, G. & Joris, I. (1995) Apoptosis, oncosis, and necrosis. An overview of cell death. Am. J. Pathol. 146, 3–15. Myaing, M., MacDonald, D.J. & Li, X. (2006) Fiber-optic scanning twophoton fluorescence endoscope. Opt. Lett. 31, 1076–1078. Oleinick, N.L., Morris, R.L. & Belichenko, I. (2002) The role of apoptosis in response to photodynamic therapy: what, where, why, and how. Photochem. Photobiol. Sci. 1, 1–21. Seibel, E.J. & Smithwick, Q.Y. (2002) Unique features of optical scanning, single fiber endoscopy. Lasers Surg. Med. 30, 177– 183. Yelin, D., Bouma, B.E., Yun, S.H. & Tearney, G.J. (2004) Double-clad fiber for endoscopy. Opt. Lett. 29, 2408–2410. Zhang, Y., Akins, M.L., Murari, K., Xi, J., Li, M., Luby-Phelps, K., Mahendroo, M. & Li, X. (2012) A compact fiber-optic SHG scanning endomicroscope and its application to visualize cervical remodeling during pregnancy. Proc. Natl. Acad. Sci. 109, 12878–12883.
C 2014 The Authors C 2014 Royal Microscopical Society, 254, 13–18 Journal of Microscopy
doi: 10.1111/jmi.12119
Received 26 July 2013; accepted 8 February 2014
Fibre-optical microendoscopy M. GU, H. BAO & H. KANG Centre for Micro-Photonics, Faculty of Science, Engineering and Technology, Swinburne University of Technology, Hawthorn, Victoria, Australia
Key words. Fibre-optical fluorescence imaging, microendoscopy, microsurgery.
Summary Microendoscopy has been an essential tool in exploring micro/nano mechanisms in vivo due to high-quality imaging performance, compact size and flexible movement. The investigations into optical fibres, micro-scanners and miniature lens have boosted efficiencies of remote light delivery to sample site and signal collection. Given the light interaction with materials in the fluorescence imaging regime, this paper reviews two classes of compact microendoscopy based on a single fibre: linear optical microendoscopy and nonlinear optical microendoscopy. Due to the fact that fluorescence occurs only in the focal volume, nonlinear optical microendoscopy can provide stronger optical sectioning ability than linear optical microendoscopy, and is a good candidate for deep tissue imaging. Moreover, one-photon excited fluorescence microendoscopy as the linear optical microendoscopy suffers from severe photobleaching owing to the linear dependence of photobleaching rate on excitation laser power. On the contrary, nonlinear optical microendoscopy, including two-photon excited fluorescence microendoscopy and second harmonic generation microendoscopy, has the capability to minimize or avoid the photobleaching effect at a high excitation power and generate high image contrast. The combination of various nonlinear signals gained by the nonlinear optical microendoscopy provides a comprehensive insight into biophenomena in internal organs. Fibre-optical microendoscopy overcomes physical limitations of traditional microscopy and opens up a new path to achieve early cancer diagnosis and microsurgery in a minimally invasive and localized manner.
Introduction The development of optical microscopy has advanced investigations into diagnosis of microstructures, especially in biological samples. However, a biopsy procedure is normally required to remove tissue for its examination owing to the Correspondence to: Min Gu, Centre for Micro-Photonics, Faculty of Science, En-
low movement flexibility and large size of optical microscopy, which introduces intrusion to animal or human, and affects original cellular structures. To address this issue, a miniature instrument is highly desired to substitute bulky microscopy instruments for in vivo diagnostics. Microendoscopy has received a great deal of interest in turning a cumbersome imaging head of microscopy into a small probe at the millimetre scale (Dickensheets et al., 1996; Seibel et al., 2002; Yelin et al., 2004; Myaing et al., 2006; Bao et al., 2008; Hoy et al., 2008; Kiesslich et al., 2008) because of the improvement of the following aspects. First, optical fibres including single-mode fibres and double-clad fibres (DCFs) (Yelin et al., 2004; Bao et al., 2008, 2009) are used to deliver light into tight space, including cells residing within hollow tissue tracts or within solid organs, and allow for handheld imaging in freely moving animals. Second, laser-scanning mechanisms employ microfabricated scanners with fast scanning rates. Third, microlenses are adopted for the imaging purpose which can greatly reduce the system size. With the integration of these three components, compact microendoscopy can be formed and be applied to in vivo imaging and even microsurgery. Today, fibre-optical microendoscopy has been well developed with good performance. Three-dimensional (3D) imaging capability with diffraction-limited resolution and a large field of view up to 475 μm × 475 μm has been demonstrated (Bao et al., 2008). The advent of new fabrication technology reduces weight and size of probe heads to subgram and millimetre (Engelbrecht et al., 2008; Zhang et al., 2012). Together with the above properties, fast scanning frame rates make real-time in vivo detection possible (Engelbrecht et al., 2008). In this review, two categories of single fibre-based optical microendoscopy (linear optical microendoscopy and nonlinear optical microendoscopy) are illustrated, which are defined with respect to the relationship between photon absorption and fluorescence generation. Moreover, their applications in biomedical areas are exploited.
Linear optical microendoscopy
gineering and Technology, Swinburne University of Technology, P. O. Box 218, Hawthorn, 3122 VIC, Australia. Tel: +61 3 9214 8776; fax: +61 3 9214 5435; e-mail: [email protected]
C 2014 The Authors C 2014 Royal Microscopical Society Journal of Microscopy
In linear optical microendoscopy, one-photon excited fluorescence (OPEF) is usually utilized as a detection signal. A
14
M. GU ET AL.
Fig. 1. OPEF images of goblet cells at (A) surface, (B) 44 μm, (C) 88 μm, (D) 132 μm and (E) 176 μm under the mouse intestine epithelium. Size of the images: 200 μm × 200 μm (Bao et al., 2009).
488-nm continuous-wave laser beam is coupled into the core of a single-mode fibre (SMF) and is delivered to a microendoscopy probe. Inside the probe, multiple-element microlenses focus a laser beam from the output of the SMF into a specimen. OPEF generated at the laser focal spot is collected backward through the microlenses and the core of the SMF. The core of the SMF acts as a small aperture for illumination, as well as the spatial pin hole for confocal imaging. Only fluorescence returning from the focus can be efficiently recaptured into the SMF for detection, eliminating complicated optical alignments required in bulky optical confocal microscopes. Based on two-dimensional scanning with horizontally and vertically driven coils, the z-axis movement can be performed by a shape memory alloy to achieve 3D imaging built up with a scanning area of 500 μm × 500 μm and a penetration depth of 250 μm (Kiesslich et al., 2008). Images are acquired point by point at selected depth and reconstructed with a computer, allowing for 3D reconstruction of topologically complex objects (Goetz et al., 2007). Intestinal metaplasia of the stomach is a premalignant lesion of gastric cancer and is considered to be the earliest indicator of potential malignant progression to gastric cancer. Goblet cells appearing in the epithelial layer of the colon and intestine are an indicator of pathognomonic for intestinal metaplasia. As a consequence, 3D imaging of goblet cells is of significance for early cancer detection. The capability of this compact OPEF microendoscopy has been demonstrated in 3D viewing of human and mouse gastrointestinal system in vivo and differentiating epithelial gaps and goblet cells (Bao et al., 2009). A mouse intestine with many goblet cells is used as a model of intestinal metaplasia. Between 0.3 and 0.4 mL of the fluorescein (1% solution diluted in saline) was intravenously injected into the mouse intestine to enable the visualization of most cells. The appearance of goblet cells can be shown as dark void structures because of the mucin contained in their cytoplasm. Figure 1 displays the 3D confocal microendoscopy images of goblet cells under the mouse intestine epithelium. Although the image at a depth down to 176 μm is obtained, only the surface image of mouse intestine is clear. Moreover, severe photobleaching of the fluorescein occurs in the presence of high excitation energy fluence because the photobleaching rate for one-photon excitation has a linear
Fig. 2. OPEF image of the mouse intestine with 440 μW power on the sample. (A) First scan, (B) scanned for 5 min and (C) scanned for 10 min. Size of the images: 200 μm × 200 μm (Bao et al., 2009).
dependence on the excitation laser power. Figure 2 presents OPEF images of the mouse intestine with 440 μW power on the sample after different amounts of the scanning time. It can be seen that the OPEF intensity is decreased dramatically with the rising of the scanning time, confirming that high laser power induces high photobleaching in OPEF imaging. Nonlinear optical microendoscopy Nonlinear optical microendoscopy employs signals generated from nonlinear optical processes including two-photon excitation (TPE), second-harmonic generation (SHG) or other third-harmonic generation to visualize cellular layers and intracellular organs with deep penetration depth. These signals provide valuable information for diagnosing diseases at an early stage and understanding complex mechanisms of biophenomena in living animals (Denk et al., 1990; Brown et al., 2003; Campagnola et al., 2003). Similar to linear optical microendoscopy, nonlinear optical microendoscopy can use an optical fibre to offer a comfortable distance between a target and optical and electronic hardware for imaging internal organs. Although an SMF was first tried for fibre-optic two-photon fluorescence imaging system (Bird et al., 2002), its main drawbacks are (1) severe temporal and spectral broadening of ultrashort pulses due to the chromatic dispersion and the nonlinearity and (2) the insufficient collection of nonlinear signals associated with the low numerical aperture. Photonic crystal fibres utilizing a unique manner of light guidance have demonstrated capabilities of improving both the efficiencies of the excitation beam delivery to samples and nonlinear signal collection. Apart from the research into optical fibres, a C 2014 The Authors C 2014 Royal Microscopical Society, 254, 13–18 Journal of Microscopy
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Fig. 3. Schematic diagram of nonlinear microendoscopy (Bao et al., 2008).
miniature scanning mechanism including microelectromechanical systems mirrors and a piezoelectric or an electromagnetic actuators have found their applications in scanning the light coupled from fibres in the compact microendoscopy head to generate a high-resolution two-dimensional image (Fu et al., 2007). To date, another kind of fibres, a DCF has been incorporated with a tubular piezoelectric actuator to form a microendoscope, which has been proven to be an excellent candidate for nonlinear optical imaging (Bao et al., 2008, 2009). Figure 3 shows the schematic setup of a typical nonlinear optical microendoscope based on a DCF (Bao et al., 2008). A laser beam with ultrashort pulses is used as an excitation source. Near-infrared wavelengths can be selected to excite nonlinear optical signals for a deep penetration depth because of its low absorption and scattering coefficients in tissue. In order to compensate for chromatic dispersion of the optical fibre, a prechirp unit such as a grating pair is adopted. Then the laser beam is coupled to the core of the DCF, which supports a singlemode light propagation, and delivered to a handheld probe for nonlinear excited fluorescence generation. The nonlinear signal from a sample is collected by the DCF in the backward direction and separated from the excitation light by a dichroic mirror before coupled by an objective via a multimode fibre to a photomultiplier tube. In a handheld probe, a custom design lens (diameter 3 mm, numerical aperture 0.35) which consists of multiple elements to accurately correct image aberration is employed for imaging, and a 3D micro-scanner is designed for scanning the DCF line by line and moving the DCF in the axial direction to realize images at different penetration depth. The nonlinear signal is synchronized with the 3D micro-scanner R by a modified image processor FIVE1 (Optiscan Pty Ltd, Victoria, Australia) to build up an image dimension frame C 2014 The Authors C 2014 Royal Microscopical Society, 254, 13–18 Journal of Microscopy
by frame up to a field of view of 475 μm × 475 μm × 250 μm (Bao et al., 2008). In the two-photon excited fluorescence (TPEF) microendoscopy system, two photons are absorbed simultaneously by fluorophores to excite an electron into a higher energy state, from which it decays and emits a fluorescence signal. As the intensity of the TPEF has a quadratic dependence on the peak intensity of the excitation light, the TPE can be highly localized at the focus. In this way, out-of-focus fluorescence is blocked and optical sectioning is realized. These features provide the ability to reduce photobleaching and photodamage of specimen and realize 3D imaging and prediagnosis of various kinds of diseases (Denk et al., 1990; Flusberg, 2005). To make a comparison between OPEF and TPEF techniques, the same specimen used in OPEF imaging is employed for TPEF imaging. The wavelength of 790 nm is employed for TPE, which coincides with the peak of TPE spectrum of fluorescein. Figure 4 shows the TPEF images of mouse intestine with the nonlinear microendoscopy (Bao et al., 2009). Although both OPEF microendoscopy images and TPEF microendoscopy images show clear surface images of mouse intestine, more goblet cells can be observed from TPEF imaging. Moreover, the gland structure of OPEF imaging starts to be blurred at 132 μm deep (Fig. 1D) while the gland structure can still be identified at 176 μm deep from TPEF imaging, demonstrating that TPEF has a higher sectioning ability. Less photobleaching is another advantage of TPEF imaging over OPEF imaging. From Figure 5, it can be seen that no obvious change of TPEF intensity is presented after scanning for 5 and 10 min at the excitation laser power of 38 mW, which is approximately two orders of magnitude larger than the excitation power used in OPEF imaging.
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Fig. 4. TPEF images of goblet cells at (A) surface, (B) 44 μm, (C) 88 μm, (D) 132 μm and (E) 176 μm under the mouse intestine epithelium. Size of the images: 200 μm × 200 μm (Bao et al., 2009).
Fig. 5. TPEF images of the mouse intestine with 38 mW power on the sample. (A) First scan, (B) scanned for 5 min and (C) scanned for 10 min. Size of the images: 200 μm × 200 μm (Bao et al., 2009).
SHG, another stream of nonlinear optical processes, involves simultaneous interaction of two photons with noncentrosymmetrical structures without absorption, producing radiation at exactly half of the excitation wavelength. Unlike the TPEF imaging, neither excitation of endogenous florescent molecules nor phototoxicity effects or photobleaching is observed in the SHG process. In particular, SHG polarization anisotropy utilizing polarization-dependent SHG signals is favourable in determining orientations and organization degrees of proteins in tissues (Brown et al., 2003; Campagnola et al., 2003). The establishment of SHG as a nondestructive imaging modality holds a promise for both basic research and relevance to clinical pathology. The SHG imaging with fibreoptical microendoscopy has a substantial impact on studies in biology and medicine, including tissue organization, wound healing, myofilament assembly, muscle development and disease, aging and the division cycles of normal and cancerous cells. By combining TPEF imaging and SHG imaging, the morphological and functional information of biological structures are obtained to get a comprehensive understanding of samples in vivo. The separate investigations into these two kinds of information can be realized by switching two filters suitable for detecting TPEF and SHG signals only. Figure 6 reveals twochannel images of a mouse-tail tendon with high resolution (Bao et al., 2010). It is obvious that the difference between the SHG images and the TPEF images can be identified. The TPEF images show the distribution of fluorescein in the mousetail tendon absorbed from the bloodstream, while SHG images display noncentrosymmetrical orientations and structures of proteins in the mouse-tail tendon.
Fig. 6. Two-channel images of a mouse-tail tendon by a nonlinear microendoscope. (A), (D) SHG images. (B), (E) TPEF images. (C), (F) Combination of two-channel images. Red: TPEF image. Green: SHG image. The size of the images: 150 μm × 150 μm. Arrows are the polarization directions of the excitation laser beam (Bao et al., 2010).
Microendoscopy for microsurgery A lot of efforts have been made to ease pains of people suffering from cancer. Chemotherapy, radiation therapy and surgery are the most commonly used methods as cancer management options in hospitals. However, shortcomings associated with them, including causing toxicity to adjacent healthy tissue and removing a wide surgical margin or a free margin, always produce side effects to patients. As a consequence, an approach to curing cancer cells in a minimally invasive and observable manner without damage to the surrounding area is essential. One of the promising treatment methods is microsurgery of cancer cells using femtosecond pulsed laser beams and optical microscopy, which has allowed for ablation of cancer cells and 3D imaging through thick tissue media. However, the utilization of bulky optics and multilaser systems has been a main hurdle for femtosecond-laser-based microscopes to be potentially useful in in vivo and noninvasive environments with a desirable 3D selectivity and specificity. Nonlinear optical microendoscopy is a powerful tool for 3D in vivo imaging and can be utilized to deliver a femtosecond pulsed beam for microsurgery (Hoy et al., 2008; Gu et al., 2010). Transferrin-conjugated gold nanorods can be C 2014 The Authors C 2014 Royal Microscopical Society, 254, 13–18 Journal of Microscopy
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Fig. 7. Three-dimensional imaging and localized necrosis of cancer cells in the collagen matrix tissue phantom mixed with propidium Iodide. Two-photon photoluminescence image of HeLa cells labelled with biofunctional gold nanorods at (A1) the collagen surface, (B1) 20 μm and (C1) 40 μm below the surface. HeLa cells were shown to be selectively killed by the focused excitation laser beam at (A2) the collagen surface, (B2) 20 μm and (C2) 40 μm below the surface (Gu et al., 2010).
attached specifically to molecular abnormalities in certain tumors. Under this circumstance, they are employed not only as a targeted therapeutic ‘drug’ for specific cancer treatment, but also as targeted ‘light’ for identifying cancer cells. The cancer treatment can be undertaken by inducing instant damage through the necrosis process, or by stopping cell proliferation through the apoptosis process. Typically, necrosis occurs when strong external factors are applied. As a result, cell membrane cannot be maintained. On the contrary, apoptosis is a much less severe cell destruction process without tissue inflammation and it can be controlled by both intra- and extracellular signals (Majno et al., 1995; Oleinick et al., 2002). By controlling the power level of a single femtosecond pulsed laser beam, minimally invasive single cancer-cell necrosis and apoptosis are realized, respectively (Fig. 7) (Li et al., 2008; Gu et al., 2010). Figures 7(A1)–(C1) show TPE photoluminescence images of cancer cells (HeLa). The images clearly reveal individual cancer cells at different depths. The HeLa cells without gold nanorod labelling cannot be detected by the microendoscope. With the help of this visualization method of cancer cells, the gold-nanorod enhanced cancer-cell necrosis process can be investigated. To this end, live HeLa cells labelled by transferrinconjugated gold nanorods were placed in the collagen matrix tissue phantom mixed with propidium iodide. Propidium iodide can be used to examine the death of cells after being irradiated by an excitation laser beam because it only stains the nucleuses of dead cells. The whole process of the single cancer-cell microsurgery was completed by imaging the cancer cells first, and focusing the laser beam to a selected single cell for a period of time (Fig. 7). The two-photon-excited photoluminescence image (Fig. 7A1) of the live cancer cells at the surface of the collagen was obtained using the laser beam at the wavelength of 790 nm, which coincides with the longitudinal surface plas C 2014 The Authors C 2014 Royal Microscopical Society, 254, 13–18 Journal of Microscopy
mon resonance wavelength of the transferrin-conjugated gold nanorods. These cells forming an ‘L’ shape were selected. The excitation laser beam at 5 mW was focused on each of those chosen cells with a dwelling time of 0.5 s. Six minutes after the entire selected cells were illuminated, the nonlinear fluorescence microendoscope was used to examine the cell death. The excitation wavelength was changed to 740 nm (Fig. 7A2), which is the peak two-photon absorption wavelength of propidium iodide. Figure 7(A2) shows a clear ‘L’ shape in the field of view and confirms that the selected cells irradiated with a dwelling time of 0.5 s were dead without destructing the rest of the cells. The localized necrosis effect below the surface of the collagen can be observed at different penetration depths. Figures 7(B1), (B2), (C1)and (C2) present two-photon-excited photoluminescence images at 20 and 40 μm below the surface, respectively. It can be seen that the cells outside the selected laser irradiated area were not affected at all, indicating that the microendoscope can be used for the highly localized and selective 3D necrosis treatment of cancer cells, especially at early stages before it metastasizes. The power of 35 mW, which is far below the medical safety level of 100 mJ cm–2 (American Laser Institute, 2000), is required to damage nongold-nanorod-labelling cancer cells (Gu et al., 2010). It should be noted that the power and energy threshold for cancer-cell necrosis are significantly decreased by using the biofunctional gold nanorods. A 100% death rate of a cancer cell can be confirmed through necrosis or apoptosis processes when the laser power is higher than 3 mW, corresponding to an energy fluence of 0.53 mJ cm–2 (Gu et al., 2010). Conclusion The research work reviewed in this paper has demonstrated the applications of fibre-optical microendoscopy in imaging
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fields. 3D optical sectioning abilities through cells have been confirmed. Compared with OPEF microendoscopy, the advent of TPEF and SHG microendoscopy technique paves a way to improve endoscopic image contrast and image resolution and observe functional and morphological changes of microstructures in vivo. With compact microendoscopy, endoscopists can be provided with clear and robust visualization of biological samples and the detection of small and subtle dysplastic lesions to enhance the early detection, diagnosis and staging of gastric cancers. This, in turn, will facilitate the early detection of cancers at the stage of dysplasia and the enhanced cure rate of cancer patients. In this regard, two specific functionalities of fibre-optical microendoscopy, superresolution imaging (Gu et al., 2014a) and cellular manipulation or tweezing (Gu et al., 2014b), are highly desirable.
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