Bertrand ViellerobeMauna Kea Technologies, 9 Rue d’Enghien, 75010 Paris, France
http://www.maunakeatech.com/
Abstract: We report the design and implementation of spectroscopicand multicolor imaging capabilities into a fibered confocalfluorescencemicroscope (FCFM) already capable ofin vivo imaging. The real timeimaging device and the high resolution fiber probe make this system thefirst reported capable of performing multi color detection in the field ofFCFM. The advantages of the system will allowin vivo morphologicaland functional imaging. Preliminary experiments were carried out in tissuesamples to demonstrate the potential of the technique. The quality of theaxial sectioning achieved in the confocal fluorescence spectroscopy modeis demonstrated experimentally, and applications to multicolor imaging areshown.
References and links1. E. Laemmel, M. Genet, G. Le Goualher, A. Perchant, J.F. Le Gargasson and E. Vicaut, “Fibered Confocal Fluo-
rescence Microscopy (Cell-viZioTM ) Facilitates Extended Imaging in the Field of Microcirculation: A Compari-son with Intravital Microscopy,” J. Vasc. Res.41,400-411 (2004).
2. A.R. Rouse, A. Kano, J.A. Udovich, S.M. Kroto and A.F. Gmitro, “Design and demonstration of a miniaturecatheter for a confocal microendoscope,” Appl. Opt.43,5763-5771 (2004).
3. K. Carlson, M. Chidley, K.B. Sung, M. Descour, A. Gillenwater, M. Follen and R. Richards-Kortum, “In vivofiber-optic confocal reflectance microscope with an injection-molded plastic miniature objective lens,” Appl.Opt.44,1792-1797 (2005).
4. P.M. Lane, A. Dlugan, R. Richards-Kortum and C.E. MacCaulay, “Fiber-optic confocal microscopy using aspatial light modulator,” Opt. Lett.25,1780-1782 (2000).
5. T.D. Wang, M.J.Mandella, C.H. Contag and G.S. Kino, “Dual-axis confocal microscope for high-resolutioninvivo imaging,” Opt. Lett.28,414-416 (2003).
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8. H. Miyajima, K. Murakami and M. Katashiro, “MEMS Optical scanners for microscopes,” IEEE J. QuantumElectron.10,514-527 (2004).
9. K. Sokolov, J. Aaron, B. Hsu, D. Nida, A. Gillenwater, M. Follen, C. MacAulay, K. Adler-Storthz, B. Korgel,M. Descour, R. Pasqualini, W. Arap, W. Lam and R. Richards-Kortum, “Optical systems forin vivo molecularimaging of cancer.,” Technol. Cancer Res. Treat.2, 491-504 (2003).
10. A. Perchant, G. Le Goualher, M. Genet, B. Viellerobe and F.Berier, “An integrated fibered confocal microscopysystem forin vivo and in situ fluorescence imaging - applications to endoscopy in small animal imaging,” inProceedings of the IEEE International Symposium on Biomedical Imaging: From Nano to Macro, 2004.
11. A.R. Rouse and A.F. Gmitro, “Multispectral imaging with a confocal microendoscope,” Opt. Lett.25,1708-1710(2000).
12. G. Le Goualher, A. Perchant, M. Genet, C. Cave, B. Viellerobe, F. Berier, B. Abrat and N. Ayache, “Towards op-tical biopsies with an integrated fibered confocal fluorescence microscope,” Lecture Notes in Computer Science3217(II):761-768, Springer (Medical Image Computing and Computer Assisted Intervention), 2004.
13. P. Vincent, I. Charvet, L. Bourgeais, L. Stoppini, N. Leresche, J-P. Changeux, R. Lambert, P. Meda, D. Paupardin-Tritsch, ”Live imaging of neural structure and function by fibered fluorescence microscopy,” EMBO Reports7,11, 1154-1161, 2006.
14. L. Thiberville, S. Moreno-Swirc, T. Vercauteren, E. Peltier, C. Cave, G. Bourg-Heckly, ”In vivo imaging of thebronchial wall microstructure using fibered confocal fluorescence microscopy,” Am. J. Respir. Crit. Care Med.1, 175, 22-31, 2007.
1. Introduction
Laser scanning fiber-optic confocal microscopy provides the ability to perform non or mini-mally invasive subsurfacein vivo andin situ imaging at cellular level. Different technical ap-proaches, based on fiber bundles [1, 2, 3, 4, 10] or single fibers [5, 6, 7, 8] have been developedby various groups [9]. Confocal imaging can be reflectance-based [3, 7, 8, 9], using backscat-tered light from within the tissue, or fluorescence-based, using fluorescent light generated bychemical probes labeling specific tissue microstructures [1, 2, 4, 5, 6, 9, 10, 11]. The lattertechnique has emerged as the leading modality in the field of biological imaging, due to thehigh sensitivity and specificity of fluorescence probes. Theavailability of a growing numberof target-specific dyes and fluorescence proteins spanning the visible and near infrared spec-tral range has led to a new generation of benchtop microscopes featuring multiple excitationwavelengths and detection channels, capable of measuring simultaneously the distribution ofseveral fluorescent probes, such enabling multilabeling studies. Besides, in addition to imagingthe integrated fluorescence emission in the detection channel spectral windows, the knowledgeof the complete emission spectrum from a localized region brings additional capabilities; spec-tral information is relevant for detecting changes in the microenvironment associated to spec-tral shifts, helping unmix signals from overlapping fluorophores, assessing the contribution oftissular autofluorescence. . . Applicative results can be found in the ref [14] where direct iden-tification of the autofluorescence components of the bronchial mucosa was made possible bythe combination of imaging and spectroscopy in a single color mode. Then, the incorporationof multicolor imaging and spectroscopic analysis capabilities in fibered confocal fluorescencemicroscopy systems would considerably extend the potentialities of in vivo, in situ imaging. Toour knowledge, all current fibered systems are single wavelength excitation devices, only one ofthem exhibiting multispectral imaging and 2 excitation lasers [12], but with a slit-scanning so-lution and only one laser used at a time. The goal of this studywas to demonstrate the conceptof such a point scanning system, featuring two excitation wavelengths and two independentdetection channels, with an additional spectroscopic channel, capable of real time operation.In order to allow the use of a large number of fluorophores, twoprototypes were developed,each featuring two excitation wavelengths: 405 nm/488 nm and 488 nm/638 nm. The principleof operation is based on the fibered confocal fluorescence microscopy technology developedby Mauna Kea Technologies (Paris, France) which is already implemented in several commer-cialized instruments with single wavelength illuminationmode at 488 nm or 660 nm for small
#78339 - $15.00 USD Received 21 December 2006; revised 26 March 2007; accepted 26 March 2007
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InjectionSystem
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Fig. 1. Lay-out of the fibered confocal fluorescence spectroscopy and multicolor imagingsystem.
animal imaging [13] and for microendoscopy in the gastrointestinal tract and in the lungs [14].
2. System description
The scheme of the experimental set-up is shown in figure 1. A solid-state laser diode at 488 nm(Laser Sapphire from Coherent) and a second laser at 638 nm (Laser Cube from Coherent),are scanned by two mirrors on the proximal face of a fiber bundle. Horizontal line scanning isperformed using a 4 kHz oscillating mirrorM1 (GSI Lumonics) while a galvanometric mirrorM2 (GSI Lumonics) performs frame scanning. The resulting frame rate, limited by the numberof scanned lines, reaches 12 images/s for a 896× 640 pixels image size. The afocal systems(G = 1) ensure the conjugation of the scanning mirrors with the entrance pupil plane of theinjection system.
One of the main challenges was to obtain a high coupling efficiency of the laser beam withthe micron-sized optical fibers. For that purpose, a specificopto-mechanical connector wasdesigned to connect the fiber bundle to the laser scanning unit. The connector gives a highrepeatability in the focus position, as well as the optimal injection, along the whole imagefield. The fiber bundle is blocked in rotation and translationwithin a few microns. Finally,the optical beam was optimized to ensure that only one fiber isinjected at a time, each fibersequentially acting as a point source and a point detector. Thus, the system is working in aflying spot mode (point to point reconstructed image). The light injection module (LIM) isa custom made optical system composed of several doublets and triplets of lenses acting asa microscope objective. It was specifically designed to workon a large field of view (FOVtypically 600µm) with minimized aberrations (WFE< λ/5) , particularly in terms of spherical(< 2µm) and chromatic aberrations (< 9µm). Thus, this optical system is compatible with a
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large range of excitation wavelengths, from 400 nm to 700 nm.Its numerical apertureNA = 0.5and point spread function (PSF) were calculated to get the best injection rate into the fibers,around 70 %, the PSF fitting at best the diameter of the fundamental mode of the optical fibers(around 1.8µm) . Effectively, the best coupling efficiency between the excitation laser beamand the fibers, at proximal and distal ends, is obtained when the NA of the beam approximatelymatches the NA of the best-excited propagation mode in the fiber (0.22 for the TEM00 mode).Typically, the PSF is ranging from 1.2µm to 1.5µm depending on the wavelength.
The microprobes are composed of a connector (mentioned above), a fiber bundle and anoptical head. The fiber bundle (Fujikura, FIGH-30-650S) is made of 30,000 fiber cores. Thelateral resolution of the microprobe is limited by the fiber bundle inter-core distance of 3.3µmand the fiber core diameter of 1.9µm. At the distal end of the fiber bundle, the light beam isfocused in a plane located at a given depth in the biological tissue.
Two types of fiber bundle probes were developed:
Probe I:miniaturized (overall diameter 1.4mm), the fibers being directly in contact with the tissue. Thisprobe provides images of a 15µm thick tissue layer located immediately below the surface,with a lateral resolution of 3.5µm [13]. The field of view is adjustable from several hundredmicrons up to the useful entire diameter (600µm). The scanned surface is adjustable, with anominal size of 280µm×400µm. In general, the field of view is equal to the scanned surfacedivided by the magnification of the distal optics. Then, in the case of the Probe I, the field ofview is simply equal to the scanned surface.
Probe II :not miniaturized yet, based on an achromatic optical head (figure 2) composed of a×25 objec-tive O1 (Melles Griot, NA = 0.32, f ′ = 160mm), a field lens (Melles Griot, F ′
= 80mm) anda×100 objectiveO2 (Melles Griot, NA = 1.3, f ′ = 160mm, oil immersion, 0.17 cover slip).The resulting magnification (tissue to fiber) is 4. Fluorescence light collected by the distal op-tics propagates through the bundle using all available modes. The best NA for the optics whichcollects this light must thus be larger than the geometric NAof the fibers (0.42) on the prox-imal side. On the other hand, the geometric NA around 0.32 on the distal side clearly insuresa maximum coupling efficiency for the fluorescence coming back from the tissue. This probeprovides images with a field of view of 80µm×80µm of a 2µm thick tissue layer. The depthis adjustable from the surface of the tissue down to approximately 100µm. This z scanningcapability is realized by translating the specimen. This feature is not included inside the distalhead, meaning that adjustable depth of observation will notbe implementedin vivo.
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Fig. 2. Scheme of probe II.
A miniaturized version of the probe II is currently being assembled and tested (overall diame-ter = 4.2mm). The standard optical design of this probe is based on 4 lenses, one doublet and
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Table 1. Specifications of the probe II’s miniaturized version.
Fig. 3. Typical images obtainedin vivo with the miniaturized version of probe II. Healthycolonic crypts stained with fluorescein can be seen. FOV is 240µm. The working distanceis 30 µm.
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3 singlets. One lens is divergent in order to lower the chromatic aberrations. The high NA isobtained via one half-ball lens at the distal end of the micro-objective. The main specificationsof this fibered microprobe are listed in table 1. Image quality is diffraction limited within awide spectral range from 488 nm to 700 nm. Several working distances ranging from 20µmand 100µm can be set depending on the application. Earlyin vivo imaging results have beenobtained, demonstrating the confocal capability of this miniaturized probe, as shown in figure 3.The detection optical path is divided into two channels by a beamsplitter enabling simulta-neous imaging and spectroscopy: 80 % of the fluorescent signal being used for imaging and20 % for spectral analysis. The imaging signal is detected bytwo avalanche photodiodes(Hamamatsu), each one being associated with a pinhole, in two distinct spectral ranges(λ1 < λ < λ2 andλ2 < λ < 800nm) via a dichroic filterD3 (see table 2). The pinhole diameter,20µm, was calculated to reject the fluorescent light coming from adjacent fibers (encirclingthe excitation fiber). The detected optical signal consistsof a complex mixing of multipleradiations that can be splitted into two majors contributions: the fluorescence signal emittedby the tissue and the background signal generated by the interaction of the laser with theoptical fibers of the bundle (Rayleigh scattering, Raman scattering and Autofluorescence).Their levels are depending on the laser power injected into the fibers of the bundle. Typically,the incident power on the specimen is in the range of 3-6 mW depending on the fiber probe type.
Table 2. Dichroic and rejecting filters properties in the case of 488/635 nm architecture.
Type of filter Cut-off wavelength BandwidthD1 High-pass Dichroic 500 nm 250 nmD2 Dual-band Dichroic 500 nm 110 nm
A customized hardware is responsible for mirrors control and synchronous signal digitiza-tion. Images are rebuilt starting from this signal by an in-house built software described inthe MICCAI conference paper [12] which is a reference for medical image computing. Whilescanning the fiber bundle, the input light is modulated in amplitude by the injection rate ineach individual fiber of the bundle. This modulation createsa honeycomb effect on the rawimages. The task of the processing module is therefore to restore the images by removing thefiber bundle modulation, the scanning distortions and any unwanted signal such as the fiberautofluorescence.
The spectroscopic channel consists of an optical fiber (100µm core) coupled to a gratingspectrometer (USB2000, Ocean Optics) with an entrance slit width of 100µm achieving a spec-tral resolution better than 4 nm. The spectrometer is controlled through the Spectra Suite soft-ware (Ocean Optics). After the beamsplitter, the fluorescence signal is focused into the opticalfiber which serves as a confocal limiting aperture. On one hand, the axial resolution is obviouslymuch better for the imaging channel compared to the spectroscopic one, due to the smaller pin-hole size. On the other hand, the lateral resolution of the spectroscopic channel depending ondifferent parameters (pinhole size, scanned area, integration time of the spectrometer) cannotbe compared directly to the imaging channel. The fluorescence spectrum is recorded between500 nm and 750 nm. As excitation light generates autofluorescence and Raman scattering sig-nal within the fiber, a spectrum of this background was acquired prior to each experiment andsubtracted from the tissular emission spectrum. Then the spectrum is corrected for the spectral
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sensitivity of the spectrometer detector and optical setup. The spectrometer integration time isset to 80 ms, identical to the image scanning time, the spectra acquisition being synchronizedwith the frame trigger.
(a) (b) (c)
(d) (e)
Fig. 4. Images and spectra of a fixed human cervix sample stained with a 100µM solutionof DiA ( Invitrogen) (a,d) and with a 100µM solution of POPO-1 (Invitrogen) excited re-spectively at 488 nm and 405 nm (c,e). (b) is the fusion of these two images. (Probe II, FOV:80µm×100µm). Signal losses visible on the spectra are due to filters spectral response inthe case of 405/488nm architecture system.
For each detection channel, image and spectrum are simultaneously acquired and displayed onthe computer monitor in real-time. Then images are superimposed to get a multi-colored fusionimage (figure 4). False colors were selected to provide the best contrast.
3. Results
Both channels were validated separately and measurements of the axial and lateral resolutions atboth wavelengths were performed using 2µm-diameter fluorescent beads (Estapor, FXC 200).A drop of solution of those fluorescent beads was deposited ona plate, diluted in ethanol andevaporated. The maximum of emission fluorescence intensitywas measured by bidirectionalmicrometric translation. Results are given in table 3 and figure 5.
Table 3. Lateral and axial resolution (LR, AR) measured at 488 and 638 nm, with probes Iand II, for the spectroscopic and imaging channels (micrometers).
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Micrometric translation
Fig. 5. Axial intensity profiles of a fluorescent bead showing the resolution of the spectro-scopic channel for Probe I and Probe II at 488 nm. Axial resolution isdetermined by theFWHM of the curves.
The spectral interval between the two excitation wavelengths has been chosen to be wideenough to prevent the respective fluorophores’ absorption and emission spectra from overlap-ping and the corresponding imaging channels from perturbating each other, the illumination be-ing simultaneous or sequential. The quality of axial and lateral achromatism was demonstratedby a negligible shift of the images allowing easy superimposition. Images and confocal spectrafrom various samples of fixed human tissue (oesophagus, cervix, colon, thyroid), stained withdifferent fluorophores are shown in figures 4, 6 and 7.
The results obtained in the figures 4, 6 and 7 clearly indicatethe cellular imaging capabilityof the system, giving the possibility to a direct comparisonwith standard histology. For exam-ple, images from figure 4 show cell nuclei (in blue) and cellular/nuclear membranes (in red).The regular nuclei sizes and shapes seem to be compatible with an healthy tissue, histologicalanalysis giving the same conclusion. Furthermore, images aand b from figure 6 exhibit spe-cific cellular architecture, squamous epithelium for the first one (cervix tissue) and glandularepithelium for the second one (thyroid tissue), in agreement with standard histological patterns.In the same way, parakeratosis pathology, characterized byhigh nuclei density with small sizesand elongated shapes, is illustrated in the images c, d and e from figure 6, histology giving thesame diagnosis.
The high signal to noise ratio (SNR> 30dB) obtainedex vivo clearly opens newin vivo ap-plications which requires spectral analysis such as Forster Resonance Energy Transfer (FRET).
The spectral analysis will be more relevant in the future by adding Region Of Interest illumi-nation capabilities. In this case, spectral informations as changes in the microenvironment, dyescolocalization, contribution of tissular autofluorescence will definitely be helpful. This presentwork presents only early results and proof of concept of spectral acquisition capability within amulticolor system. Recentin vivo results associating spectral information within a single colorsystem were demonstrated by Thiberville et Al [14].
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(c) (d) (e)
(a) (b)
Fig. 6. Multilabeled images from fixed squamous and glandular human tissues with probeII. Cervix (a) and thyroid (b) samples are stained with 100µM solutions of DiA (red)and To-Pro-1 (Invitrogen) (blue), respectively excited at 488 nm and 638 nm. Multilabeledimages (c,d,e) from fixed human cervix samples exhibiting different parakeratosis grades,confirmed by the histological analysis (FOV: 80µm×100µm).
(a) (b)
Fig. 7. Multilabeled images with probe I from fixed human cervix samples stained with DiA(blue) excited at 488 nm and To-Pro-1 (red) excited at 638 nm (FOV: 430µm×300µm).The density of cell nuclei (in red) indicates a high-grade parakeratosis. The lower resolutionof the Probe I cannot give a detailed of view of the cellular membranes (inblue). Thesquamous epithelium pattern is clearly visible.
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4. Conclusion
This study demonstrates the feasibility of a point scanningFCFM system with spectroscopicand muticolor imaging capabilities. Preliminary experiments clearly demonstrate the ability ofthe system to perform real-time imaging. Miniprobes, such as probe I and miniaturized versionof probe II, can be used directly or inserted in the working channel of an endoscope, allowingin vivo, in situ morphological and functional imaging. This should be of particular interest forsmall animal imaging, where the wide range of available dyesallows a large number of combi-nations for multiple fluorescence staining. Due to toxicityissues, the number of dyes approvedfor human use is obviously more limited; however fluorescentprobes such as fluorescein, cre-syl violet or indocyanine green are of common use in clinicalpractice andin vivo single colormode microendoscopy is a fast growing technique. The development of new fluorescent con-trast agents is a rapidly progressing field which should alsobenefit to molecular imaging inhuman beings. In conclusion, the development of such a FCFM system, featuring two excita-tion wavelengths with spectroscopic analysis capability,paves the way to multilabeling studiesin live animals and clinical trials as well.
Acknowledgments
We thank A. Perchant from Mauna Kea Technologies for image processing. Human tissueswere provided through a collaboration with Dr E. Peltier. The support of OSEO-ANVAR(n◦A0405120Q) is gratefully acknowledged.
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