- Email: [email protected]
Endoscope-compatible confocal microscope using a gradient index-lens system
15 February 2001
Optics Communications 188 (2001) 267±273
www.elsevier.com/locate/optcom
Endoscope-compatible confocal microscope using a gradient index-lens system J. Knittel a,*, L. Schnieder a, G. Buess a, B. Messerschmidt b, T. Possner b a
Department of Minimally Invasive Surgery, University of Tubingen, Waldhornlestrasse 22, 72072 Tubingen, Germany b GRINTECH GmbH, Schillerstrasse 1, 07745 Jena, Germany Received 2 June 2000; received in revised form 20 October 2000; accepted 5 December 2000
Abstract We present an endoscope-compatible confocal laser scanning microscope (CLSM) for medical imaging. A gradient index (GRIN)-lens system projects a magni®ed image on the distal end of a ®ber-optic imaging bundle, which transfers the image to a CLSM situated at the proximal end. As the maximum diameter of the distal optics is 1 mm, application through the accessory channel of a conventional endoscope is possible. A lateral resolution of 3.1 lm and an axial resolution of 16.6 lm is achieved. Confocal images of colonic tissue stained with a ¯uorescent marker are shown. Ó 2001 Published by Elsevier Science B.V. Keywords: Optical biopsy; In vivo histology; Endoscopy; Confocal microscopy; Minimally invasive; Cancer
Confocal laser scanning microscopy is a well established technique to image thin sections within thick tissue with high resolution non-invasively [1]. In principle a point source is used to illuminate a point on the sample and the light emitted from this point is imaged through a pinhole placed at the proper position onto a detector. Using this con®guration light coming from out-of-focus points is mostly rejected because it does not produce a point focus at the pinhole. By scanning the illuminated point over the sample a thin slice within a thick tissue sample can be imaged with high resolution and contrast. This is known as optical sectioning
*
Corresponding author. Fax: +49-7071-295569. E-mail addresses: [email protected], knittelj@ thmulti.com (J. Knittel).
and makes CLSM attractive for imaging intact biological samples. In the medical ®eld, CLSM has the potential to gather histo-pathologic information non-invasively and in real time [1±3]. In the ®eld of endoscopy, a miniaturized high-resolution imaging system could be especially helpful for a number of medical problems, such as the determination of tumor margins during a minimally invasive operation, the screening of large tissue sections, and, most importantly, the guiding of conventional biopsy. For the pathologic inspection and diagnosis of tissue, the position and the shape of the nuclei within the cells are important criteria. Therefore a resolution in the range of 1±3 lm is necessary. This can be achieved by CLSM, using ¯uorescent markers to enhance the contrast in the images by staining the tissue [4].
0030-4018/01/$ - see front matter Ó 2001 Published by Elsevier Science B.V. PII: S 0 0 3 0 - 4 0 1 8 ( 0 0 ) 0 1 1 6 4 - 0
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J. Knittel et al. / Optics Communications 188 (2001) 267±273
Recently, signi®cant eorts have been made to develop miniaturized confocal imaging systems [5±7]. In order to be useful in the clinical setting, the overall diameter of the system has to be below 2.5 mm. Thus, the microscope can be passed through the accessory channels of standard endoscopes used in the gastrointestinal tract. To the best of our knowledge, the development of a CLSM of such size has only been published by Dickensheets and Kino [7]. Compared to their microscope, the miniaturized CLSM system we report on in the following has the advantage that no active parts have to be introduced into the human body, and that the gradient index (GRIN)optics used in our system covers a broader wavelength range than the diractive optics used in DickensheetsÕs setup. A wavelength coverage of about 50 nm is essential for ¯uorescence imaging. Fig. 1 shows a sketch of the experimental setup. The system is similar to the one presented by Gmitro and Avid [8], except that we use dierent microscope objectives. A conventional CLSM (Zeiss LSM 410) is used to illuminate the proximal end of a ®ber-optic imaging bundle (FOIB), which is placed exactly at the position where a sample would normally be placed. The FOIB serves to relay images between the proximal and the distal ends of the ®ber. The spatial integrity of the image is maintained by the FOIB. By scanning the laser beam of the CLSM, the individual ®bers of the bundle are illuminated one
by one and a light spot raster scans the sample. Simultaneously, light emitted from the sample is detected by the CLSM, provided that it is coupled into the ®ber that presently illuminates the sample. Due to the small diameter of the individual ®bers, each ®ber acts like a pinhole aperture. This produces the optical section capability of the microscope. The CLSM is operated in the confocal mode to suppress stray-light emitted from neighboring ®bers of the FOIB. A multi-immersion objective (Zeiss 16 0:45) is used to couple the light into the FOIB (Sumitono, IGN08/30, Japan). Glycerin with a refractive index of 1.47 is used as an index matching ¯uid to reduce re¯ections at both end faces of the FOIB. The GRIN-lens system is mounted onto a 3D microscope stage. The GRIN-lens system is composed of two gradient index rod lenses with dierent focal lengths and numerical apertures (NAs) (see Fig. 2) and a diameter of 1.0 mm. The objective lens was produced by silver/sodium ion exchange in glass [9], yielding a focal length of 0.93 mm and a NA of 0.5. The NA of the lens results from the index gradient ranging from 1.547 at the margin to 1.626 at the center of the lens. Refractive index pro®les were measured by a modi®ed refracted-near-®eld method [10]. The second imaging GRIN rod lens was fabricated by lithium/sodium ion exchange in glass [11], resulting in a focal length of 2.41 mm and a NA of 0.2. Both lenses have plane surfaces
Fig. 1. Schematic of the experimental setup: it consists of a CLSM (left part) and a FOIB equipped with a miniaturized objective with a magni®cation of 2.6.
J. Knittel et al. / Optics Communications 188 (2001) 267±273
Fig. 2. Ray tracing design of the GRIN-lens system (OL ± objective lens; IL ± imaging lens). The working distance of the objective lens is 74 lm in water. The system has a diameter of 1 mm and a total length of 7.8 mm.
and were glued together by a UV-curing adhesive. The objective lens has a length of 2.3 mm, the imaging lens a length of 5.5 mm yielding a total length of 7.8 mm. For endoscopic application, the overall length of the in¯exible part of the microscope system has to be below 15 mm. Otherwise, the system cannot be introduced through the working channel of an endoscope, as it is curved at the beginning and at the end of the endoscope. The system is designed to realize a magni®ed image of the object plane in its exit surface plane. The object plane is located in a working distance of 74 lm to the input surface plane of the objective lens. This value is valid for an immersion medium with the refractive index of 1.33 (water). According to the telescope principle used, the magni®cation of 2.6 results from the ratio of the focal lengths of both lenses. The optical design of the GRIN-system minimizes the spherical aberration and the ®eld curvature. The calculated root mean square (RMS) wavefront aberration of objective and collimator lens is 0.095 and 0.150 waves respectively. As the aberrations have opposite sign, the whole system has a wavefront aberration (RMS) of 0.06 waves in the center of the ®eld. Nearly diraction-limited performance in 90% of the lens aperture was proven experimentally by Shearing interferometry yielding an RMS wavefront error of 0.09 waves. The GRIN-lens system produces a calculated spot size with FWHM (full-width at half maximum) of 1.3 lm in the center of the object ®eld and of 1.4 lm (sagittal plane) and 1.9 lm (tangential plane) at the border of the object ®eld, which has a diameter of 280 lm. Adjacent spots sampled by successively illuminating individual ®bers have a
269
distance of 1.1 lm on the sample. Therefore, neighboring sample points overlap. The theoretical spot sizes and RMS wavefront aberrations were calculated with a ray tracing program (ZEMAX, Focus Software, Tucson, USA) using the Huygens point spread function algorithm and by placing a point source in the image plane. The refractive index n r at the radial distance r from the center of the gradient index material was modeled with the following formula: n r
p n0 n2 r 2 n4 r 4
The coecients n0 , n2 , n4 were determined by a modi®ed refractive near-®eld method (objective lens: 2.64, 1.17, 0.241, collimating lens: 2.34, 0.173, 0.0257) [10]. The lateral resolution of the entire system depends on the FOIB, the GRIN-lens system and the scanning system at the proximal side used for illumination and detection. In the following we neglect the in¯uence of the proximal scanning system. This was veri®ed experimentally by using dierent objective lenses 10 0:35, 16 0:45, 20 0:5 for input coupling at the proximal side. No signi®cant change of resolution was measured. Obviously, the FWHM of the intensity distribution of the scanning laser spot has to be smaller than half the distance between individual ®ber elements to prevent coupling of energy into multiple ®bers, which would result in a reduced lateral resolution. The FOIB limits the lateral resolution due to the limited number of image points, irregular ®ber packing and coupling of the illumination in neighboring ®bers (cross-talk). Sabharwal et al. [6] described these eects with an eective PSF (point spread function). The eective PSF of the ®beroptic bundle used in this study has an FWHM of 2.8 lm [6]. This corresponds to an FWHM of 1.1 lm at the sample due to the magni®cation of the GRIN-lens system. The convolution of the PSF of the FOIB and the GRIN-lens system respectively yields a PSF with a FWHM of approximately 2.8 lm for the whole system taking into account that the light passes two times through the system and assuming a Gaussian shape for the PSF of the ®ber bundle and the GRIN-lens system respectively.
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In this case the FWHM of the PSF of the entire system is given by the following equation: FWHM system
r 2 2 FWHM lens2 2 FWHM fiber2 M
where M is the magni®cation of the lens system. It can be concluded that reducing the aberrations of our GRIN-lens system is the most eective way to improve the resolution. Further improvement can be achieved by increasing the magni®cation of the lens system. However, this would reduce the ®eld of view. To visualize the lateral resolution of the microscope, we used ¯uorescent beads (T-7283, Molecular Probes, Eugene, USA) with a diameter of 4.2 lm. The beads were excited at a wavelength of 633 nm. Fluorescence was detected using a longpass ®lter with a cuto wavelength of 635 nm (LP 635, Zeiss, Germany). The ¯uorescence image of three beads is shown in Fig. 3. The diameter of each bead is represented by about three to four ®bers of the bundle. The nuclei of cells have a
Fig. 3. Three ¯uorescent beads with a diameter of 4.2 lm imaged with the microscope. For comparison a circle with a diameter of 4.2 lm was inserted. The ®gure illustrates the sampling of the object with individual ®bers. The ®eld size is 60 60 lm2 .
typical diameter of about 5 lm. Therefore, our microscope should be able to resolve nuclei. The pixel-related appearance is caused by the FOIB and not by aberrations of the optics. The ®ber bundle consists of 30,000 single multi-mode ®bers with a diameter of 2 lm each. The distance between the ®bers is about 3 lm. As there is a ®nite distance between the ®bers, part of the image does not contain any useful information. To suppress the pixilate image appearance, either image processing techniques, i.e. Gaussian low pass ®lters, could be applied, or the proximal end of the ®ber may be sampled with a laser beam that does not resolve individual ®bers [6]. In Fig. 3 no ¯uorescence intensity is detected in the regions between the individual ®bers. This signi®es that there is no disturbing energy transfer through that FOIB when the illuminating light spot is situated midway between two individual ®bers. The lateral resolution was measured from an image of the edge of a coverslip immersed in a solution of dextran (D-7136, Molecular Probes Eugene, USA). The ¯uorescence of the dye was excited at 488 nm. Due to the pixilated appearance of the image, the edge width determined from a single line ¯uctuates strongly. Therefore, we averaged along the direction of the edge to obtain an average edge function. By averaging 55 lm along the edge of the coverslip, a 10±90% edge resolution of 3.3 lm was obtained (see Fig. 4a). For a Gaussian spot the lateral resolution de®ned as its FWHM can be shown to be 0.925 times the 10± 90% edge width. This yields a lateral resolution of 3.1 lm. This FWHM is within 10% of the FWHM predicted by our model. Nevertheless, it is far away from the theoretical value of 0.43 lm for an aberration free objective with a NA of 0.5 [12]. The axial resolution of our miniaturized CLSM was determined by the standard method of moving a plane mirror through the focus of the microscope, while detecting the re¯ected light intensity. The curve is shown in Fig. 4b. The axial response intensity distribution is typical for spherical aberration [13]. The FWHM is 16.6 lm. The theoretical resolution for an ideal CLSM with in®nitely small pinhole and an objective with a NA of 0.5 is 2.8 lm. The signi®cant deviation observed is not
J. Knittel et al. / Optics Communications 188 (2001) 267±273
Fig. 4. (a) The lateral resolution of 3.1 lm (FWHM) was calculated form the edge resolution of an image of a coverslip. The value was averaged over 55 lm along the edge to reduce pixel noise and (b) the axial resolution was measured by the standard method of moving a mirror through the focus of the microscope. At a wavelength of 633 nm the measured axial resolution (FWHM) was 16.6 lm.
caused by the FOIB, but by aberrations of the GRIN-lens system. This was demonstrated experimentally by replacing the GRIN-lens system by two conventional microscope objectives [14]. Furthermore, there is a mismatch between the NA of the FOIB of 0.35 and the NA of the collimating GRIN lens of 0.2. For ideal confocal operation, the NA of the FOIB should be smaller than the NA of the collimating GRIN lens [15]. For medical applications, the microscope has to be able to resolve cellular structures within intact tissue. Recently, we demonstrated that images
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from gastrointestinal tissue taken by confocal ¯uorescence microscopy yield clinically relevant information [4]. Thus this mode of operation was used for the following experiments. Fig. 5a shows an image of colonic tissue of a rat. The tissue was stained with the ¯uorescence marker SYTO 61 using a concentration of 10 lm (S-11343, Molecular Probes, Eugene, USA) for 2 h immediately after being taken from the animal. The tissue was rinsed with water to wash o remaining dye afterwards. The image was taken approximately 20 lm below the surface of the tissue. The exact depth is not known, as the tissue sample adheres to the GRIN lens and moves in an uncontrolled fashion when the axial position of the miniaturized microscope is changed. The image was taken in 2 s by sampling the proximal end of the ®ber bundle with 1024 1024 image points. A HeNe laser with a wavelength of 633 nm was used to illuminate the sample with a power 290 lW. The image is shown in the negative contrast mode: areas shown dark emitted the highest amount of ¯uorescent light. This improves the clarity and simpli®es the comparison of confocal and conventional histologic images. According to the ¯uorescent dye used, the nuclei were preferably stained and showed the most intense ¯uorescence. They are found to be aligned in a circle around the border of glandular openings within the intestinal mucus membrane. This is typical for normal, i.e. healthy, colonic tissue. The glandular openings are called crypts. For the purpose of comparison we imaged the tissue sample shown in Fig. 5a also with a conventional CLSM, using a 40 1:2 water immersion objective (Zeiss, Germany). The corresponding image is shown in Fig. 5b. According to the manufacturer a lateral resolution of 0.3 lm and an axial resolution of 1 lm can be achieved with this objective. In comparing the two images the eects of the better resolution are obvious. With the conventional CLSM the nuclei, which have a diameter of about 5 lm, are well separated. Using the miniaturized microscope they partially overlap. In contrast to what is shown in Fig. 5a the lateral resolution of our miniaturized microscope should be sucient to separate the nuclei. This was demonstrated by imaging ¯uorescent beads with a
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Fig. 5. (a) Colonic tissue of a rat imaged with the miniaturized CLSM approximately 20 lm below the surface. The picture was smoothed with a Gaussian low pass ®lter (FWHM 1.5 lm). The ®eld diameter is 280 lm. (b) Same sample imaged with a conventional CLSM using a 40 1:2 water immersion objective. The image was taken 20 lm below the tissue surface and has a size of 319 160 lm2 .
J. Knittel et al. / Optics Communications 188 (2001) 267±273
diameter of 4.2 lm, as shown in Fig. 3. However, due to the axial resolution of 16.6 lm of the miniaturized confocal microscope, the ¯uorescence signals of nuclei from dierent image planes overlap. Thus, the nuclei in Fig. 5a are not well separated. We are currently working to improve the optical system by reducing the spherical aberration [13]. We have demonstrated experimentally the concept of a miniaturized CLSM suitable for diagnostic endoscopy. The endoscopic head of the system contains no active mechanical or electrical parts. A resolution of 3.1 lm lateral (FWHM) and of 16.6 lm axial (FWHM) was achieved. The microscopic system has the potential to enable imaging with cellular resolution through the working channel of a ¯exible endoscope.
Acknowledgements This research was supported in part by the BMBF grant 13N6943. We thank B. B ultmann and U. Vogel for providing access to the confocal microscope (Zeiss LSM 410).
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References [1] M. Rajadhyaksha, R.R. Anderson, R.H. Webb, Appl. Opt. 38 (1999) 2105. [2] F. Koenig, S. Gonzalez, W.M. White, M. Lein, M. Rajadhyaksha, Urology 53 (1999) 853. [3] H. Inoue, T. Igari, T. Nishikage, K. Ami, T. Yoshida, T. Iwai, Endoscopy 32 (2000) 439. [4] J. Knittel, S.M. Kr ober, H. Raestrup, M.O. Schurr, L. Schnieder, G.F. Buess, in preparation. [5] R. Sch utz, K. D orschel, G.J. M uller, Proc. SPIE 3197 (1997) 223. [6] Y.S. Sabharwal, A.R. Rousse, L. Donaldson, M.F. Hopkins, A.F. Gmitro, Appl. Opt. 38 (1999) 7133. [7] D.L. Dickensheets, G.S. Kino, J. Microelectromech. Syst. 7 (1998) 38. [8] A.F. Gmitro, D. Avid, Opt. Lett. 18 (1993) 565. [9] B. Messerschmidt, U. Possner, A. Pfeil, T. Possner, Inorganic optical materials, Proc. SPIE 3424 (1998) 88. [10] R. Goering, M. Rothhardt, Application of the refractednear-®eld technique to multimode planar and channel waveguides in glass, J. Opt. Commun. 7 (1986) 82±85. [11] N. Haun, D.S. Kindred, D.T. Moore, Appl. Opt. 29 (1990) 4056. [12] T.R. Corle, G.S. Kino, Confocal scanning optical microscopy and related imaging systems, Academic Press, London, 1996, p. 152. [13] C.J.R. Sheppard, M. Gu, Appl. Opt. 30 (1991) 3563. [14] R. Juskaitis, T. Wilson, T.F. Watson, Scanning 19 (1997) 15. [15] T. Dabbs, M. Glass, Appl. Opt. 31 (1992) 3030.
Optics Communications 188 (2001) 267±273
www.elsevier.com/locate/optcom
Endoscope-compatible confocal microscope using a gradient index-lens system J. Knittel a,*, L. Schnieder a, G. Buess a, B. Messerschmidt b, T. Possner b a
Department of Minimally Invasive Surgery, University of Tubingen, Waldhornlestrasse 22, 72072 Tubingen, Germany b GRINTECH GmbH, Schillerstrasse 1, 07745 Jena, Germany Received 2 June 2000; received in revised form 20 October 2000; accepted 5 December 2000
Abstract We present an endoscope-compatible confocal laser scanning microscope (CLSM) for medical imaging. A gradient index (GRIN)-lens system projects a magni®ed image on the distal end of a ®ber-optic imaging bundle, which transfers the image to a CLSM situated at the proximal end. As the maximum diameter of the distal optics is 1 mm, application through the accessory channel of a conventional endoscope is possible. A lateral resolution of 3.1 lm and an axial resolution of 16.6 lm is achieved. Confocal images of colonic tissue stained with a ¯uorescent marker are shown. Ó 2001 Published by Elsevier Science B.V. Keywords: Optical biopsy; In vivo histology; Endoscopy; Confocal microscopy; Minimally invasive; Cancer
Confocal laser scanning microscopy is a well established technique to image thin sections within thick tissue with high resolution non-invasively [1]. In principle a point source is used to illuminate a point on the sample and the light emitted from this point is imaged through a pinhole placed at the proper position onto a detector. Using this con®guration light coming from out-of-focus points is mostly rejected because it does not produce a point focus at the pinhole. By scanning the illuminated point over the sample a thin slice within a thick tissue sample can be imaged with high resolution and contrast. This is known as optical sectioning
*
Corresponding author. Fax: +49-7071-295569. E-mail addresses: [email protected], knittelj@ thmulti.com (J. Knittel).
and makes CLSM attractive for imaging intact biological samples. In the medical ®eld, CLSM has the potential to gather histo-pathologic information non-invasively and in real time [1±3]. In the ®eld of endoscopy, a miniaturized high-resolution imaging system could be especially helpful for a number of medical problems, such as the determination of tumor margins during a minimally invasive operation, the screening of large tissue sections, and, most importantly, the guiding of conventional biopsy. For the pathologic inspection and diagnosis of tissue, the position and the shape of the nuclei within the cells are important criteria. Therefore a resolution in the range of 1±3 lm is necessary. This can be achieved by CLSM, using ¯uorescent markers to enhance the contrast in the images by staining the tissue [4].
0030-4018/01/$ - see front matter Ó 2001 Published by Elsevier Science B.V. PII: S 0 0 3 0 - 4 0 1 8 ( 0 0 ) 0 1 1 6 4 - 0
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J. Knittel et al. / Optics Communications 188 (2001) 267±273
Recently, signi®cant eorts have been made to develop miniaturized confocal imaging systems [5±7]. In order to be useful in the clinical setting, the overall diameter of the system has to be below 2.5 mm. Thus, the microscope can be passed through the accessory channels of standard endoscopes used in the gastrointestinal tract. To the best of our knowledge, the development of a CLSM of such size has only been published by Dickensheets and Kino [7]. Compared to their microscope, the miniaturized CLSM system we report on in the following has the advantage that no active parts have to be introduced into the human body, and that the gradient index (GRIN)optics used in our system covers a broader wavelength range than the diractive optics used in DickensheetsÕs setup. A wavelength coverage of about 50 nm is essential for ¯uorescence imaging. Fig. 1 shows a sketch of the experimental setup. The system is similar to the one presented by Gmitro and Avid [8], except that we use dierent microscope objectives. A conventional CLSM (Zeiss LSM 410) is used to illuminate the proximal end of a ®ber-optic imaging bundle (FOIB), which is placed exactly at the position where a sample would normally be placed. The FOIB serves to relay images between the proximal and the distal ends of the ®ber. The spatial integrity of the image is maintained by the FOIB. By scanning the laser beam of the CLSM, the individual ®bers of the bundle are illuminated one
by one and a light spot raster scans the sample. Simultaneously, light emitted from the sample is detected by the CLSM, provided that it is coupled into the ®ber that presently illuminates the sample. Due to the small diameter of the individual ®bers, each ®ber acts like a pinhole aperture. This produces the optical section capability of the microscope. The CLSM is operated in the confocal mode to suppress stray-light emitted from neighboring ®bers of the FOIB. A multi-immersion objective (Zeiss 16 0:45) is used to couple the light into the FOIB (Sumitono, IGN08/30, Japan). Glycerin with a refractive index of 1.47 is used as an index matching ¯uid to reduce re¯ections at both end faces of the FOIB. The GRIN-lens system is mounted onto a 3D microscope stage. The GRIN-lens system is composed of two gradient index rod lenses with dierent focal lengths and numerical apertures (NAs) (see Fig. 2) and a diameter of 1.0 mm. The objective lens was produced by silver/sodium ion exchange in glass [9], yielding a focal length of 0.93 mm and a NA of 0.5. The NA of the lens results from the index gradient ranging from 1.547 at the margin to 1.626 at the center of the lens. Refractive index pro®les were measured by a modi®ed refracted-near-®eld method [10]. The second imaging GRIN rod lens was fabricated by lithium/sodium ion exchange in glass [11], resulting in a focal length of 2.41 mm and a NA of 0.2. Both lenses have plane surfaces
Fig. 1. Schematic of the experimental setup: it consists of a CLSM (left part) and a FOIB equipped with a miniaturized objective with a magni®cation of 2.6.
J. Knittel et al. / Optics Communications 188 (2001) 267±273
Fig. 2. Ray tracing design of the GRIN-lens system (OL ± objective lens; IL ± imaging lens). The working distance of the objective lens is 74 lm in water. The system has a diameter of 1 mm and a total length of 7.8 mm.
and were glued together by a UV-curing adhesive. The objective lens has a length of 2.3 mm, the imaging lens a length of 5.5 mm yielding a total length of 7.8 mm. For endoscopic application, the overall length of the in¯exible part of the microscope system has to be below 15 mm. Otherwise, the system cannot be introduced through the working channel of an endoscope, as it is curved at the beginning and at the end of the endoscope. The system is designed to realize a magni®ed image of the object plane in its exit surface plane. The object plane is located in a working distance of 74 lm to the input surface plane of the objective lens. This value is valid for an immersion medium with the refractive index of 1.33 (water). According to the telescope principle used, the magni®cation of 2.6 results from the ratio of the focal lengths of both lenses. The optical design of the GRIN-system minimizes the spherical aberration and the ®eld curvature. The calculated root mean square (RMS) wavefront aberration of objective and collimator lens is 0.095 and 0.150 waves respectively. As the aberrations have opposite sign, the whole system has a wavefront aberration (RMS) of 0.06 waves in the center of the ®eld. Nearly diraction-limited performance in 90% of the lens aperture was proven experimentally by Shearing interferometry yielding an RMS wavefront error of 0.09 waves. The GRIN-lens system produces a calculated spot size with FWHM (full-width at half maximum) of 1.3 lm in the center of the object ®eld and of 1.4 lm (sagittal plane) and 1.9 lm (tangential plane) at the border of the object ®eld, which has a diameter of 280 lm. Adjacent spots sampled by successively illuminating individual ®bers have a
269
distance of 1.1 lm on the sample. Therefore, neighboring sample points overlap. The theoretical spot sizes and RMS wavefront aberrations were calculated with a ray tracing program (ZEMAX, Focus Software, Tucson, USA) using the Huygens point spread function algorithm and by placing a point source in the image plane. The refractive index n r at the radial distance r from the center of the gradient index material was modeled with the following formula: n r
p n0 n2 r 2 n4 r 4
The coecients n0 , n2 , n4 were determined by a modi®ed refractive near-®eld method (objective lens: 2.64, 1.17, 0.241, collimating lens: 2.34, 0.173, 0.0257) [10]. The lateral resolution of the entire system depends on the FOIB, the GRIN-lens system and the scanning system at the proximal side used for illumination and detection. In the following we neglect the in¯uence of the proximal scanning system. This was veri®ed experimentally by using dierent objective lenses 10 0:35, 16 0:45, 20 0:5 for input coupling at the proximal side. No signi®cant change of resolution was measured. Obviously, the FWHM of the intensity distribution of the scanning laser spot has to be smaller than half the distance between individual ®ber elements to prevent coupling of energy into multiple ®bers, which would result in a reduced lateral resolution. The FOIB limits the lateral resolution due to the limited number of image points, irregular ®ber packing and coupling of the illumination in neighboring ®bers (cross-talk). Sabharwal et al. [6] described these eects with an eective PSF (point spread function). The eective PSF of the ®beroptic bundle used in this study has an FWHM of 2.8 lm [6]. This corresponds to an FWHM of 1.1 lm at the sample due to the magni®cation of the GRIN-lens system. The convolution of the PSF of the FOIB and the GRIN-lens system respectively yields a PSF with a FWHM of approximately 2.8 lm for the whole system taking into account that the light passes two times through the system and assuming a Gaussian shape for the PSF of the ®ber bundle and the GRIN-lens system respectively.
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J. Knittel et al. / Optics Communications 188 (2001) 267±273
In this case the FWHM of the PSF of the entire system is given by the following equation: FWHM system
r 2 2 FWHM lens2 2 FWHM fiber2 M
where M is the magni®cation of the lens system. It can be concluded that reducing the aberrations of our GRIN-lens system is the most eective way to improve the resolution. Further improvement can be achieved by increasing the magni®cation of the lens system. However, this would reduce the ®eld of view. To visualize the lateral resolution of the microscope, we used ¯uorescent beads (T-7283, Molecular Probes, Eugene, USA) with a diameter of 4.2 lm. The beads were excited at a wavelength of 633 nm. Fluorescence was detected using a longpass ®lter with a cuto wavelength of 635 nm (LP 635, Zeiss, Germany). The ¯uorescence image of three beads is shown in Fig. 3. The diameter of each bead is represented by about three to four ®bers of the bundle. The nuclei of cells have a
Fig. 3. Three ¯uorescent beads with a diameter of 4.2 lm imaged with the microscope. For comparison a circle with a diameter of 4.2 lm was inserted. The ®gure illustrates the sampling of the object with individual ®bers. The ®eld size is 60 60 lm2 .
typical diameter of about 5 lm. Therefore, our microscope should be able to resolve nuclei. The pixel-related appearance is caused by the FOIB and not by aberrations of the optics. The ®ber bundle consists of 30,000 single multi-mode ®bers with a diameter of 2 lm each. The distance between the ®bers is about 3 lm. As there is a ®nite distance between the ®bers, part of the image does not contain any useful information. To suppress the pixilate image appearance, either image processing techniques, i.e. Gaussian low pass ®lters, could be applied, or the proximal end of the ®ber may be sampled with a laser beam that does not resolve individual ®bers [6]. In Fig. 3 no ¯uorescence intensity is detected in the regions between the individual ®bers. This signi®es that there is no disturbing energy transfer through that FOIB when the illuminating light spot is situated midway between two individual ®bers. The lateral resolution was measured from an image of the edge of a coverslip immersed in a solution of dextran (D-7136, Molecular Probes Eugene, USA). The ¯uorescence of the dye was excited at 488 nm. Due to the pixilated appearance of the image, the edge width determined from a single line ¯uctuates strongly. Therefore, we averaged along the direction of the edge to obtain an average edge function. By averaging 55 lm along the edge of the coverslip, a 10±90% edge resolution of 3.3 lm was obtained (see Fig. 4a). For a Gaussian spot the lateral resolution de®ned as its FWHM can be shown to be 0.925 times the 10± 90% edge width. This yields a lateral resolution of 3.1 lm. This FWHM is within 10% of the FWHM predicted by our model. Nevertheless, it is far away from the theoretical value of 0.43 lm for an aberration free objective with a NA of 0.5 [12]. The axial resolution of our miniaturized CLSM was determined by the standard method of moving a plane mirror through the focus of the microscope, while detecting the re¯ected light intensity. The curve is shown in Fig. 4b. The axial response intensity distribution is typical for spherical aberration [13]. The FWHM is 16.6 lm. The theoretical resolution for an ideal CLSM with in®nitely small pinhole and an objective with a NA of 0.5 is 2.8 lm. The signi®cant deviation observed is not
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Fig. 4. (a) The lateral resolution of 3.1 lm (FWHM) was calculated form the edge resolution of an image of a coverslip. The value was averaged over 55 lm along the edge to reduce pixel noise and (b) the axial resolution was measured by the standard method of moving a mirror through the focus of the microscope. At a wavelength of 633 nm the measured axial resolution (FWHM) was 16.6 lm.
caused by the FOIB, but by aberrations of the GRIN-lens system. This was demonstrated experimentally by replacing the GRIN-lens system by two conventional microscope objectives [14]. Furthermore, there is a mismatch between the NA of the FOIB of 0.35 and the NA of the collimating GRIN lens of 0.2. For ideal confocal operation, the NA of the FOIB should be smaller than the NA of the collimating GRIN lens [15]. For medical applications, the microscope has to be able to resolve cellular structures within intact tissue. Recently, we demonstrated that images
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from gastrointestinal tissue taken by confocal ¯uorescence microscopy yield clinically relevant information [4]. Thus this mode of operation was used for the following experiments. Fig. 5a shows an image of colonic tissue of a rat. The tissue was stained with the ¯uorescence marker SYTO 61 using a concentration of 10 lm (S-11343, Molecular Probes, Eugene, USA) for 2 h immediately after being taken from the animal. The tissue was rinsed with water to wash o remaining dye afterwards. The image was taken approximately 20 lm below the surface of the tissue. The exact depth is not known, as the tissue sample adheres to the GRIN lens and moves in an uncontrolled fashion when the axial position of the miniaturized microscope is changed. The image was taken in 2 s by sampling the proximal end of the ®ber bundle with 1024 1024 image points. A HeNe laser with a wavelength of 633 nm was used to illuminate the sample with a power 290 lW. The image is shown in the negative contrast mode: areas shown dark emitted the highest amount of ¯uorescent light. This improves the clarity and simpli®es the comparison of confocal and conventional histologic images. According to the ¯uorescent dye used, the nuclei were preferably stained and showed the most intense ¯uorescence. They are found to be aligned in a circle around the border of glandular openings within the intestinal mucus membrane. This is typical for normal, i.e. healthy, colonic tissue. The glandular openings are called crypts. For the purpose of comparison we imaged the tissue sample shown in Fig. 5a also with a conventional CLSM, using a 40 1:2 water immersion objective (Zeiss, Germany). The corresponding image is shown in Fig. 5b. According to the manufacturer a lateral resolution of 0.3 lm and an axial resolution of 1 lm can be achieved with this objective. In comparing the two images the eects of the better resolution are obvious. With the conventional CLSM the nuclei, which have a diameter of about 5 lm, are well separated. Using the miniaturized microscope they partially overlap. In contrast to what is shown in Fig. 5a the lateral resolution of our miniaturized microscope should be sucient to separate the nuclei. This was demonstrated by imaging ¯uorescent beads with a
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Fig. 5. (a) Colonic tissue of a rat imaged with the miniaturized CLSM approximately 20 lm below the surface. The picture was smoothed with a Gaussian low pass ®lter (FWHM 1.5 lm). The ®eld diameter is 280 lm. (b) Same sample imaged with a conventional CLSM using a 40 1:2 water immersion objective. The image was taken 20 lm below the tissue surface and has a size of 319 160 lm2 .
J. Knittel et al. / Optics Communications 188 (2001) 267±273
diameter of 4.2 lm, as shown in Fig. 3. However, due to the axial resolution of 16.6 lm of the miniaturized confocal microscope, the ¯uorescence signals of nuclei from dierent image planes overlap. Thus, the nuclei in Fig. 5a are not well separated. We are currently working to improve the optical system by reducing the spherical aberration [13]. We have demonstrated experimentally the concept of a miniaturized CLSM suitable for diagnostic endoscopy. The endoscopic head of the system contains no active mechanical or electrical parts. A resolution of 3.1 lm lateral (FWHM) and of 16.6 lm axial (FWHM) was achieved. The microscopic system has the potential to enable imaging with cellular resolution through the working channel of a ¯exible endoscope.
Acknowledgements This research was supported in part by the BMBF grant 13N6943. We thank B. B ultmann and U. Vogel for providing access to the confocal microscope (Zeiss LSM 410).
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