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Journal of Microscopy, Vol. 207, Pt 2 August 2002, pp. 137 – 145
Received 7 November 2001; accepted 11 June 2002
Near real time in vivo fibre optic confocal microscopy:
sub-cellular structure resolved
Blackwell Science, Ltd
K. B. SUNG*, C. LIANG†, M. DESCOUR†, T. COLLIER*,
M. FOLLEN‡, A. MALPICA‡ & R. RICHARDS-KORTUM*
*Department of Biomedical Engineering, University of Texas at Austin, Austin, TX 78712, U.S.A.
†University of Arizona, Optical Sciences Center, 1630 East University Boulevard, Tucson, AZ 85721,
U.S.A.
‡The University of Texas M. D. Anderson Cancer Center, 1515 Holcombe Boulevard, Houston,
TX 77030, U.S.A.
Key words. Confocal microscopy, fibre optic, in vivo, laser, reflectance, scanning.
Summary
We have built a fibre optic confocal reflectance microscope
capable of imaging biological tissue in near real time. The
measured lateral resolution is 3 µm and axial resolution is
6 µm. Images of epithelial cells, excised tissue biopsies, and
the human lip in vivo have been obtained at 15 frames s−1. Both
cell morphology and tissue architecture can be appreciated
from images obtained with this microscope. This device has
the potential to enable reflected light confocal imaging of
internal organs for in situ detection of pathology.
1. Introduction
Optical technologies are being increasingly used to perform
real time assessment of tissue pathology in vivo. One particularly promising new technology is confocal microscopy, which
samples small volumes of tissue, producing images with micrometre resolution at depths up to several hundred micrometres
within tissue. A number of groups have used non-fibre optic
confocal microscopes to obtain reflected light images of accessible tissues. In skin, confocal imaging can provide detailed
morphological images of epithelial cell morphology and tissue
architecture throughout the entire epithelial thickness, as well
as detailed imaging of subepithelial blood flow in capillaries
(Betrand & Corcuff, 1994; Rajadhyaksha et al., 1995; Corcuff
et al., 1996; Masters et al., 1997; Rajadhyaksha et al., 1999b).
Recently, confocal imaging has been used to image neoplastic
skin lesions in vivo (Busam et al., 2001; Langley et al., 2001).
However, the application of confocal imaging to in vivo
detection of pathology in epithelial tissues other than the skin
has been limited by the difficulty in access to these organ sites.
Correspondence: Dr R. Richards-Kortum. Tel.: + 1 512 471 2104; fax: + 1
512 475 8854; e-mail: [email protected]
© 2002 The Royal Microscopical Society
A number of groups have attempted to develop flexible
endoscopes to record confocal images in vivo based on fibre
optics. Gmitro & Aziz (1993) first proposed the use of a fibre
optic bundle to transfer the scanned image plane and raster
scanning the proximal end of the bundle to produce en face
images of the sample at the focal plane. Each fibre within the
bundle serves as the point source as well as the detection
pinhole for confocal imaging. The Gmitro group has recently
developed a slit-scan confocal fluorescence microendoscope
with a miniature objective and a hydraulic focusing mechanism (Sabharwal et al., 1999). Fluorescence images of
biological samples have been obtained from cultured human
prostate cells and tissue sections, as well as mouse peritoneum
in vivo. All samples were stained with vital fluorescent dyes to
provide image contrast. In vivo fluorescence imaging has two
important limitations. First, fluorescent stains must typically
be used to yield sufficient signal for imaging. These dyes must
be non-toxic to the tissue and able to penetrate to deeper layers
within the tissue. Second, the excitation wavelength is usually
in the UV or blue range, which has a short penetration depth
within tissue because of scattering.
Fibre optic confocal microscopes imaging reflected light
from tissue could be used with longer wavelength sources
to obtain deeper penetration for in vivo imaging without the
need for fluorescent stains. Fluctuations in refractive indices
of tissue provide contrast that allows cellular and subcellular
structures to be detected (Dunn et al., 1996). Refractive indices
have been reported to be n = 1.4–1.45 for the nuclei and
n = 1.35–1.37 for the cytoplasm (Schmitt & Kumar, 1996).
Therefore, for biological imaging the system needs to be able
to detect reflected light from ∆n < 0.1 and resolve the size of
cell nuclei. One fibre-based confocal system utilized a singlemode optical fibre to deliver illumination light to the sample to
be imaged and to collect the backscattered light from the
illuminated focal volume (Dickensheets & Kino, 1998). En face
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Fig. 1. Schematic of the fibre optic confocal
reflectance microscope. The operation of the laser is
controlled by the host computer. The detected signal
from the avalanche photodiode (APD) is digitized by
a frame grabber on the computer and integrated
with timing signals from the scanning system to
provide video output on the computer monitor.
images were obtained by microfabricated scanning mirrors
and focusing optics at the distal fibre end. Juskaitis et al.
(1997) developed a reflection confocal microscope using
a fibre optic bundle and a white light source. A tandem
scanning microscope was used to obtain real time images. A
primary limitation for fibre-based confocal reflectance microscopes is that the specular reflections generated at the end
faces of the fibre are many orders of magnitude stronger than
the tissue backscattering (Yang et al., 1999). Angle polishing
the fibre faces and index matching has been used to reduce
the unwanted reflections. However, no images of biological
samples have been reported because of the low light efficiency
and sensitivity of these devices.
Recently we have developed a fibre optic confocal reflectance microscope (FCRM) capable of imaging biological tissues
with subcellular resolution. In this paper we present the
design, performance and biological image results of the first
prototype of this imaging system, which uses a fibre bundle
coupled to a conventional microscope objective. The development of the focusing optics for this system has been described
in detail (Liang et al., 2001). The basic concept of the FCRM
is similar to that described by Gmitro. However, our system
detects reflected light produced without the need for exogenous stains, and operates at longer wavelength for deeper
penetration. Two galvanometer mirrors are used to generate
a raster scan of the illumination spot at the proximal end
of the fibre bundle. The problem of low efficiency in Juskaitis’
approach has been solved by using a monochromatic laser
source and focusing the illumination beam onto individual
fibres for higher illumination intensity. The scanning mirrors
also have much higher efficiency than the tandem scanning
method. The specular reflections from fibre end faces are
reduced by index matching liquid at both fibre ends. In vitro
and in vivo images of biological specimens including epithelial
cells, biopsies from human uterine cervix, and human lip have
been obtained with this prototype and are presented in this
paper. The goal of our research is to develop a flexible, fibre optic
confocal microscope to image pre-cancer in internal organs
such as the cervix and oral cavity. The curable precursor of
cancers, epithelial pre-cancer, is characterized by increased
nuclear size, increased nuclear to cytoplasmic ratio, hyperchromasia and pleomorphism, which can only be assessed
through invasive biopsy. Screening and detection could be
vastly improved by technologies that image subcellular structure in vivo without need for expensive and painful tissue
removal and examination.
2. Experimental
2.1. Fibre optic confocal reflectance microscope
The system diagram of the FCRM is shown in Fig. 1. A coherent
image guide from Sumitomo (IGN-15/30, Sumitomo, Electric
USA Inc.) is located between the focusing lens L1 and the
objective optics. This image guide has 30 000 fibres, an overall
outer diameter of 2.5 mm, a length of 5 m, and a nominal
numerical aperture (NA) of 0.3. The fibres inside this bundle
have an average core diameter of 4 µm and average centre-to-centre
spacing of 7 µm. The laser beam from the illumination optics is
focused onto the proximal end face of the fibre bundle such
that only one fibre is illuminated at one time. At the distal end
of the bundle, the illumination fibre is imaged to the sample by the
objective optics. The illumination fibre also serves as the detection
pinhole in a conventional confocal microscope ( Juskaitis &
Wilson, 1992). Backscattered light from the sample goes back
through the fibre bundle and emerges from the proximal fibre
end. A secondary pinhole, located in the detection path, is adjusted
to a position that is conjugate to the illumination /detection
fibre. Backscattered light from locations other than the illumination/detection fibre will be mostly rejected by the pinhole.
Therefore, the condition for confocal imaging is achieved.
2.2. Illumination optics
The light source is a continuous wave Nd:YAG laser (Optomech
Ltd, Torrance, CA, USA) that emits near-infrared light at 1064 nm.
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This particular laser is chosen for its short coherence length so
that the effect of interference is minimized. The laser beam passes
through a spatial filter, a collimation lens and a variable iris that
limits the beam diameter. A beam splitter partially transmits
the incident beam to the first scanning mirror and partially
reflects backscattered light toward the detection arm. The
scanning system consists of two scanning mirrors, two lenses
located between the mirrors, and the electronics to drive and
monitor the mirrors. In order to achieve a frame rate of 15
frames s−1 and an adequate number of pixels in the digitized
images, the line-scan device must operate at several kHz. One
scanner that meets our requirements is a resonant scanner
(SC-30 with driver PLD-XYG, Electro-Optical Products Corp.,
New York, USA) that oscillates sinusoidally at a pre-set resonant
frequency of 7.68 kHz. The second scanner is a magnetically
driven mirror (6800HP, Cambridge Technology, Inc., Cambridge,
MA, USA) operating at 15 Hz to provide a linear frame-scan. The
two lenses between the mirrors form a telescope configuration
with two functions: to expand the incident beam and to make the
exit pupil of the first scanning mirror coincide with the second
mirror so that the design of the following optics is simplified.
The first custom-made lens system, L1, is designed to: (1)
place a diffraction-limited spot at the image plane over the
active region of the fibre bundle; (2) be used with immersion
oil to reduce the specular reflection from the fibre surface; (3)
be telecentric so that the chief ray is parallel to the optical axis
and normal incidence is obtained over the radial position on
the bundle surface; (4) have a NA of 0.3 to match the NA of the
image guide. Given the NA of L1 and the wavelength of the
source, the diffraction-limited spot size is calculated to be
4.3 µm, which is close to the core diameter of the fibres. The
optical design is aided by the software package ZEMAX and
the details have been reported previously (Liang et al., 2001).
At distal end of the bundle, the other lens system, L2, relays
emerging light from the fibre to the back aperture of a 40×,
1.15 NA, water-immersion objective (OApo 40× w/340, Olympus,
Melville, NY, USA). The design goals of L2 are the same as
those of L1 except for a larger back aperture.
2.3. Detection optics and signal to background
Backscattered light from the object follows the same path as
illumination light in reverse direction through the objective
lens, L2, the fibre bundle, and L1. The returned light is
descanned by the two scanning mirrors and is deflected to the
detection path by the beam splitter. The lens in front of the
pinhole forms an image of the proximal end of fibre bundle at the
plane containing the pinhole. The pinhole diameter is selected
to be slightly larger than the size of the image of a fibre core.
Backscattered light coming from an out-of-focus spot in the
tissue will be distributed over a number of fibres within a
certain distance to the illumination fibre. Only light travelling
back through the illumination fibre will pass through the
pinhole, ensuring optical sectioning. The detector is a
© 2002 The Royal Microscopical Society, Journal of Microscopy, 207, 137–145
139
high-speed avalanche photodiode (APD) with a preamplifier
module from Hamamatsu (C5460, Hamamatsu, Bridgewater,
NJ, USA). This APD module has an APD gain of 10–300, and a
responsivity of 1 × 105 V W−1 at 1064 nm when the gain is 30.
One of the biggest challenges for the FCRM is the overwhelming specular reflection from the surfaces of the illumination fibre. This background light is intrinsic to the system
because the fibre end faces are conjugate to the detection
pinhole. Cleaving the fibre end faces at an oblique angle will
not work well because the fibres within the bundle will be at
different locations other than the desired focal plane of either
lens system L1 or L2. An alternative, pursued here, is to use
index-matching oil to reduce the background so that the
extremely faint signal from the tissue can be detected.
Complete elimination of the background cannot be achieved
because the fibres consist of two materials, the core and
cladding. Immersion oil with a refractive index halfway between
the core and cladding indices has been used at the proximal
(L1) end. The index difference between the oil and the fibres
is 0.015. At the distal (L2) end the index of immersion oil
matches the index of the fibre core so specular reflection from
the distal fibre surface can be neglected.
The signal-to-background ratio of the FCRM is predicted to
be 0.4 assuming that the signal is produced by ∆n = 0.1 interface and a total attenuation of e−2 in the tissue. The signal-tonoise ratio (S/N) is predicted to be 500, assuming a maximum
power of 0.15 W at the input of L1. These calculations take
into account the transmission of L1, the fibres, L2, and the
objective lens. The main sources of noise are shot noise of the
photodiode and thermal noise of the preamplifier electronics.
2.4. Image formation
The FCRM is a point-scanning device. In other words, at any
instant, only one (or none) small region in the tissue is illuminated and the backscattered light from that region is detected.
In order to form a 2D image, we need to scan the focus over a
plane that is perpendicular to the optical axis, record the
signal from each sampled region, and reconstruct the image
based on the collected data. In this case the signal is a timevarying voltage from the APD preamplifier, which corresponds to intensity of backscattering from the sample. We use
a frame grabber (MV-1000, MuTech, Billerica, MA, USA) to
digitize the analogue signal and convert it to VGA-compatible
format so that we can view the images on a computer
monitor. Because the fibres in the bundle are not regularly
packed, and the digitizer only works at a fixed sampling rate, it
is not possible to selectively save signal that is from the object
through the fibres and discard signal that is from the cladding or interfibre filling areas. Therefore, we oversample the
APD output signal at a very high frequency of 10.4 MHz so
that at least five points are sampled within each fibre.
Image formation is accomplished by synchronizing the
operations of the frame grabber and the scanning mirrors. The
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controller of the scanning mirrors provides two outputs that
correspond to the positions of the line-scan and frame-scan
mirrors. The zero crossings of these two position outputs are
detected and used to generate two TTL-compatible pulses that
are sent into the frame grabber to provide the H and V synchronization pulses, respectively. Each H sync pulse triggers
the frame grabber to start a new line and each V sync pulse
triggers a new frame. The frame grabber can be programmed
to have the proper delay time between the triggering signals
and the actual start of a line or a frame so that images corresponding to the raster scan are reconstructed. Images are
updated at a rate of 15 frame s−1 on the computer monitor and
no drift is observed between individual frames.
2.5. Spatial resolution
The lateral resolution of the FCRM cannot be described as
0.46λ/NA as in a conventional confocal microscope because
the image of the object is pixelated by the fibres. Following the
definition of resolution as the smallest separation between two
object points that can be visually resolved, the lateral resolution of the FCRM cannot be better than the separation
between two spots that are illuminated by two adjacent fibres.
The transverse magnification from the tissue to the distal fibre
end is 3.8, given that the NAs of the objective lens and L2 are
1.15 and 0.3, respectively. The distance between two adjacent
illumination spots in the tissue can be calculated as 7 µm/
3.8 = 1.8 µm. Because the diffraction-limited spot diameter in
the tissue is only 1.1 µm, the illumination spots do not overlap
in the tissue. Therefore, the lateral resolution of the FCRM
is limited by sampling rather than by diffraction. A lateral
resolution of 3 µm is estimated and is sufficient to resolve
the size of a cell nucleus.
The axial resolution of the FCRM can be calculated following
the method reported by Gu et al. (1991). In their model a single-mode fibre with the parameter A = (2πa0r0/λd )2 is used for
both illumination and detection. The field distribution of a
single-mode fibre with is approximated by a Gaussian profile
with a radius r0, at which the intensity is e−1 of the maximum;
a0 is the aperture radius and d is the focal length of lens system
L2. The axial resolution can be calculated from the axial optical coordinate u1/2, which is related to parameter A. The parameter A is calculated to be 6 using the specifications of the
system set-up, predicting a full width at half maximum
(FWHM) axial resolution of 2 µm.
2.6. Samples imaged
The performance of the FCRM was tested by imaging standard
samples including a mirror, a glass Ronchi grating and
polystyrene microspheres. Then the FCRM was used to image
various biological samples including rat breast cancer cells
in vitro, human epithelial cells scraped from oral mucosa, ex vivo
biopsies taken from human cervix, and the human lip in vivo.
Fig. 2. Images of a Ronchi grating oriented (a) horizontally and (b)
vertically. The grating has 500 cycles/inch, which corresponds to widths
of 25 µm for both the bright and the dark regions.
Biopsy specimens were excised from patients undergoing
colposcopic examination at the M.D. Anderson Cancer Center
in Houston, Texas. Informed consent was obtained from each
patient, and the study was reviewed and approved by the
Internal Review Boards at the University of Texas M.D. Anderson Cancer Center and the University of Texas at Austin. These
biopsies were put in tissue growth media (DMEM without
phenol red) for 6–8 h before being imaged at the University of
Texas at Austin. For all biological samples both still images in
bitmap format and videos in AVI format were saved. Background images were recorded along with regular images, and
the corresponding background images were subtracted from
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Fig. 3. Measured spatial resolutions of the system. (a) A line profile was plotted across an edge of the grating image. The edge response was measured as
1.8 µm from the distance between 10% and 90% intensity marks. The line profile was averaged over a line width of 100 µm. (b) The axial resolution was
measured by moving a mirror through focus; the FWHM value of the average intensity was 6 µm. The lines plotted between measured points are only to
suggest continuity and do not represent any kind of model of the measured data.
the regular images in order to reduce the effects of residual
specular reflection. In raw images, fibres at the left and right
edges appear to be extended horizontally because the line scan
pattern of the resonant scanner is sinusoidal instead of linear.
The image distortion was corrected by image processing and
the contrast and brightness of the resultant images were
adjusted for better presentation.
3. Results
Figure 2 shows two images of a glass Ronchi grating orientated
both horizontally and vertically. The target has widths of
25 µm for both the bright and dark regions. The field of view
can be found to be about 180 µm × 160 µm from these
images. The intensity is not uniform for all fibres over the
bundle, which is mainly because of variations in coupling
efficiency of fibres.
The lateral resolution was measured by imaging a Ronchi
grating and calculating the distance marked by 10% and 90%
intensity across the edge. A line profile from a grating image
is shown in Fig. 3(a). The line profile is averaged over a line
width of 100 µm to smooth the intensity variations resulting
from the pattern of fibres. The 10–90% distance is found to be
1.8 µm, which is the same as the calculated distance between
adjacent illumination spots in the object.
The axial resolution was measured by moving a planar
mirror in 2 µm steps through the focus of the system. The
average intensity versus the axial position is shown in Fig. 3(b)
and the FWHM is 6 µm. The predicted FWHM axial resolution
is 2 µm using Gaussian approximation for single-mode fibres.
The discrepancy between the measured and predicted values
is partly attributed to aberrations in the optics and multimode
propagation of light in the fibres. Higher order modes can also
propagate in the fibres because the fibre parameter V of the
fibres that we used is 3.5, given NA = 0.3 and core radius =
© 2002 The Royal Microscopical Society, Journal of Microscopy, 207, 137–145
Fig. 4. Image of 4.3 µm polystyrene microspheres immersed in water.
These spheres are comparable to cell nuclei in size. Scale bar is 20 µm.
2 µm. The second order modes have been shown to have larger
axial response than the fundamental mode ( Juskaitis &
Wilson, 1992).
The spatial resolution measured from planar objects (mirror
and grating) is sufficient to resolve cellular and subcellular
structures in epithelial tissue. In order to test the system
performance on small objects, polystyrene microspheres with
4.3 µm diameter were imaged (Fig. 4). These microspheres
provide a good sample for testing the system because the
nuclei of human epithelial cells are usually 5–15 µm in diameter. The image shows that the microspheres are well resolved
by the FCRM.
Epithelial cells in suspension were imaged by the FCRM.
Figure 5(a) shows an image of cultured MTC cells (rat breast
cancer cells) that were washed of growth media, suspended in
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phosphate buffered saline (PBS), placed on top of a thick layer
of gelatin and then covered by a coverslip. The gelatin layer
provides a substrate for the cell suspension and avoids any
stray background from the saline/gelatin interface because
the refractive index of gelatin is approximately matched to
saline. Figure 5(b) shows an image of human epithelial cells
taken from the oral mucosa. The cells were scraped from the
inside of the cheek of one of the authors (R.R.K.) and immediately smeared onto the gelatin substrate. PBS and 6% acetic
acid solution were added and a coverslip was placed. The
bright features in the images are identified as cell nuclei based
on the size and the contrast.
The optical sectioning ability of the FCRM was tested by
imaging ex vivo biopsy specimens. Immediately before imaging,
6% acetic acid solution was added to enhance the contrast
of the cell nuclei (Drezek et al., 2000). The image plane was
parallel to the tissue surface and the specimen was imaged at
different depths from the surface to about 150 µm below the
surface. The maximum depth of imaging is currently limited
by working distance of the objective lens. Figure 6(a)–(d)
show images of a colposcopically normal cervical biopsy with
the image plane located at about 40, 80, 100 and 150 µm
below the tissue surface, respectively. Images of an abnormal
cervical biopsy from the same patient are shown in Fig. 6(e)–
(h), with the image plane 40, 60, 90 and 120 µm beneath the
tissue surface, respectively. The cell nuclei are clearly resolved
and useful information such as nuclear density, nuclear area
Fig. 5. Confocal reflectance images of (a) cultured MTC cells (rat breast
epithelial cells) in PBS and (b) human epithelial cells taken from inside of
cheek and immersed in PBS and acetic acid. Scale bars are 20 µm.
Fig. 6. (a)–(d) Images of a colposcopically normal cervical biopsy with the image plane (a) 40, (b) 80, (c) 100 and (d) 150 µm below the tissue surface.
(e)–(h) Images of a colposcopically abnormal cervical biopsy from the same patient with image plane (e) 40, (f) 60, (g) 90 and (h) 120 µm below the
tissue surface. Images were acquired within 8 h after excision. Six per cent acetic acid was added to enhance the contrast. Scale bar is 20 µm.
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Fig. 8. In vivo images of human epithelial cells in the lip were acquired
with 6% acetic acid added to the lip before examination. The image plane
was close to the surface but the exact depth was not available. Scale bar is
20 µm.
Fig. 7. (a)–(c) Images of a tissue slice from a colposcopically normal
cervical biopsy. (d)–(f ) Images of a tissue slice from a colposcopically
abnormal cervical biopsy from the same patient. The thickness of the
slices is about 200 µm. The field of view is near the basement membrane
in (a) and (d), at the middle of epithelium in (b) and (e), and near the top of
epithelium in (c) and (f ). Scale bar is 20 µm.
and nuclear/cytoplasmic ratio can be extracted from the
images. In general, a decrease in signal level is observed when
the image plane is moved deeper into the tissue, which is
expected because light is attenuated (mostly scattered) in
tissue. Variations in signal level and contrast, however, exist
from sample to sample. The source of these variations may be
intrinsic to tissues or dependent on the process of sample
preparation, or a combination of the two.
Tissue sections 200 µm thick were cut from cervical biopsies in the orientation perpendicular to the surface of the
epithelium. Six per cent acetic acid solution was added prior to
imaging. Figure 7(a)–(c) show images of a tissue section from
a normal biopsy and Fig. 7(d)–(f ) show images of an abnormal
tissue section from the same patient. For both tissue sections
the images are arranged such that the stroma region is on the
left and top of the epithelium is on the right. The basement
membrane of epithelium is shown clearly in Fig. 7(a) and (d).
© 2002 The Royal Microscopical Society, Journal of Microscopy, 207, 137–145
The nuclear density in the images taken from the abnormal
biopsy is clearly increased as compared to the images taken
from the normal biopsy. The area of individual nuclei in the
images of the abnormal tissue is also greater than that of the
normal tissue.
The FCRM was used to image human epithelial cells in the
lip in order to assess the feasibility of in vivo imaging by this
approach. Figure 8 shows an image of epithelial cells in the lip
of one of the authors (K.B.S.). The maximum output power
that entered the tissue was less than 40 mW, which was
comparable to that used by Rajadhyaksha et al. (1999a). Six per
cent acetic acid solution was added to enhance the contrast of
the cell nuclei. Images were easily blurred by motion of the
volunteer, so videos were saved into AVI files and individual
frames of still images were extracted from the video files. The
same procedure of background subtraction and contrast
enhancement were applied to the images afterwards. The cell
nuclei were clearly resolved in the images and showed images
with higher contrast than those of the ex vivo biopsies.
4. Discussion
The results shown here provide evidence that backscattered
light confocal imaging of biological tissues can be carried out
in near real time through flexible fibre optic bundles. The
spatial resolution, sensitivity and signal-to-background of the
FCRM are sufficient to permit imaging of subcellular structures in epithelial tissues. Images of cervical biopsy specimens
(Figs 6 and 7) show significant differences in nuclear density
and area between the colposcopically normal and abnormal
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biopsy. Morphological information such as nuclear to cytoplasmic
ratio can be extracted from confocal images (Collier et al.,
2000) and used to distinguish between pre-cancerous and
normal tissue (Collier et al., 2002). The results suggest that
such a device would be useful for in situ detection of pathology.
The outer diameter of this prototype is 3.8 cm; this is small
enough so that it could be used to image the cervix during
colposcopy (Utzinger et al., 2001). We are currently developing
a second prototype which incorporates a miniature objective
lens, along with an appropriate mechanism to hold the objective lens stably against the tissue and provide a controllable
axial scanning. This device should enable imaging of the oral
cavity and other internal organ sites.
The lateral resolution of the FCRM is limited by the pixelated
nature of the fibre bundle. According to the Nyquist theorem,
for fully reconstructing the image we need to sample the image
at a spatial frequency at least two times that corresponding to
the diffraction limit of the objective optics (Webb & Dorey,
1995). Here the sampling frequency posed by the fibres is
much lower than the required Nyquist frequency. The
incidence of undersampling may cause aliasing in the images.
However, for our application we are interested in the area,
density and distribution of cell nuclei. In epithelial tissue the
cell nuclei are 5–15 µm in diameter and about 10–20 µm
apart. The sampling pitch (1.8 µm) of the FCRM is sufficient to
locate cell nuclei and resolve the areaof a nucleus to certain
accuracy.
The use of a fibre bundle with multimode fibres in a laser
scanning confocal reflectance microscope may result in a high
degree of speckle ( Juskaitis et al., 1997). Our approach is to
use a laser source with a short coherence length to reduce
interference between reflections from various surfaces in the
light path and backscattered signal from tissue. For epithelial
tissue which the FCRM is designed to image, there is no solid,
highly reflective planar surface to generate considerable level
of speckle. This is confirmed by the images of biological
samples in which no prominent speckle is observed.
A resonant galvanometer is selected to achieve the highspeed line scan because of its optical and mechanical simplicity, fairly low cost, high reflection efficiency, adjustable scan
angles and negligible variations from one scan line to the next.
The major disadvantage of using a resonant galvanometer is the
sinusoidal nature of the scan. In the FCRM, only the central
part of the scan in one direction is used as the active line,
resulting in a low duty cycle of about 40%. Non-linearity is
noticeable at the left and right edges of raw images. The
image distortion was corrected by image processing for all
images presented in the Results section. A software subroutine
was made to correct the distortion in the image of the
grating (Fig. 2a). The same subroutine was applied to all
images of biological samples. Another solution is to acquire
the pixels at intervals that are not uniform in time but
correspond to equal intervals in image space along the lines
(Montagu, 1991).
Index matching at the proximal and distal surfaces of the
fibre bundle sufficiently reduces the specular reflections
produced there to enable acquisition of high contrast images.
Image contrast is degraded somewhat by the residual specular
reflection from the proximal fibre surface. This background
can be further reduced by polishing the fibre bundle end face
at a small angle instead of perpendicular to the optical axis,
provided that the angle is small enough so that the focal spot
of L1 remains nearly diffraction-limited over the active areas
of the bundle. A constant background level can be subtracted
from the video signal by blocking the dc component of the
APD output. However, noise generated by the electronic
components needs to be carefully controlled because the electronics work in the radio frequency range. In practice, moving
the pinhole axially off its optimal position can also reduce the
background. The pinhole will no longer be conjugate to the
illumination/detection fibre so part of the specular reflection
from the fibre surface will be blocked. The trade-off is, of
course, worse spatial resolution.
The images presented here have a pixelated appearance
because of the fibre bundle. Gaps between fibres within features
can be filled by image processing techniques such as low
pass filtering, median filtering or morphological operations
(Thiran & Macq, 1996). Interpolation can also be utilized to
produce regular, smooth-looking images. Because of the rapid
acquisition speed, these are difficult to implement in real time.
The pixelation is perceived most strongly in the still images
presented here. The raw videos displayed on the computer
screen are perceived to be much less pixelated. Persistence of
vision provides a time-averaging effect on the image sequence
so the effects of noise and pixelation are reduced. Small movements of the object across the field of view greatly facilitate
perception and recognition of small or dim features over a
constant background in the videos. Therefore, the current system
set-up is preferable for a real time diagnostic imaging device.
5. Conclusions
We have built a FCRM that has sufficient resolution and sensitivity to image biological samples in cellular and subcellular
level at half video rate. No fluorescent stain is needed because
backscattered light from the tissue is detected. Images of
epithelial cells, excised tissue biopsies and the human lip
in vivo have been obtained. The evident differences in nuclear
morphology between normal and abnormal biopsy images
suggest that the FCRM has diagnostic value for detection of
cervical pre-cancers. We believe that FCRM will be highly
useful for the recognition and monitoring of pathology in
epithelial tissues in vivo.
Acknowledgements
This project is supported by an NIH grant (R01 CA82880).
The authors thank Ina Pavlova for sample preparation.
© 2002 The Royal Microscopical Society, Journal of Microscopy, 207, 137 – 145
JMI1049.fm Page 145 Wednesday, July 24, 2002 10:22 PM
I N V I VO F I B R E O P T I C C O N F O CA L M I C RO S C O P Y
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