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Hybrid Small Animal Imaging System
Combining Magnetic Resonance Imaging
With Fluorescence Tomography Using
Single Photon Avalanche Diode Detectors
Florian Stuker, Christof Baltes, Katerina Dikaiou, Divya Vats, Lucio Carrara, Edoardo Charbon, Member, IEEE,
Jorge Ripoll, and Markus Rudin*
Abstract—The high sensitivity of fluorescence imaging enables
the detection of molecular processes in living organisms. However,
diffuse light propagation in tissue prevents accurate recovery
of tomographic information on fluorophore distribution for
structures embedded deeper than 0.5 mm. Combining optical
with magnetic resonance imaging (MRI) provides an accurate
anatomical reference for fluorescence imaging data and thereby
enables the correlation of molecular with high quality structural/functional information. We describe an integrated system
for small animal imaging incorporating a noncontact fluorescence
molecular tomography (FMT) system into an MRI detector.
By adopting a free laser beam design geometrical constraints
imposed by the use of optical fibers could be avoided allowing for
flexible fluorescence excitation schemes. Photon detection based
on a single-photon avalanche diode array enabled simultaneous
FMT/MRI measurements without interference between modalities. In vitro characterization revealed good spatial accuracy of
FMT data and accurate quantification of dye concentrations. Feasibility of FMT/MRI was demonstrated in vivo by simultaneous
assessment of protease activity and tumor morphology in murine
colon cancer xenografts.
Index Terms—Biomedical imaging, magnetic resonance imaging
(MRI), optical imaging, optical tomography.
I. INTRODUCTION
OMBINING two or more imaging modalities that provide complementary information on tissue morphology
and tissue-specific molecular processes is an attractive concept
for improving diagnostic specificity and patient care. For example, the combination of positron emission tomography (PET)
C
Manuscript received December 14, 2010; accepted January 30, 2011. Date
of publication February 10, 2011; date of current version June 02, 2011. This
work was supported in part by the Swiss National Foundation under Grant SNF
3100-112835 and Grant 310030-126029, and in part by the European Union
under Project FP7 FMT-XCT. Asterisk indicates corresponding author.
F. Stuker, C. Baltes, K. Dikaiou, and D. Vats are with the Institute
for Biomedical Engineering, University and ETH Zurich, 8093 Zurich,
Switzerland (e-mail: stuker@biomed.ee.ethz.ch; baltes@biomed.ee.ethz.ch;
dikaiou@biomed.ee.ethz.ch; vats@biomed.ee.ethz.ch)
L. Carrara and E. Charbon are with the Aqua Group, EPFL, 1015 Lausanne,
Switzerland (e-mail: ing.luciocarrara@gmail.com; edoardo.charbon@epfl.ch).
J. Ripoll is with the Institute of Electronic Structure and Laser—FORTH,
Vassilika Vouton, 71110 Heraklion, Greece (e-mail: jripoll@iesl.forth.gr).
*M. Rudin is with the Institute for Biomedical Engineering, University and
ETH Zurich, 8093 Zurich, Switzerland and also with the Institute of Pharmacology and Toxicology, University Zurich, 8057 Zurich, Switzerland (e-mail:
rudin@biomed.ee.ethz.ch).
Digital Object Identifier 10.1109/TMI.2011.2112669
and X-ray computed tomography (CT) has emerged as a sensitive tool in cancer diagnostics [1]. More recent developments
relate to hybrid PET/MRI systems as clinical and preclinical
imaging devices [2]. Combinations with MRI are attractive due
to the high quality structural information provided as a result of
high soft-tissue contrast. Dedicated MRI methods also yield accurate information on tissue physiology that might be intimately
linked to the underlying molecular processes. Moreover, MRI
does not involve ionizing radiation like X-ray CT.
Mapping of molecular processes in vivo demands for high
sensitivity of the imaging modality as provided e.g., by PET.
However, applications of PET may be limited by the short halflife of commonly used positron emitting isotopes, which requires on-site synthesis of PET tracers and correspondingly access to a cyclotron and radiochemistry facility. In contrast, fluorescence markers are in general stable and, when emitting in
the near-infrared range, suited for imaging structures deeply embedded in tissue. Therefore, fluorescence imaging has emerged
as an alternative to PET for small animal imaging yielding comparable molecular sensitivity. In view of the strengths of the two
modalities the combination of MRI and fluorescence imaging
constitutes an attractive concept for preclinical research [3].
In designing a hybrid FMT/MRI instrument, two strategies
can be pursued: 1) an integrated solution with the FMT setup
placed within the magnet or 2) sequential measurements on two
independent systems using e.g., an animal support compatible
with both modalities to minimize registration errors [4]. Advantages of the integrated approach are that both measurements
are inherently registered due to fixed instrument geometry and
can be performed simultaneously under identical physiological
conditions. Furthermore, the displacement of soft tissues due
to movement of the subject between different instruments is
avoided. However, several technological challenges are to be
met by combining FMT and MRI. In fluorescence imaging, photons emanating from the sample are commonly detected using
cooled charge-coupled detectors (CCD), which do not operate
at the high magnetic field strengths of animal MRI scanners.
Hence, a hybrid system based on CCDs requires the detectors
to be located outside of the magnet, and the use of optical fibers
for excitation and fluorescence detection [5], [6]. Fiber-based
systems are inherently limited by fixed geometry and relatively
small number of source-detector pairs [7], which compromises
spatial resolution. Moreover, space requirements of fiber bundles may not be compatible with dimensions of high-field small
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Fig. 1. Experimental setup. Schematic of the FMT/MRI instrumentation. The illumination module comprises the laser source, a resonator and the scan head for
deflecting the beam across the object. The sample platform comprises the SPAD detector, the focusing lens and the filter wheel used for selecting the wavelength
for detection. A semi-cylindrical RF coil is used in transceiver mode. The laser beam is deflected to the sample surface using a mirror.
animal MRI scanners [8]. Finally, signal loss at fiber couplings
may be detrimental when guiding low intensity light from the
sample surface via an optical fiber to the detector.
Fluorescence detectors that operate in high magnetic fields
would enable a fundamentally different solution. Detectors
could be positioned in close proximity to the fluorescent source
as an integral part of the radio-frequency (RF) antenna used for
the detection of MRI signals. Avalanche photodiode detectors
are not susceptible to high magnetic fields as demonstrated in
PET/MRI systems [9] and are often used in other biomedical
applications [10]. In this work we describe a hybrid FMT/MRI
system based on a noncontact FMT setup [11], [12] using
single-photon avalanche diodes (SPAD) designed to fit inside
the 120 mm gradient bore of a small animal MRI scanner
operating at a magnetic field strength of 9.4 T. Devoid of optic
fibers, the system offers maximal flexibility with regard to
excitation schemes and sample placement. The performance of
the hybrid FMT/MRI system was characterized using phantoms
with optical parameters mimicking those of biological tissue
that comprised one or several fluorescent sources. Aspects
investigated were the accuracy of the spatial in-plane and depth
information derived from FMT data as well as the sensitivity
and linearity of the optical detectors. Feasibility for in vivo
imaging was assessed by simultaneously studying protease activity and tumor morphology in a murine colon cancer model.
II. MATERIALS AND METHODS
A. Instrumentation
The experimental setup consisted of an illumination module
located outside and a sample platform inside the magnet [Fig. 1].
The illumination module consists of a fiber coupled 670nm continuous wave laser (B&W Tek, Newark, DE) for excitation. The
multimode fiber (core diameter 100 m,
) was connected to a numerical aperture matched collimation lens (Thorlabs, Munich, Germany). The pinhole (diameter 0.5 mm) behind
the lens was mapped with an anti-reflectance coated spherical
singlet lens (
mm, Melles Griot, Bensheim, Germany)
by a 2f image on the subject. To achieve a focal length of 2 m (required due to the dimensions of the MR magnet) the laser beam
was guided through an optical resonator mounted in-between
pinhole and focusing lens. The resonator consisted of two coated
economy front surface mirrors (Thorlabs, Munich, Germany)
placed 400 mm apart and was traveled five times. The light
was then guided to a scan head (Scanlab, Puchheim, Germany),
where it was deflected by two galvanometric driven mirrors to
allow scanning of sources on the sample surface. These mirrors
deflected the beam by 90 from the input direction. The beam
was then directed through an 1800-mm-long tempered epoxy
resin tube of 72 mm inner diameter, which could be separated
at the center for better handling. All components were fixed on
a movable home-built aluminum breadboard placed outside the
mT to ensure
bore of the magnet in a magnetic stray field
the proper function of the components.
At distal end of the tube the sample platform was fixed. A
coated front surface mirror (Edmund Optics, Karlsruhe, Germany) was placed at the back flange of the sample platform
[Fig. 2(a)] located in the isocenter of the MRI, to deflect the
beam by an angle of approximately 70 on the surface of the object, passing through the rectangular window of the MRI transceiver surface coil. A reflection geometry setup was chosen,
STUKER et al.: HYBRID SMALL ANIMAL IMAGING SYSTEM COMBINING MAGNETIC RESONANCE IMAGING WITH FLUORESCENCE TOMOGRAPHY
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white light illumination during the acquisition of reference surface images. All parts used for the setup were made from materials compatible with operation in high magnetic fields.
B. Detector Details
Fig. 2. Setup components. (a) Side view of the sample platform. (b) Bottom
view of the detector PCB. (c) Sketch of the SPAD architecture.
which allowed the use of source intensity profiles for image reconstruction [13]. Diffusive light patterns on the surface of the
animal were recorded for each illumination point. Light photons emanating from the surface were selected according to
their wavelength using filters placed in the filter wheel. High
quality bandpass filters (Semrock, Rochester, NY) with peak
wavelengths of 660 and 720 nm for measurements at the excitation and fluorescence wavelength have been used. Filters
were characterized by
% transmission over a range of
nm. A small 4-mm-diameter fixed focus lens (Edmund Optics,
Karlsruhe, Germany) was mounted to the printed circuit board
(PCB) in front of the detector array, yielding a field-of-view
(FOV) of 8 8 mm at a focal distance of 33 mm [Fig. 2(b)]. A
custom made 32 32 array of single-photon avalanche diodes
(SPAD) [14] bonded on the PCB was used as photon detector
[Fig. 2(c)]. A small LED stripe at the back flange was used for
The detector consisted of a SPAD array with 32 32 pixels
implemented in a 0.35 m CMOS technology [15]. Each pixel
of dimension
m comprised a SPAD, a quenching circuit, and a 1-bit counter, all implemented with ten NMOS and
two PMOS transistors, resulting in a total pitch of only 30 m
and a fill factor of 3.14% (area ratio). The SPAD is a pn-junction biased above a breakdown voltage (
V) in
order to operate in time-uncorrelated photon counting mode. A
cathode bias voltage VOP of 21 V was applied to operate in
Geiger mode yielding an excess bias voltage VE of 3.3 V. At
this excess value, a maximum photon detection probability of
35% (fill factor not included) was measured with a sensitivity
spectrum ranging from 350 to 850 nm. Calculating the overall
sensitivity (fill factor times photon detection probability) revealed that about one in one hundred photons is detected. The
image sensor array operates according to the rolling shutter principle. Each row is selected via the row decoder and read independently for all the columns during one clock cycle. The row
is subsequently reset for initiating a new integration period. A
clock frequency of 48 MHz was used due to firmware limitations, thus the minimum integration time in our experiment was
2.66 s. The total exposure time was set by the selected iterations to build a frame. Since each pixel produces single bit
digital information, an external FPGA chip-set is used to reconstruct a multi-bit image by incrementing 1024 32-bit counters sequentially row-by-row. The chip-set was not designed to
tolerate high magnetic fields and was thus placed outside the
bore of the magnet along with the circuit needed for communication with the computer. The connection between the chip-set
and PCB was achieved via conventional flat ribbon cables. The
SPAD detector array was tested inside the MRI system to assess the potential interference of a high static magnetic field of
9.4 T. The signal-to-noise ratio (SNR) of 45 dB is dominated
by Poisson noise, while the maximum dynamic range achieved
at 12 frames per second was 90 dB. The median of the noise
dark count rate (DCR) per pixel measured over the entire array
is 140 Hz at room temperature and can be reduced by cooling
of the chip to
C to DCR a median of 98 Hz. DCR was unaffected by the magnetic field.
C. MRI Details
All MR experiments were carried out on a Bruker BioSpec
94/30 (Bruker BioSpin MRI, Ettlingen, Germany) horizontal
small animal MR system operating at 400 MHz equipped with
a gradient system capable of generating a maximum strength of
400 mT/m with a minimum rise time of 80 s. A home-built
rectangular surface coil (20 24 mm) was used for RF signal
transmission and reception. The coil made of a flexible PCB
substrate with standard copper layer was designed for mouse
imaging and therefore curved to a cylindrical radius of 15 mm.
A window of 13 20 mm was cut out of this single loop coil to
allow for optical measurement (illumination and detection) in
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reflectance mode from the top. Three-dimensional MR imaging
was performed using a gradient echo sequence with the parammm ; matrix dieters: field of view FOV
mensions MTX
; voxel dimension VOX
m ; excitation pulse angle
; echo
time/repetition time TE/TR
ms; number of averages
NA
resulting in a scan time = 15 min 22 s. In addition
separate acquisitions with zero excitation pulse angle were performed for estimation of noise levels. In vivo experiments were
performed using a two-dimensional multislice MR sequence
with the following parameters: 14 slices of 700 m thickness;
FOV
mm MTX
; pixel dimension
PIX
m TE/TR
ms; NA
,
yielding a total acquisition time of 11 min 5 s.
D. Phantom Preparation
For the phantom experiments two tissue-mimicking siliconebased phantoms were prepared. Phantom (A) was used to characterize spatial resolution and concentration accuracy of the
optical system and for the combined measurements with the
MRI. It comprised holes for dye administration separated with
center-to-center distances of 2, 3, 4, and 5 mm in a plane parallel to the surface at a depth of 1 mm. Phantom (B) was used
for depth analysis and comprised holes parallel to the surface
at depths of 1, 2, and 3 mm with respect to the surface with an
in-plane separation of 5 mm. The holes could be filled with fluorescent dyes. Both phantoms were based on a room temperature vulcanizing (RTV) silicone (Wacker Silicone, Munich, Gercm
many). To simulate optical tissue properties (
and
cm ), TiO particles (Alfa Aesar, Karlsruhe, Germany) and carbon black powder (Alfa Aesar, Karlsruhe, Germany) were added as scattering and absorption agents, respectively. Resulting optical properties were measured using a near
infrared spectrometer (OxiplexTS, Champaign, IL) at 690 nm.
The phantoms were cuboid shaped with a thickness of 10 mm
and contained holes of 1.5 mm diameter for administration of
the fluorescent dye at chosen positions. The dye was filled into
small glass capillaries (outer diameter: 1.5 mm inner diameter:
0.6 mm) and inserted into the holes.
E. Subcutaneous Tumor Animal Model
Pathogen-free female BALB/C (CAnN.Cg-Foxn1 nu /crl)
nude mice (8–10 week old, weighing 20–25 mg) were obtained from Charles River Laboratory (Sulzfeld, Germany) and
housed in a controlled environment (23 C, 12 h/12 h light/dark
cycle) with unlimited access to water and chlorophyll free
food (Kliba Nafag, Kaiseraugst, Switzerland). All experiments
were performed in accordance to the Swiss Veterinary law
(License no. ZH 172–2008. C51-cells, a colon cancer-derived
secondary cell-line, were used to generate tumors.
cells
were injected subcutaneously on the left thigh flank of the nude
mice. The animals were monitored every second day for their
body weight and tumor size. Animals with tumors of a size
3 3 mm and larger were selected for imaging experiments.
F. Fluorescent Agents
The commercially available cathepsin-activatable near infrared (NIR) fluorescent imaging agent ProSense 680 (VisEn
IEEE TRANSACTIONS ON MEDICAL IMAGING, VOL. 30, NO. 6, JUNE 2011
Medical, Bedford, MA), a fluorophore conjugated graft polymer
of polyethylene glycol (PEG) and poly-lysine, was administered intravenously via the tail vein in all in vivo imaging
studies. A single dose of 13 nmol was given followed by a
second dose of 13 nmol, 24 h after the first injection. The
FMT/MRI measurements were done 24 h after the second dose
of ProSense 680.
The near-infrared fluorescent oxazine dye AOI987 used in
the phantom experiments was prepared according to published
procedures [16].
G. Fluorescence Data Acquisition
The silicone-based phantoms were used for characterizing the
performance of the FMT/MRI system. For each measurement,
one or two fluorescent samples consisting of an aqueous solution of AOI987 [16] at various concentrations were included at
defined depths and with variable spacing. Initially the sample
was placed on the imaging platform within the FOV of the optical system and under the RF surface coil of the MRI system.
For excitation, the laser beam was scanned across a grid with
7 7 illumination spots with dimensions and the exact location of the grid adjusted operator-interactively in order to match
the ROI on the sample surface. We have chosen an integration
time of 2–4 s per source position depending on the fluorescent
signal strength. This results in a total acquisition time between 5
and 9 min. For each source position an image was recorded resulting in a
array of images, where
and
are
the number of pixels in x and y and
the number of sources.
For calibration purposes the sources were measured on a white
paper placed in the focal plane of the camera in order to get the
center position of the sources and the pixel dimensions in centimeters. After that the optical acquisition was started and images were recorded at two wavelengths corresponding to the excitation (660 nm) and emission maximum (720 nm) of AOI987.
For in vivo measurements the illumination scheme and filter
settings were kept the same.
H. Fluorescence Data Reconstruction
For image reconstruction the normalized Born approximation was used, which has been shown to yield robust data by
considering the ratio of the measured emission and excitation
images [17]. The values of this ratio were used in the forward
model [18] accounting for the fluorescence signal on the sample
surface. The reconstruction was performed by inverting this
forward model using an algebraic reconstruction technique
(ART). This algorithm is an iterative method used in computed
tomography [19] and adapted for imaging diffusive media with
boundary restrictions [20]. An initial assumption for the fluorophore distribution is iteratively improved using a least square
minimization procedure to minimize the difference between
reconstructed and measured data. The reconstruction yielded
a three dimensional map of the fluorescent source distribution
within the subject. Simulations were performed using the same
forward model and reconstruction algorithm.
III. RESULTS
Quantitative analyses have been carried out for assessing the
crosstalk between modalities and the accuracy of in-plane and
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Fig. 3. Spatial resolution in a phantom. (a)–(c) Reconstructed fluorescence intensity (arbitrary units) from simulated data in xz-projection for 4 mm (a), 2 mm
(b) and 0.8 mm (c) center-to-center tube separation. The tubes are indicated as red circles. The insets show the line profiles at
mm. (d), (e) Reconstructed
fluorescence intensity (arbitrary units) from experimental data in xz-projection for 4 mm (d) and 2 mm (e) center-to-center distance. The inset show the line profiles
at
mm. (f) Plot of the extracted tube separations for simulated data (blue circles) and experimental data (green stars) versus the theoretical separations (red
dashed line).
depth reconstruction. Moreover the linearity between the fluorescence intensity and the local dye concentration was investigated. Finally simultaneous hybrid measurements were performed with the combined system in a phantom experiment and
in a first in vivo animal study.
A. Crosstalk
In hybrid systems potential crosstalk between the imaging
modalities is an issue, i.e., sensitivity might be compromised as
compared to the sensitivity of standalone devices. Also, interference might cause image artifacts or image distortions. We calculated the signal-to-noise ratio (SNR) of the FMT by dividing the
signal of a ROI in fluorescent raw image by the standard deviation of noise value extracted from a ROI in a background image,
which was recorded in the absence of fluorescence excitation,
and where only ambient light and detector noise contribute to
the signal. The SNR was evaluated using the previously deand during the
scribed phantom prior
MR image acquisition
and we did not
observe any degradation in
within error limits due to insertion of the optical setup into the MRI system. Similarly, the
SNR of the MR measurements was evaluated with and without
simultaneous optical measurement (
versus
), which did not reveal an
effect on the MRI sensitivity. For MRI SNR calculations the
mean signal intensity in a sample ROI located at a fixed location within the FOV was divided by the standard deviation of
the intensity in a ROI comprising only contributions from noise.
Moreover, we did not observed any artifacts introduced by the
optical detector in the MR images since the SPAD was positioned at a distance of 40 mm from the RF coil and outside the
FOV of the MRI data acquisition. Correspondingly, potential artifacts caused by the SPAD were not visible in the MR images.
All MR sequences can be run by the setup, the only limitation
is the inhomogeneous excitation by the surface coil. We therefore conclude that crosstalk between modalities in FMT/MRI
system is negligible.
B. Spatial Resolution
A critical parameter in imaging is spatial resolution. In MRI
voxel dimensions are typically of the order of 150–500 m
depending on the acquisition parameters, whereas resolution
in FMT is inferior, typically 1–2 mm isotropic. The FMT
spatial resolution in our hybrid system was evaluated using
phantom (A) comprising two identical fluorescent tubes with
variable spacing (2–5 mm). Simulations revealed that sources
separated by more than 0.8 mm could be clearly discriminated
[Fig. 3(a)–(c)] by analyzing the line profile in x direction. To
analyze the line profile the reconstructed volume was projected
to the xz-plane. The line profile is drawn through the peak value
of the embedded tubes. To determine the spatial separation a
gauss fit was applied to each peak in the line profiles and the
distance between the peak value of the two fitted Gaussian
profiles was taken as the separation distance. If two clearly
distinguishable peaks could be identified in the line profile the
two signals were considered as spatially separable. Experiments
were limited by the diameter and the wall thickness of the capillaries used. The experimental dispersion of the fluorescence
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intensity was slightly larger than predicted by the simulations
[Fig. 3(d)–(e)] and some artifacts were observed, which are
a result of imperfect reconstruction due to deficiencies like
the forward model (assuming a fluorescence point source embedded in an optically homogeneous medium), data collection
(detector noise) and SNR which contributes to the uncertainty
of the data lead therefore to these artifacts. Comparing the full
widths at half maximum of the FMT/MRI setup for simulated
and experimental data, we estimate the FMT in-plane spatial
resolution of the FMT/MRI setup to be 0.9 mm. A linear correlation has been found between reconstructed experimental data
slope
; Fig. 3(f)]
and true tube separation [
indicative of the accuracy of the in-plane spatial localization.
C. Depth Resolution
Quantitative analysis of dye concentrations depends on accurate reconstruction of the depth of the fluorescent source, which
was analyzed for a depth range of 1–3 mm using both simulations and experimental data. To analyze accuracy of depth information a line profile was drawn in z direction across the maximum intensity of the volume projected onto the xz-plane. The
distance between the phantom surface and the peak value of a
fitted Gaussian line was taken as a measure of depth. For this
experiment phantom (B) was used comprising one fluorescent
tube in the investigated depth. The dispersion of fluorescence intensity increases slightly with increasing depth [Fig. 4(a)–(d)],
as photon scattering scales with their path length through tissue.
The observed artifact in Fig. 4(c) is a result of imperfect reconstruction due to different deficiencies like the forward model,
data collection and SNR which contributes to the uncertainty
of the data as already discussed. Comparing reconstructed with
true depths demonstrated a good correlation up to depth values
of about 3 mm [Fig. 4(e)]. If depth of the fluorescent source
exceeds 3 mm, the actual depth is underestimated in the reconstructed image data sets. This can be attributed to reflection geometry setup, which emphasizes contributions from superficial
structures. Simulated data reveal that the limited field of view in
the current setup has only minor influences on the accuracy of
depth localization. By keeping the imaging parameters constant
and doubling the FOV the accuracy of the depth localization of
the FMT setup only slightly increases by 2%.
D. Quantification Analysis
A principal motivation for using tomographic imaging is the
ability to derive quantitative information on the local dye concentrations. Two aspects are of interest: 1) the minimal dye concentration that can be detected (detection limit), and 2) the linearity of the relationship fluorescence intensity versus dye concentration. Experiments were carried out using a tissue phantom
with a fluorescent source embedded at a depth of 1 mm. The correlation between AOI987 concentration and fluorescence intensity reconstructed revealed a linear relationship [
;
Fig. 4(f)] for a range from 0.36 to 25 picomoles, with a detection limit of approximately 0.3 picomoles of AOI987 dissolved
in a volume of 1.5 l.
Fig. 4. Depth resolution in a phantom. (a), (b) Reconstructed fluorescence intensity (arbitrary units) from simulated data in xz-projection for 2 mm (a) and
1 mm (b) below the surface. The insets show the depth line profiles. The tubes
are indicated as red circles. (c), (d) Corresponding reconstructed fluorescence
intensity (arbitrary units) from experimental data in xz-projection for 2 mm (c)
and 1 mm (d) depth. The inset show the depth line profiles. (e) Plot of the extracted tube depth for simulated data (blue circles) and experimental data (green
stars) versus theoretical depth (red dashed line). (f) Reconstructed mean fluorescence intensity versus AOI987 concentration in the tube of the phantom.
E. Hybrid Imaging
The feasibility of hybrid FMT/MRI was investigated using a
tissue phantom [Fig. 5(a)–(d)] and in vivo using tumor bearing
mice [Fig. 5(e)–(j)]. Due to the fixed geometry of the two detector systems a phantom measurement using fluorescent tubes
visible in both modalities enabled the definition of the affine
transformation to register the FMT data using the MRI images
as reference coordinate system. Voxel dimensions and the slice
positions for the FMT data were adjusted accordingly. The same
affine transformation could then be applied to in vivo data sets
as the FMT/MRI geometrical arrangement was left unchanged.
For these experiments the acquisition time for both modalities for phantom and in vivo measurements were on the same
time scale allowing to measure time dependent processes simultaneously.
A sketch of the silicone slab used for phantom experiments is
shown in Fig. 5(a) where one of four holes is filled with the fluorescent dye and the slice position is indicated. It was shown that the
overlay image of MRI [Fig. 5(b)] and FMT [Fig. 5(c)] sections revealed an accurate localization of the reconstructed fluorescent
tube from the FMT data [Fig. 5(d)]. For in vivo proof-of-principle of simultaneous FMT/MRI measurements a murine tumor
model was used. C51 cells (colon cancer-derived cell-line) were
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Fig. 5. Simultaneous MRI and optical data acquisition. Phantom: (a) Schematic drawing of the phantom (A) (gray) with one whole filled with a fluorescent dye
(green). The slice displayed in the FMT image is indicated in light blue). (b) MRI slice of the phantom. (c) Reconstructed slice (fluorescence intensity distribution)
from the optical data set. (d) Overlay of MR image and reconstructed fluorescence intensity distribution. The white horizontal line indicates the top surface of
the phantom. In vivo: (e) MRI slice of a tumor bearing mouse. (f) Corresponding reconstructed fluorescence intensity from the simultaneous measured optical
signal. (g) Overlay of the two slices from the two different modalities. The white horizontal line indicates the topmost tangent plane to the tumor surface. (h)–(j)
Detailed view of the tumor region.
subcutaneously implanted in the flank of a nude BALB/C mouse.
The majority of tumors are known to express high levels of proteases (protein degrading enzymes), which are involved e.g., in
the degradation of the extracellular matrix to enable tissue invasion and metastasis formation [21]. An activatable probe (ProSense 680) that generates fluorescent signals only upon cleavage
by proteases [22], [23] was used to assess the level of protease activity by FMT. A representative MR image [Fig. 5(e)] and the corresponding FMT section displaying fluorescence intensity reflectiveofproteaseactivity[Fig.5(f)]havebeenregistered [Fig.5(g)].
They show regions of high protease activity within the tumor
tissue. The edge of the slab used for the FMT reconstruction corresponded to the tumor tangential plane in the MR image oriented
parallel to the focal plane of the detector [indicated by the white
horizontal line in Fig. 5(e)–(g)]. An enlarged region of the registered FMT and MR image gives an insight into the fluorescence
dye distribution (protease activity) within the structural image
provided by MR [Fig. 5(h)–(j)]. Interestingly, protease activity
measured in FMT was not homogeneous across the tumor. Histological analysis confirmed heterogeneity of protease activity
throughout the tumor.
IV. DISCUSSION
Proper interpretation of molecular and/or cellular information
derived from molecular imaging studies requires the registration
to an anatomical reference that is commonly recorded using different imaging modalities. A straightforward solution is a sequential imaging strategy, which requires a sample support that is compatible with both modalities and that comprises fiducial markers
as landmarks for the registration process. An advantage of this
setup is that it provides optimal performance of each standalone
system without compromising image quality. Disadvantages are
that sequential measurements take inherently longer which puts
demands regarding the maintenance of physiologically stable
conditions throughout the measurement period. In addition, when
studying fast concurrent processes with different modalities, sequential sampling is not feasible. Moreover, soft-tissue may move
during the translocation between imaging modalities, which demands for nonrigid body registration procedures that are intrinsically difficult. Thus true hybrid systems are highly attractive for
biomedical research applications, as they provide simultaneous
multiplexed information in aninherentlyregisteredmannerdueto
concurrent measurement with both modalities in a rigid configuration. Combinations with MRI, which provides a high soft-tissue
contrast and has become a standard method both in preclinical and
clinical structural and functional imaging, are challenging due to
the hostile environment of a high magnetic field. Hybrid optical
imaging-MRI systems are typically fiber-based and used in combination with large bore clinical MRI with gradient systems of
50–60 cm inner diameter to account for the space requirements
by the source and detector fiber bundles. The use of a clinical
scanner for studies in small rodents is suboptimal: lower field
strength and weaker magnetic field gradients would compromise
the spatial resolution achievable. This might be sufficient if MRI
data are only used as reference indicating major anatomical structures; yet studies with MRI providing complementary information as independent imaging modality will suffer. Moreover, the
rigid fiber geometry limits the flexibility of the optical imaging
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setup: the number of source-detector pairs is limited, definitely
much smaller than for typical array detectors (with typically
k pixels) and the excitation schemes/points are predefined by the
position of excitation fibers. Fiber-based systems are typically
operated with the light guiding fibers in direct contact with the
sample. This renders the experimental procedure cumbersome
requiring manual adjustments of individual fibers. The non-contact system presented in this work allows overcoming these limitations. Excitation can be varied by adjusting the parameters of
the scanning device, while the fixed distance between the SPAD
array detector and the MRI surface coil, which is directly placed
over the animal’s body surface, ensures maintenance of optical
focus (apart from the surface curvature) irrespective of the body
region sampled. The focal plane was set according to a tangential
plane on the animal surface perpendicular to the objective lens.
In a hybrid system combining two complementary imaging
modalities, system performance of each modality should be not
or only minimally degraded with regard to their standalone performance, though some performance loss as compared to optimized standalone systems, which might use different detector
technologies, might be acceptable in view of the benefits of the
multimodality readouts. In a combined FMT/MRI system, the
MRI setup defines the design of the integrated FMT system: the
detector must be functional at high magnetic fields and has to
fit into the small bore of the gradient system (120 mm inner diameter in our case). The use of an MRI-compatible SPAD array
comprising 32 32 detectors led to some loss in performance as
compared to standalone FMTs using cooled CCD cameras with
regard to sensitivity, field-of-view and the number of source-detector pairs available. Nevertheless, the setup is still superior
to fiber-based rigid geometry systems with regard to experimental flexibility and amount of independent data generated for
reconstruction. Moreover, its performance is expected to improve with new generation SPAD arrays. The spatial resolution provided by the FMT/MRI setup was found to be comparable to that of microPET systems [24], [25]: the in-plane spatial resolution derived from phantom experiments was at least
2 mm, whereas simulations indicated that it should be possible
to resolve two fluorescent sources of equal strength separated
by 0.8 mm. The linear dependence of measured and theoretical source separation for a range of 0.8–5 mm center-to-center
distance reflects the accuracy of the reconstruction algorithm.
On the other hand, the estimation of the source depth from the
sample surface was found to be systematically smaller than the
theoretical depth both for simulated and experimentally found
data for depth values exceeding 3 mm. Two factors contribute to
this: Due to diffusive light propagation the intensity is reduced
to a level that makes detection difficult (in particular in view of
autofluorescence contributions from surface structures) and secondly, the small field-of-view prevents accurate sampling of the
intensity distribution of the surface which is critical for proper
reconstruction of the source location and distribution. Though
simulations have revealed that by increasing the field of view to
16 16 mm, the accuracy of depth reconstruction can only be
marginally improved.
The claim of molecular imaging is to provide noninvasive
quantitative information on molecular processes in the intact
organism. These events commonly occur at a low frequency;
IEEE TRANSACTIONS ON MEDICAL IMAGING, VOL. 30, NO. 6, JUNE 2011
hence any molecular imaging modality should provide inherent high sensitivity. Calibration experiments revealed that
the FMT/MRI system was capable of detecting subpicomole
amounts of dye. Moreover, fluorescence intensity depended
linearly on the amount of fluorescent dye in the sample volume,
an important prerequisite for quantitative studies. The amount
of dye used covered the ranges typically reported for in vivo
applications.
The true strength of a hybrid system is that high resolution
structural images can be annotated with molecular information
derived from an inherently registered data set. The in vivo study,
which allowed identifying regions within a subcutaneous tumor
that are high in protease activity, clearly illustrates the potential of the approach. While the MRI signature clearly indicates
tumor heterogeneity based on differences in MRI contrast (domrelaxation times), the FMT inforinated by differences in
mation adds a molecular component to it: high protease activity is commonly associated with invasive (malignant) tumor
behavior [26]. The heterogeneity of fluorescence activity confirmed by histological analysis reflects the sensitivity of our
hybrid imaging system. Even more attractive FMT/MRI applications would be the correlations between physiological/functional MRI readouts (e.g., vascular permeability as a target for
tumor angiogenesis [27]) with FMT derived molecular information which demand for simultaneous measurement. The prerequisite for such applications is that the time resolution of the
slower of the two modalities is still sufficient to capture the biological process. Sequential measurements involving transfer of
the sample from modality A to B might not be possible, as this
would be too time consuming (including adjustments of individual modalities) so the sample state would be different.
Aside from collecting multiplexed data, hybrid systems
can offer improved data reconstruction for the low-resolution
modality, in our case for FMT. This has been used in PET/CT
systems, where the CT map is used to compute the tissue
attenuation correction required for reconstructing PET data
[28]. Correspondingly, in an FMT/MRI system, the anatomical
information provided by MR can be fed into both the forward
and the backward model of the FMT reconstruction [29], which
should render the FMT reconstruction more accurate. However,
the use of imaging priors bears the risk of over-constraining the
FMT reconstruction.
V. CONCLUSION
In conclusion, a novel hybrid setup for simultaneous acquisition of FMT and MRI data compatible for a small animal MRI
system has been described (Figs. 1 and 2). The interference between the two modalities was found negligible. Reconstruction
of fluorescence data revealed accurate localization of fluorescent sources, whereas the spatial resolution was demonstrated
experimentally to be in the order of 2 mm (see Fig. 3) whereas
based on the width of the individual profile we estimated that
sources separated by 0.9 mm should be distinguishable. Limitations in depth resolution may be attributed to the inherent
surface weighting of the reflectance geometry approach and to
the limited array size (32 32 pixels) and the limited field of
view of the current detector array. Improvements in depth reconstruction might therefore be achieved by using the setup in
STUKER et al.: HYBRID SMALL ANIMAL IMAGING SYSTEM COMBINING MAGNETIC RESONANCE IMAGING WITH FLUORESCENCE TOMOGRAPHY
transmission mode, which would in addition reduce the interference by autofluorescence. The next generation of detector
arrays will also have significantly increased matrix dimensions
and should due to the increased number of source detector pair
improve reconstruction accuracy. Hybrid FMT/MRI is sensitive
and we determined a sub-picomole detection limit for sources
embedded at a depth of 1 mm below the surface [Fig. 4(f)].
For this imaging domain the system performance was found to
be quantitative, as reflected by the linear relationship of fluorescence signal versus dye concentration. With its potential for
annotating structural/physiological with quantitative molecular
information FMT/MRI constitutes an attractive hybrid modality
for experimental biomedical researchers.
ACKNOWLEDGMENT
The authors gratefully acknowledge the contribution of
R. Keist for cell culturing, K. H. Altmann for AOI987 preparation and C. Ruettimann for fruitful discussions.
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