IEEE TRANSACTIONS ON NUCLEAR SCIENCE, VOL. 52, NO. 3, JUNE 2005
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Architecture of a Dual-Modality, High-Resolution,
Fully Digital Positron Emission
Tomography/Computed Tomography (PET/CT)
Scanner for Small Animal Imaging
Réjean Fontaine, François Bélanger, Jules Cadorette, Jean-Daniel Leroux, Jean-Pierre Martin, Jean-Baptiste Michaud,
Jean-François Pratte, Stéfan Robert, and Roger Lecomte
Abstract—Contemporary positron emission tomography (PET)
scanners are commonly implemented with very large scale integration analog front-end electronics to reduce power consumption,
space, noise, and cost. Analog processing yields excellent results
in dedicated applications, but offers little flexibility for sophisticated signal processing or for more accurate measurements
with newer, fast scintillation crystals. Design goals of the new
Sherbrooke PET/computed tomography (CT) scanner are: 1) to
achieve 1 mm resolution in both emission (PET) and transmission
(CT) imaging using the same detector channels; 2) to be able to
count and discriminate individual X-ray photons in CT mode.
These requirements can be better met by sampling the analog
signal from each individual detector channel as early as possible,
using off-the-shelf, 8-b, 100-MHz, high-speed analog-to-digital
converters (ADC) and digital processing in field programmable
gate arrays (FPGAs). The core of the processing units consists of
Xilinx SpartanIIe that can hold up to 16 individual channels. The
initial architecture is designed for 1024 channels, but modularity
allows extending the system up to 10 K channels or more. This
parallel architecture supports count rates in excess of a million
hits/s/scintillator in CT mode and up to 100 K events/s/scintillator
in PET mode, with a coincidence time window of less than 10 ns
full-width at half-maximum.
Index Terms—Computed tomography, digital signal processing,
field programmable gate array (FPGA), positron emission tomography.
I. INTRODUCTION
O
NE goal of contemporary research in biological imaging,
centered on the genetically-modified mouse in this
postgenomic era [1], is to achieve better resolution images with
Manuscript received June 3, 2003; revised May 12, 2004. The work was supported by the Natural Science and Engineering Research Council of Canada and
Le Fonds Québécois de la Recherche sur la Nature et les Technologies.
R. Fontaine, F. Bélanger, J.-D. Leroux, J.-B. Michaud, and S. Robertis are
with the Université de Sherbrooke, Sherbrooke, QC J1K 2R1, Canada (e-mail:
Rejean.Fontaine@USherbrooke.ca).
J. Cadorette is with the Centre Hospitalier Universitaire de Sherbrooke
(CHUS), Sherbrooke, QC J1H 5N4, Canada.
J.-P. Martin is with the Department of Physics, Université de Montréal, Montréal, QC H3C 3J7, Canada.
J.-F. Pratte is with the Instrumentation Division, Brookhaven National Laboratory, Upton, NY 11973-5000, USA.
R. Lecomte is with the Animal PET Imaging Group, Department of Nuclear
Medicine and Radiobiology, Université de Sherbrooke, Sherbrooke, QC J1H
5N4, Canada.
Digital Object Identifier 10.1109/TNS.2005.850484
less exposure and shorter acquisition time. Current positron
emission tomography (PET) scanners designed for small animal imaging achieve a spatial resolution that makes it possible
to detect lesions just over 2 mm in size. For applications in
the genetically modified mouse, a 1-mm accuracy would be
required all across the field of view of the camera. Furthermore, future PET scanners should be designed to be able to
accommodate new incoming detector and signal processing
technologies that will help enhance its accuracy. Currently,
one major problem in PET scanner designs resides in their
low upgrading capacity resulting mainly from their analog
electronics architecture. Up until very recently, digital signal
processing technologies could not be considered due to size
and power limitations, and effort in PET design was focused on
the necessity to reduce data processing and power using analog
electronics with high-level channel encoding. Unfortunately,
this approach makes it difficult to integrate new technologies
and new scientific discoveries.
In another respect, with the use of new, highly specific
PET radiopharmaceuticals that are being developed to very
selectively target desired tissues, cells, or cell subsystems, it
is becoming increasingly difficult to accurately localize the
uptake structure highlighted by the PET tracers within the
body [2], [3]. One must rely on anatomical details obtained
with another imaging modality, either on a different scanner
or with dual-modality systems incorporating a PET camera
and a computed tomography (CT) scanner [4]–[6]. This is
currently achieved by using two sets of sensors placed side by
side: one for PET imaging and the other one for CT imaging.
These devices are cumbersome and expensive. Moreover,
they suffer limitations due to sequential imaging and possible
subject motion between successive scans. To overcome these
problems, the concept of multimodality imaging on the basis
of a common detection system based on high-performance
avalanche photodiode (APD) detectors was first proposed in
1998 [7]. However, due to the complexity of the signal processing from the common PET/CT detector front-end, such an
approach would be very difficult to implement in a fully analog
scanner design.
We propose to realize this original scanner concept using a
novel electronic architecture relying heavily on digital signal
processing. This approach offers much more flexibility and
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IEEE TRANSACTIONS ON NUCLEAR SCIENCE, VOL. 52, NO. 3, JUNE 2005
Fig. 1. PET/CT scanner basic operation principle.
opens new possibilities for more sophisticated signal processing. The counterpart is higher power consumption and a
huge amount of data to process in real time.
II. PET AND CT IMAGING PRINCIPLES
PET radiopharmaceuticals enable the investigation of biochemical processes at the cellular and molecular level in vivo.
-fluorodeoxyglucose, an analog of glucose,
For example,
will accumulate faster in cancer cells that are more metabolically active than normal cells, making it possible to detect tumors and determine their grade. The disintegration of the PET
radioisotope within tissues produces a positron that will soon
annihilate with an electron, emitting two collinear 511-keV photons 180 apart. The goal of the PET scanner is to detect and
count a large number of these annihilation photon pairs in coincidence to eventually reconstruct the image of the radioactivity
distribution using tomographic techniques (Fig. 1). Typically,
to a few mCi (10–100 MBq) of radioacseveral hundreds
tivity is injected to the subject under study, yielding individual
detector singles count rate in the range of 10-100 kcps.
In CT imaging, the transmission of X-rays from an external
source through the subject is used to obtain an image of the
tissue density. Tomographic data are acquired by rotating the
X-ray source around the subject while recording the X-ray flux
transmitted through the tissues in opposite detectors, as illustrated in Fig. 1. As the flux per unit area can be fairly large
photons
[8]), CT detec(typically
tors conventionally operate in current mode.
III. PET/CT SCANNER REQUIREMENTS
A prerequisite for using the same detection system for PET
and CT imaging is to be able to detect and count individual highenergy annihilation photons (511 keV) and low-energy X-rays
(30–60 keV) with the same detector and front-end electronics.
This has been shown to be achievable with APD-based multilayered crystal detectors [7] and fast, low-noise, high dynamic
range front-end electronics [9]. One further requirement is that
the signal from every individual detector must be processed independently to cope with the high-event rate in CT mode and
the crystal identification in PET mode.
The accuracy and contrast of PET images are dependent on
a number of physical and electronic parameters, including the
number and size of sensors, the detector speed and front-end
electronic noise, and the signal processing sophistication and
computational power which affect the ability to perform correction on acquired data.
The size of individual detectors essentially determines the intrinsic spatial resolution in PET, which is given to a first approximation by half the detector dimension [10], [11]. In small
diameter scanners used for animal PET, the parallax error resulting from the depth of interaction (DOI) within detectors has
a nonnegligible effect on resolution. Parallax is the uncertainty
on the position of the virtual line traced between thin and tall
coincident detectors when viewed at an angle. Image fuzziness
increases trigonometrically with parallax as the coincidence occurs further away from the center of cylindrical scanners, unless
DOI is measured to reduce location uncertainty roughly to the
same value in all directions. One way to measure DOI is to use
a phoswich assembly of scintillation crystals having different
decay times, which can be identified by analyzing their signal
pulse shape [12]–[17]. The major challenge with this approach
resides in the capacity to perform accurate DOI identification
with minimal signal processing and CPU power.
The detector coincidence timing resolution and time skew between detectors mostly determine the time window width that
can be used for coincidence detection of annihilation photon
pairs in PET. A narrow time window is important to limit the
rate of random events, which can seriously degrade the image
signal to noise ratio at high-count rates. With APD coupled to
bismuth germanium oxide (BGO) detectors and conventional
analog timing techniques, a time window larger than 20 ns must
be used [18]. A narrower window of less than 10 ns can be considered with faster, high-luminosity scintillators, such as APD
coupled to lutetium oxyorthosilicate (LSO) [19], [20]. Using
digital techniques, a time window of the order of 10 ns is targeted with APD-BGO detectors [21].
In single photon counting CT mode, only the number of lowenergy X-ray events per pixel needs to be recorded. Besides
spatial resolution, which is determined by the size of individual
detector pixels and the geometry of the X-ray source, the most
critical parameter of a CT scanner based on PET detectors will
thus be its count rate capability. To deal with realistic X-ray
fluxes from a low-power X-ray tube, the single photon count rate
.
of individual detectors should reach at least 2
As the X-ray energy is one order of magnitude lower than the
PET 511 keV radiation, the front-end analog electronics must
be designed with a wide dynamic range and a short pulse decay
time to accommodate the large amplitude range and high rate of
detector signals.
IV. EXISTING PET ARCHITECTURES
Most existing PET scanners are built around analog subsystems in order to extract only fairly basic information from the
detected events, like detector address, timestamp, and energy
FONTAINE et al.: ARCHITECTURE OF A DUAL-MODALITY, HIGH-RESOLUTION, FULLY DIGITAL PET/CT SCANNER
693
Fig. 2. A typical analog PET scanner front-end architecture.
(Fig. 2). After an annihilation photon hits a crystal detector (like
BGO or LSO), the scintillation light re-emitted by the crystal is
collected by a photomultiplier tube or an APD that transforms
and amplifies this faint light signal into a measurable charge
pulse. This charge pulse is further amplified and converted into a
voltage pulse by means of a charge sensitive preamplifier (CSP),
and then fed to the front-end analog processing electronics. A
timestamp is generated by a leading edge discriminator (LED)
or a constant fraction discriminator (CFD) [13]. The LED uses
a fixed threshold to trigger the timing pulse, which results in
a simple design but the timestamp is amplitude dependent. An
amplitude independent timestamp can be obtained with the CFD
by setting the triggering threshold at a fixed fraction of the maximum signal amplitude, but at the cost of a more complex design. Moreover, the CFD internal delay must be tailored to the
scintillator decay time used. The performance of these two time
discrimination techniques degrades in presence of noise [13],
[19], [20]. The fast timing pulse generated by the discrimination
circuit is fed to analog or digital units for coincidence verification. Most PET scanners also extract the energy of the signal to
select true photoelectric events and reject events scattered in the
object or detection system.
V. NOVEL PET/CT ARCHITECTURE
The most promising approach to overcome several of the
problems enumerated above and to achieve sufficient system
flexibility to be able to perform CT scans and PET scans using
the same detectors is to use digital signal processing. The
proposed PET/CT architecture, depicted in Fig. 3, consists
of three main subsystems. The data acquisition subsystem
(DASS) includes the analog front-end, the analog-to-digital
converters (ADC), the digital signal processing units, a 500-V
APD bias supervisor unit, a clock controller, and a power unit.
The coincidence detector subsystem (CDSS) performs the secondary sorting of data for coincidence identification. Finally,
the image processing software performs the data normalization,
corrections, and rebinning for tomographic reconstruction of
the image.
In the DASS, each sensor has its own data processing
channel composed of the scintillation crystal and APD, the
CSP followed by an anti-aliasing filter, the ADC, and a field
programmable gate array (FPGA)-based processor unit (Fig. 3).
The MAX1198 100-MHz ADC from Maxim [22] was selected
for its low power consumption and high sampling rate, which
is necessary for accurate timing with high-speed crystals like
) or LuAP (
) [12], [13].
LSO (
To perform all needed computations for PET and CT data processing, the FPGA architecture is also critical. The digital signal
processing unit is based on the XC2S300E Spartan-IIE from
Xilinx Inc. This FPGA includes 64 kb of distributed RAM, four
delay lock loops, and between 93–300 K equivalent gates [23].
This FPGA can handle clock frequencies beyond 200 MHz,
enough for the first phase of the prototype system. The FPGA is
preferred to a digital signal processor for its capacity to perform
simple distributed computational tasks and its low cost per I/O.
The ACQuisition FPGA architecture (Fig. 4, exploded from
Fig. 3) is composed of 16 free running ADC controllers based on
a local 150-B first in, first out (FIFO), some I/O port controllers,
and a level detector. When the level detector detects a signal with
enough energy, the local FIFO stores up to 150 samples as one
acquisition frame. The acquisition length is based on the minimum time needed to sample the signal rising edge of the slowest
for BGO – at the output of the CSP.
scintillator –
At the end of data sampling, a finite state machine will transfer
the data from the FIFO to the main RAM where different processes may take place. In the case of multilayer crystals, the first
process consists in determining the time stamp generation and
energy computation. Subsequently, DOI is performed by identifying the crystal of interaction using signal theory methods [24].
Although, the proposed system design could possibly be applied
to a wider range of DOI solutions, only phoswich detectors are
being considered for now. For the time stamp generation, digital
versions of analog LED and CFD circuits, as well as more complex techniques that are accessible in a digital architecture can
be implemented and assessed [21]. It is worth pointing out that
the DOI and timing discrimination circuits can be readily modified and reprogrammed in the FPGAs in order to experiment
various strategies or optimize data processing for other crystal
scintillators. Once the computation is completed, the channel
number is attached to the data and sent through a parallel 8-b
port to the sorting FPGA. This XC2S300E SpartanIIE FPGA
performs a preliminary timestamp sorting of the data and prepares data words for transmission through a high-speed serial
port to the CDSS. In CT mode, data are simply histogrammed
in counters activated by the level detector. Counters can be read
out on request from the main computer.
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IEEE TRANSACTIONS ON NUCLEAR SCIENCE, VOL. 52, NO. 3, JUNE 2005
Fig. 3. High-level architecture of the dual-modality PET/CT scanner and exploded view of one DASS board architecture.
VI. SELF-ADJUSTING APD POLARIZATION
APDs need to be properly reverse biased (between about 350
and 500 V) for optimum timing performance with slow scintillators like BGO, and for maximum signal-to-noise ratio in
measuring low-energy X-rays. The DASS includes a 500-V regulator that can adjust the bias voltage for APDs. The system
is supervised by a microcontroller implemented with a voltage
protection filter against sudden drop or surge of the bias voltage.
The subsystem design is depicted in Fig. 5.
The microcontroller controls the polarization voltage with
an 8-channel, low-power, 10-b TLV5608 digital-to-analog converter (DAC) serially linked with a SPI protocol. DAC voltage
is amplified by a factor of 200, and regulated by a linear regulator built around an operational amplifier and a low-pass filter.
The low-pass filter acts as an additional security in the system to
ensure slow APD biasing at a slew rate not exceeding 5 V/s. A
analog multiplexer is used to sense the regulator output
low
voltage and a 12-b ADC embedded in the microcontroller samples the individual APD bias through a voltage divider to ensure
that the biasing circuit is working correctly. The microcontroller
can also send data to the scanner main computer for operator
monitoring.
VII. ACCURACY OF THE TIMESTAMP
To achieve coincidence timing resolution better than 10 ns,
the accuracy of the timestamp is very important. In analog
PET systems, the detector ring is divided into sectors within
which no coincidence is allowed. All events triggered within
one sector are ORed and forwarded to coincidence units that
compare events from opposite sectors through AND gates
[18]. This approach requires a huge number of connectors and
properly shielded cables to ensure noiseless transmission paths
between the detector front-end and the coincidence units. In
the architecture proposed, high-frequency, low-cost Ethernet
cables replace analog cables. The timestamp synchronization is
FONTAINE et al.: ARCHITECTURE OF A DUAL-MODALITY, HIGH-RESOLUTION, FULLY DIGITAL PET/CT SCANNER
Fig. 4.
Exploded view of the acquisition FPGA (DASS subsystem), showing the various digital processing blocks.
Fig. 5.
The APD bias regulator and voltage supervisor.
the critical part for the accuracy of the coincidence system. For
a 1024-sensor PET scanner, 16 DASS boards will be equally
distributed around the ring path and each board will be precisely
synchronized to provide accurate timestamps. Three factors can
influence the accuracy: the delay induced by components and
cables; the rise time and fall time of the clock signal; and noise.
The architecture proposed to overcome this problem is to distribute a single low-frequency, low-jitter, high-precision clock
source (16 MHz) to the entire system through an off-the-shelf,
low-voltage positive emitter-coupled logic (LVPECL) buffer
located on the CDSS (Fig. 6). Noise immunity is ensured by
the use of differential clock distribution within the system. On
the DASS board, a differential LVPECL frequency synthesizer
is multiplying the 16 MHz to the required 100 MHz. Then low
voltage zero delay buffers provide the 100 MHz clock to all
FPGAs and ADCs. With such a topology, the worst theoretical
clock skew between components on different DASS boards is
850 ps, while the typical value should not exceed 200 ps. Since
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events are saved with timestamp and detector addresses, clock
skew between different detector channels can be estimated by
analyzing the coincidence time distribution for each detector
pair. A software calibration will be performed in order to
correct residual time skew originating from cable length and
component delay.
VIII. CONCLUSION
A dual-mode PET/CT concept using the same detectors
has been proposed. An architecture offering the opportunity
to integrate new digital signal processing principles in the
medical imaging field has been designed. The new proposed
PET/CT scanner design, which is characterized by excellent
accuracy, modularity, flexibility, and self-adjusting systems,
will provide unprecedented performance and reliability for
molecular imaging research applications.
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Fig. 6. Distribution of the clock for a scanner with 1024 channels.
ACKNOWLEDGMENT
The authors would like to thank the Université de Sherbrooke
for the technical support in this project. Finally, the authors
want to acknowledge the scientific contribution of P. O’Connor,
Brookhaven National Laboratory, in the first phase of this
project.
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