Topical Symposium on Advanced Molecular Imaging Techniques
In The Detection, Diagnosis, Therapy And Treatment Follow-Up of
Prostate Cancer. Palazzo Barberini, Rome - Italy, Dec. 6-7, 2005
Physica Medica
The Silicon Photomultiplier for Application in PET
N. Belcari1, A. Del Guerra, D. Herbert2, S. Moehrs
Department of Physics, University of Pisa and INFN-Pisa, Pisa, Italy
Abstract
The Silicon PhotoMultiplier (SiPM) represents an interesting advance among the family of phodetectors and could soon be a rival to the traditional
PMT for the development of dedicated PET imagers. The SiPM is a densely packed 2D array of Geiger-mode APD microcells, each one having
individual resistive quenching and multiplexed outputs. In this way the SiPM acts as a linear, high-gain photodetector for moderate photon flux
(Nphotons<Nmicrocells). We have characterised and tested SiPM samples from CPTA (Russia). We present a summary of the measurements of the device
primary operating characteristics and the results of its application to the readout of scintillators commonly used for PET.
KEYWORDS: Silicon Photomultiplier, SiPM, PET, breast cancer, prostate imaging
1
Address for correspondence: Nicola Belcari, PhD, Department of Physics, University of Pisa, Largo B. Pontecorvo, 3
I-56127 Pisa, Italy. E-mail: belcari@df.unipi.it
2
Now at SensL, Lee House, Riverview Business Park, Bessboro Road, Blackrock, Cork, Ireland
1
1. INTRODUCTION
During the last twenty years Positron Emission
Tomography (PET) has moved from a distinguished research
field in physiology, cardiology and neurology to become a
major tool for clinical investigation in oncology. Whole body
PET with 18F-FDG is nowadays a standard technique for the
diagnosis and staging of cancer [1]. In recent years there has
been an increasing focus on dedicated PET systems designed
for specific applications. Three of such applications are small
animal imaging for pre-clinical investigation, Positron
Emission Mammography (PEM) for breast cancer study and
prostate cancer imaging. Thanks to the technological
advances in high resolution detectors, especially developed
for the small animal imaging devices, many research groups
are working on the construction of dedicated devices for
PEM with high sensitivity and high spatial resolution [2-4]
based on similar technologies.
On the shadow of PEM, prostate imaging devices are at
the early stage of development [5], but they are expected to
show a rapidly growing technological advance. One of the
primary requirements for dedicated PET systems for prostate
imaging is that they should provide the maximum sensitivity,
minimum dead time and a good time resolution (reduction of
random count rate), whilst also providing a high spatial
resolution (about 2 mm FWHM). The issues for prostate
imaging are very similar to that for PEM, i.e., the position of
the source is more or less known (relatively small organ) and
there is a strong background source (bladder for prostate
imaging as opposed to heart for PEM) in the vicinity of the
signal source. In addition the imaging of the prostate is rather
difficult due to the strong scattering contribution from the
body (internal organs).
2. REQUIREMENTS FOR DEDICATED
INSTRUMENTATION
The three mentioned applications (small animal, breast
and prostate imaging) differ in optimal geometry and
absolute performance requirement. However, they all gain
from the development of novel detectors with improved
performance with respect to those used in the clinical
scanners. Current clinical PET systems are still based upon
the block detector concept that relies on quadrant light
sharing over single channel PMTs and is limited in terms of
spatial and timing resolution [6]. To achieve improved spatial
resolution, some systems now use position sensitive PMT’s
coupled to pixelated matrices of scintillators. The
development of novel photodetectors has a fundamental role
for the future of nuclear imaging. The possible solutions
proposed by various research groups have a two-fold
approach. On one side systems overcoming the coding
limitation have been introduced through APD based solutions
(one-to-one coupling) [7]. On the other hand, detectors based
on a scintillator block coupled to high granularity position
sensitive photodetectors (Anger camera principle) are under
study [8]. In the latter case, the limitation introduced by the
finite size of the crystal elements could be overcome by
measuring with high precision the center of gravity of the
light spot. For this latter approach a small semiconductor
photodetector with high gain is desired to produce compact,
high-resolution, high-sensitivity cameras with an adequate
flexibility for the construction of organ oriented scanners at a
reasonable cost. The so-called Silicon Photomultipier (SiPM)
is a good candidate for fulfilling these criteria.
3. THE SILICON PHOTOMULTIPLIER (SiPM)
The Silicon Photomultiplier (SiPM) [9-10] is a
silicon diode detector that shows great promise as a
photodetector for scintillators for both one-to-one and Anger
camera approaches [11-12]. The development of this device
started about 15 years ago at MEPHI (Moscow, Russia) and
has seen significant technological progress, in particularly in
recent years [13]. In short, the SiPM is a densely packed
matrix of small, Geiger-mode avalanche photodiode (GAPD)
cells (typically ~40 x 40 m2), with individual quenching
resistances for each cell. As a consequence of the Geigermode operation of each cell, these detectors have potentially
very fast timing, high gain (105 - 106) at low bias voltage
(~50V), an excellent single photoelectron resolution and a
low noise. They are rugged and insensitive to magnetic fields.
All the cell outputs are connected in parallel to produce the
summed signal. This microcell structure of the SiPM gives a
proportional
output
for
moderate
photon
flux
(Nphotons<Nmicrocells).
We have evaluated some SiPM samples from CPTA Russia,
[9] (figure 1), for the readout of scintillating crystal. The
testing involves the characterization of basic SiPM
parameters (gain, dark rate, I-V, stability, breakdown, timing)
as well as their performance as a photodetector for scintillator
readout. The measured gain is linear with the voltage over
breakdown and is of the order of 5105. We have measured a
single photoelectron resolution of the order of 15% (figure 2).
A typical limitation of the SiPM is the relatively strong dark
count rate, i.e., the number of photon counts per second
registered by a SiPM in the absence of impinging light
(figure 3). This effect is mainly due to the thermally
generated charge carriers, which depends upon density of
defects and volume. In this case a signal corresponding to a
single cell that fires is measured. However, multiple
photoelectrons are observed in the structure of the dark rate,
more than as expected by a pure Poisson distribution. This is
an evidence of optical cross-talk between microcells [14]. For
our samples, this cross talk effect was shown to be present
only at the level of a few percent. The study of the dark rate
revealed that the rate is rather high but can be completely
eliminated by putting a threshold at the level of a few
photoelectrons.
For the readout of a single finger crystal of LSO we have
tested a SiPM optimized for blue light. The result found was
an improvement with respect to previous measurement with a
green sensitive device [12]. The improvement observed is
due to the fact that the quantum efficiency of the SiPM
optimized for the blue light better matches the LSO output
light spectrum. The energy spectrum is shown in figure 4.
The measured energy resolution is about 37% (FWHM) at
511 keV. This spectrum has been obtained with a high over
bias, thus maximizing the probability of avalanche and hence
the photon detection efficiency (PDE). Even with the blue
sensitive device, and high over bias, the detected light yield is
rather low, of the order of 90 photoelectrons for a 511 keV
deposit. The poor overall PDE of our sample reduces both the
energy resolution capability and its intrinsic spatial
resolution. For these reasons, obtaining a good PDE requires
some special attention and is the subject of current SiPM
developments.
2
4. SiPM PET: MONTE CARLO SIMULATIONS
A new PET camera module design has been proposed by the
Department of Physics, University of Pisa [15]. The proposed
camera consists of a stack of modular layers, each one being
a relatively thin (5 mm), continuous slab of LSO (or LYSO)
scintillator viewed by an array of compact of 1 mm2 active
area SiPM detector elements on a 1.5 mm pitch. To provide
sufficient detection efficiency as well as intrinsic depth of
interaction (DOI) information, three of these modular layers
are stacked together to form a single detector head with a
total active thickness of 15 mm (figure 5). The DOI
information (z-coordinate) corresponds to the layer where the
position read out takes place and can be considered as a
discrete signal.
Geant4 based simulations [16] have been carried out to
optimize the design of a single detector head. For centered,
normally oriented annihilation photons we expect a gammaray detection efficiency, defined as the fraction of
annihilation photons that deposit at least 50 keV, of about
70%. Backscattering in the same detector head is less than
5%. Here, we consider an annihilation photon to be
backscattered if it deposits energy in an inner layer (closer to
the center) after it has already interacted in an outer layer of
the same head.
We obtain a spatial resolution of ~0.4mm FWHM in the
center of the detector, which degrades towards the detector
edges. Similarly, we obtain a very small displacement error in
the center of the detector, which increases significantly
towards the detector edges. Using the skewness- and
barycenter-based
log-likelihood
method
[15],
the
displacement error remains well below the maximum parallax
error of 1 mm over the whole detector surface. It should be
noticed that a very low quantum efficiency of 2.5%, as
measured for our sample of SiPMs, was used in this
simulation.
5. SUMMARY
The SiPM represents a compact device that has a behavior
similar to a traditional PMT but with the advantage of a faster
response, lower operational voltage and the design flexibility
of a semiconductor based detector. However, the SiPM still
needs to be improved in terms of photon detection efficiency
optimised for blue-green, especially in comparison to PMT,
so as to be successfully used for scintillator readout. The
development of SiPM matrices and compact electronics is
strongly suggested for a successful implementation of the
SiPM technology in PET instrumentation. Our Monte Carlo
simulations indicate that a simple PET camera design based
on a continuous crystal could yield both high intrinsic spatial
resolution and high sensitivity. The continuous crystal also
provides continuous sampling giving better quantification
and flexible data binning.
The characteristics and performance of the SiPM justifies the
interest of various research groups world wide in the
development of SiPM based photodetectors. In fact, the
extreme compactness of such photodetector makes it a
flexible and suitable solution for the design of modular, low
cost and non-conventional geometry scanners as required for
dedicated PET applications, like prostate PET imaging. As a
bonus the non-sensitivity of SiPM to magnetic field could
allow the construction of multi-modality PET-MR scanners
that represent the new frontier for the field of oncology
imaging.
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3. Figure Captions
Figure 1: The CPTA SiPM test device. On the left, a
close up of the microcell structure; on the right, the SiPM
mounted on a test package, as used for testing.
Figure 2: Typical single photoelectron spectrum from a
SiPM
Figure 3: Dark rate as a function of the threshold level
for various bias voltages, obtained with a SiPM produced by
CPTA.
Figure 4: Energy spectrum of a SiPM coupled with a
LSO crystal 1 mm  1 mm  10 mm illuminated with a 22Na
source
Figure 5: Design of a detector head composed by three
layers of a SiPM matrix + LSO scintillator slab.
Figure 3
Figure 1
Figure 4
Figure 2
Figure 5
4