Download Title: Multimodal TOF PET and/or SPECT probe operating in low

Title: Multimodal TOF PET and/or SPECT probe operating in low-field MRI for diagnosis and follow up of
Prostate Cancer
Acronym: MUPROPESCAM
Prostate cancer and the role of imaging
Prostate carcinoma, the second most common cause of cancer death among men in western countries, is a
heterogeneous disease, it ranges from asymptomatic to rapidly progressive systemic malignancy. The
prevalence of prostate cancer is so high that it could be considered a normal age-related phenomenon (1).
Four main clinical challenges exist in prostate cancer treatment and management—diagnostic accuracy; risk
stratification, initial staging, active surveillance, and focal therapy; prostate-specific antigen relapse after
radiation therapy or radical prostatectomy; and assessing response to therapy in advanced disease (1). Better
understanding and managing this disease is needed. The issue is how anatomic, functional, and molecular
imaging techniques might add to more accurate characterization of disease at initial biopsy; how better
staging or evaluation of early response to therapy would allow better patient management, improving
effectiveness of therapies or avoidance of unnecessary treatments; and how and which imaging tools can best
be used to inform prostate cancer clinical trial designs and accelerate the evaluation of potential novel
breakthrough therapies. Traditionally, imaging has played a relatively minor role in the management of
prostate cancer, especially of the localized ones but it may play a crucial role in clarifying several important
issues. In patients with negative systematic biopsies, imaging may be helpful in localizing prostate cancers
that would otherwise be missed. For patients already diagnosed with cancer, imaging can assist in better
defining the location and margins of the tumor. Finally, there is increasing interest in performing imaging
guided focal prostate therapy. New imaging techniques will be critical to the success of this new strategy.
Current standard imaging techniques, such as ultrasound, MRI, CT, and nuclear medicine, cannot detect early
disease, and they provide limited information for disease staging. Several promising emerging techniques are
under investigation, either alone or in conjunction with standard imaging techniques but they provide less
than precise predictions. It is evident the need to consider fundamental changes in the approach to the
diagnosis of prostate cancer. MUPEMPROCA project propose a new approach to the diagnosis and follow up
of prostate cancer: the use of a multimodality imaging detector that can play a crucial role by merging
anatomical and functional details from simultaneous PET or SPECT and MRI scans. MRI is the ideal
complement to radionuclide imaging techniques for soft tissues and can have adjunct value by using
dynamics of contrast enhancement. Nevertheless, due to sub-optimal prostate imaging geometries, generic
scanners prevent separation of the signal from surrounding organs with sensitivity, spatial resolution and
contrast inferior to what is achievable with dedicated prostate imagers. MUPEMPROCA project is aimed at
developing an endorectal PET-TOF and SPECT MRI probe with new radiotracers and (possibly) low field
MRI. Exploiting the TOF capability allows an increase in the SNR/NECR and also permits elimination of
bladder background (8). The internal probe will be used in coincidence with an external dedicated detector
and/or a standard PET ring. Performance is dominated by the endorectal detector with improvements in both
spatial resolution and efficiency. The electronics must measure coincidences with a precision of 300 ps or
less, and be small enough to be connected to the internal detector. For compactness, MRI compatibility and
timing resolution, Silicon Photomultipliers (SiPM) will be used. One of the major drawback of MRI exams
comes from the high cost. Another factor is the lack of appropriate MRI biopsy guiding methods which are
needed when evaluating lesions observed only by MRI. A low-field open MRI unit in multimodality with
PET and SPECT could be a solution to these problems: lower equipment costs lead to lower prices and the
open configuration allows easy access in biopsies. A comparison of high field MRI scanner performances
with a low field one’s will be performed. Using low field scanner would decrease the cost dramatically, so
facilitating the diffusion of the proposed multimodal detector in the clinical practice without sacrifying the
performances of the detector system. An inherent limitation of low-field magnets for diagnostic MR imaging,
the reduced signal-to-noise ratio attainable compared with that attainable at higher field strengths can be
overcome with the use of MR contrast agents that possess higher r1 relaxivity compared with conventional
MR contrast agents. Low field MRI scanner with contrast enhanced technique showed to perform as well as
1
high field one’s and even better in some cases (9,10). Moreover low field would minimize the impact of the
PET/SPECT detectors on the MRI scanner and vice versa. The main issue with radionuclides scanners for
prostate is the lack of good specific imaging agents. New promising radio tracers already proved to be
successfull on animals. An important, crucial merit of the project is the search for new radiotracers, and study
of toxicity for humans application. Tests on big animals and human clinical trials will be performed. In this
way the proposed multimodal detector will demonstrate unprecedented ability in early diagnosis of the
prostate cancer providing also the right tools for guiding biopsy and follow up of the therapy.
PET TOF and SPECT endorectal probes
SPECT probe
The imaging performance of a single photon camera is dominated by its collimator. As the parallel-hole
collimator efficiency is independent of the collimator to source distance but the spatial resolution is linearly
proportional to the collimator to source distance, optimal performance is achieved by placing the camera as close
to the prostate as possible. This suggests a transrectal geometry, as shown in Figure 1, as it places the front face
of the collimator less than ~ 1cm from the prostate. Looking at the anatomy of the prostate and at the most
probable sites of the prostate cancer (the periphery of the organ (13) ) the distance of the lesion vary form ~ 1 to
~ 5 cm. A suitable standard camera geometry is shown in Figure 2.a. The small space available, the need for
maximizing the efficiency and the dimension of the prostate brought us to consider different geometry (Fig. 1 c).
In the parallel hole collimator geometry we propose to insert the scintillator pixels into the collimator tungsten
holes, for ~ 1/3 of the collimator thickness. This would allow a compact detector while getting an even higher
efficiency and probably better spatial resolution (12). Full simulation will be needed to optimized the design this
setup. From Fig. 1.a one can easily understand that to avoid angular undersampling of the prostate (at least) 2
scans should be performed. This problem can be overcome with the geometry showed in Fig. 2. Using fan beam
diverging collimator would allow making only one scan without angular undersampling. Preliminary calculation
show the advantage of this geometry (Fig.2 b). We are considering also other collimator techniques, namely a
multipinhole collimator with distance detector - collimator possibly variable to allow fine tuning the
performance of the detector as function of the distance lesion to detector. This probe can work standalone or can
be used in conjunction with one or more external panels and/or with an external gamma camera (see net
paragrapgh for the rationale) (Neal should comment and complete for the multipinhole collimator option and
anything else needed, Stan should make arguments coming from his measurements (focal plane tomography also
for single photon?) and anything else). The work and the merits of all the team members should be quoted (see
the citation of Neal work for the concept of probe + external PET). Figure have to be re-done
QuickTim eﾪ e un
decom pr essor e
sono necessar i per visualizzar e quest'im m agine.
Fig. 1. a . b . c.
2
QuickTim eﾪ e un
decom pr essor e
sono necessar i per visualizzar e quest'im m agine.
QuickTim eﾪ e un
decom pr essor e
sono necessar i per visualizzar e quest'im m agine.
Fig. 2 a . Layout of the fanbeam collimator. b. Comparison between parallel hole, diverging and fan beam
collimator efficiency and spatial resolution
PET probe
Conventional, external ring PET geometries suffer from two limitations for prostate imaging. First, since both
annihilation photons must be detected in PET, attenuation of photons emitted from the centrally-located prostate
is severe. Second, spatial resolution in PET images reconstructed from external ring scanners is relatively
modest (5–8 mm). It is challenging to detect uptake in tumors smaller than ~1ml in volume, which translates to
a 12.5mm Ø spherical lesion. External dedicated “ring” PET scanner optimized for prostate imaging have been
recently been constructed moving the ring close to the patient to maximize detection efficiency for a given
detector size, and angles the ring to minimize the effects of attenuation. Reconstructed resolution in phantom
studies is reported at 4 mm FWHM (5). Nevertheless photon attenuation is still an issue and image distortions
outside the prostate, due to the scanner design, will likely render detection of metastatic lesions in lymph nodes
more difficult than with conventional scanners. At a 2005 workshop on technology for prostate imaging held in
Rome (organized by the PI and other team members (see reference)), two relevant designs for optimized PET
prostate imaging were presented. Both made use of an endorectally placed miniature PET detector working in
coincidence with an external detector. Advantages of placing a high resolution PET detector close to the prostate
are several. Attenuation is reduced since one of the annihilation photons need only traverse the small distance
from the prostate to the internal detector. Spatial resolution can be high (as shown below) and is relatively
insensitive to the intrinsic performance of the external detector. And finally, because the orientation and
translation of the internal probe/detector relative to the external detector is tracked as a function of time, motion
artifacts from the region of the prostate can be compensated in the image reconstruction process. The two
proposed instruments are shown in Figs. 3 a.b.c.
QuickTim eﾪ e un
decom pr essor e
sono necessar i per visualizzar e quest'im m agine.
QuickTim eﾪ e un
decom pr essor e
sono necessar i per visualizzar e quest'im m agine.
3
Fig. 3. a. b. Endorectal probe with one (or two) external panel detectors. b. Endorectal probe and an external
standard PET. One or more panel detectors can be added for high efficiency in detecting linphonodes.
In the first a high-resolution endorectal detector works in conjunction with a planar external detector. The second
employs a similar high-resolution internal detector but mates it to a conventional, external ring PET scanner.
Both approaches have merit. We will propose an hybrid. Let’s start considering the probe and the external PET.
One significant reason for this choice is that the probe—itself being a PET detector—can be connected with
present PET instruments at a point of modularity. It can be incorporated as an add-on detector to present and
future PET instruments in the same way that specialized coils extend the capabilities of existing MRI
instruments at low incremental cost. Another significant reason is that use of the conventional scanner allows
acquisition of complete tomographic data rather than limited-angle tomographic information. The limited angle
tomographic problem could be mitigated by using 2 detectors profiting of the TOF capability (see later in the
TOF paragraph and in ref. 6). The probe should not significantly alter existing PET imaging protocols for
detection of metastatic disease and in addition will provide high resolution information within the prostate that
may allow earlier detection of aggressive disease. The advantages of the other geometry (probe + one (or more)
external detectors)) over conventional PET geometries parallel those for the transrectal single gamma system.
By placing a detector transrectally, it is extremely close to the prostate, and the external detectors is also placed
as close to the patient’s abdomen as possible. This increases the sensitivity and also has the potential to improve
the spatial resolution. When the object is much closer to one detector in a PET camera than the other, the spatial
resolution is determined by the closest detector. If this detector has higher resolution than conventional detector
(while maintaining high efficiency), then high-resolution, high-efficiency imaging can be achieved. In order to
maintain high spatial resolution, the transrectal probe needs to measure the third coordinate with accuracy
similar to what is measured in the other two coordinates. This will be solved by using a detector that measures
the “depth-of interaction” (references). Use of small PET detectors having high resolution in conjunction with
more conventional PET rings was proposed independently by Clinthorne and by Park and also has been explored
by Tai and co-workers for high resolution breast imaging among other application (references). The high and
low resolution data should be combined in the image reconstruction process. Because the tomographic
measurement process already entails significant uncertainty as to where an annihilation occurs along each lineof-response (LOR), spatial resolution is primarily determined by uncertainty in directions transverse to the LOR.
The component of uncertainty (FWHM for an equivalent Gaussian response) due to detector resolution at any
point along the LOR in the transverse direction is approximately
R D≈ 2 . 35
2
1− α sin 2 θ 1 σ 2D1 cos 2 θ 1 σ 2C1
α 2 sin 2 θ 2 σ 2D2 cos 2 θ 2 σ 2C2
where θ1 and θ2 are the angles of LOR incidence on each of the detectors, σDi and σCi, quantify the rms
uncertainty for each detector in the depth and circumferential directions, and α is the fractional distance along
the LOR at which the uncertainty is desired. For example, if the distance from the prostate probe to a source in
the prostate is 3cm and the total LOR length is 40 cm, α = 3/40. Note that when α is small, spatial resolution is
dominated by the characteristics of the detector 1, which in this project is the prostate probe. Assuming the
prostate imaging probe is capable of achieving 1mm spatial resolution in the circumferential direction and that
the resolution of detector 2 is 6 mm FWHM, resolution at 3 cm from the probe face would be approximately
1.03 mm FWHM for normally incident photons. Spatial resolution (as quantified by LOR uncertainty) as a
function of distance from the probe for an ~80 cm diameter external ring having nominal resolution of 6mm
FWHM is plotted in Fig. 4a.for internal probe resolutions of 1, 2, and 3mm FWHM. Note that resolution can be
excellent even at 8 – 10 cm from the probe face. Acolinearity of the annihilation radiation is not a significant
issue. The following expression for resolution-loss introduced by acolinearity is easily derived from geometric
arguments with the small-angle approximation tanθ ≈ sinθ ≈ θ where the angle is in radians and the uncertainty
due to residual momentum at annihilation in soft-tissue is 0.0088 radians FWHM:
R A≈
d 1d 2
× 0 .0088= α 1− α d 1 d 2 × 0 . 0088
d1 d 2
4
In the expression d1 and d2 are the distances (mm) from the point of annihilation to each detector. It
is easily verified that the effect is negligible within 5 cm of the probe (0.38 mm for the above
example, which adds in quadrature with other uncertainties).
6
20000
18000
16000
Events Detected
Resolution (mm FWHM)
5
4
3
3 mm
14000
BGO+LSO Probe 1
12000
BGO Only Probe 1
10000
BGO+LSO Probe 2
8000
BGO Only Probe 2
6000
2 mm
2
4000
2000
1
0
Probe resolution = 1mm FWHM
0
0
5
10
15
20
25
30
35
Axial Extent (# 5mm rings)
0
5
10
15
20
25
30
35
40
Distance from probe face (cm)
Fig. 4 a. Spatial resolution of the system as function of the resolution of the probe. b. Efficiency of the system as
function of the dimension of the probe.
The potential for a high resolution endorectal PET prostate probe to improve detection and characterization of
prostate cancer is high. Practical uncertainties that remain will be addressed in this project: (1) technology
necessary to create a compact high-resolution and clinically safe endorectal PET detector, (2) the ultimate effect
of the high resolution information on reconstructed images, which because of the limiting imaging geometry
does not provide an isotropic increase in spatial resolution, and (3) performance of a prostate probe
demonstrator in phantom imaging studies, in animal model and clinical studies. Preliminary Work has been
already done by the PI and the other members of the team, in many directions (hardware and software
(references (IEEE (Franco) and others (Stan, Neal)).
Time Of Flight (TOF)
Our aim is to build the PET probe with very good timing resolution capability (< 300 ps FWHM). One of the
major advantages of improved coincidence timing is the reduced random event rate. In PET, the random event
rate for an individual chord is given by R2=2R1R2T, where R is the random event rate for that chord, and R1
and R2 are the single event rates for two detector elements that form that chord, and T is the hardware
coincidence timing window width. The total number of random events in the image is the sum over all the
chords, thus is proportional to T. The mean contribution to the image from random events can be measured and
subtracted, but the noise resulting from the statistical variations in this rate remains. The practical effect of the
residual noise from random coincidences depends on the imaging situation and the task. However, it can be
estimated using the noise equivalent count rate (NECR), a common figure of merit for comparing tomograph
performance. The NECR is given by NECR = T2/ (T+S+2R) where NECR is the noise equivalent count rate, T is
the true coincidence event rate, S is the scattered event rate, and R is random event rate. Each positron could be
directly determined by accurately measuring the difference in arrival times of the two annihilation photons. In
other words, the position of the positron would be constrained to a point rather than a line, so three-dimensional
images could be obtained without a reconstruction algorithm. The accuracy of the measured position along the
line is x=c/T where x is the position error, c is the speed of light and T is the error in the timing
measurement. A timig resolution of 300 ps, would constraint the positron position to a line segment
approximately ~ 4.5 cm long (< dimension of the prostate). Constraining the positron position to a line segment
cm long do not improve the spatial resolution, but it reduces the statistical noise in the reconstructed image if the
line segment is shorter than the size of the emission source. This multiplicative reduction factor (corresponding
to the reduction in noise variance) is given by f = D/x = 2D/cT where D is the size of the emission source,
c is the speed of light, and is the timing error (7,8,11). For organs the size of the brain (D ~ 20 cm), this factor is
greater than unity (implying some noise reduction) for timing resolution 1.3 ns. For whole-body imaging, the
object size is larger (D ~ 35cm) and f >1 for timing resolution < 2.3 ns. Fig. 5 shows the improvement of SNR
5
and NECR as function of the timing resolution. The origin of this noise reduction can be understood with the
following arguments. With the conventional filtered backprojection algorithm, the fundamental datum is a
chord—a line joining the two detector elements that simultaneously observe 511 keV photons. The filtered
backprojection algorithm reconstructs an image by backprojecting—incrementing each pixel that lies on that
chord by an amount proportional to the number of counts measured in that chord. Placing activity in every pixel
along that chord introduces some blurring, but this blurring is removed (modulo statistical noise) by filtering the
data before it is backprojected. The time difference corresponds to a position along the chord; measurement error
implies a significant uncertainty in this position. Therefore, not every pixel along that chord is incremented by
the same amount when backprojecting. Each pixel is incremented by an amount proportional to the probability
(given the measured time difference and the timing resolution) that the annihilation occurred at that pixel, as
shown in Fig. 5
Fig. 5 Time of Flight and advantages in terms of SNR for different sizes and different timing resolution
With non-TOF reconstruction, coincident events measured in a single chord contribute to all of the image pixels
along that chord, not just the pixel from which the source truly originated. The reconstruction filter removes the
mean contribution to other pixels, but statistical fluctuations in the measurement data cannot be removed and
contribute to noise in all the pixels. With TOF reconstruction, coincident events contribute only to those pixels
that are near (i.e., are within a distance consistent with the timing resolution) the correct pixel, therefore the
statistical fluctuations from the measurement data contribute to a much smaller number of image pixels. When
TOF reconstruction is used, the noise due to randoms continues to diminish as the timing resolution improves.
The effective coincidence window width for computing the noise due to random events (when TOF
reconstruction is used) is the TOF measurement resolution, even though the hardware coincidence window is 4
ns. A lower limit on the improvement from TOF reconstruction is easily computed. If we assume that the
effective source diameters for the true, random, and scatter events are all equal to the phantom diameter (this is
accurate for the trues but underestimates the improvement in randoms and scatter), the variance reduction
factors f = D/x = 2D/ct) for the trues, the scatters, and the randoms are equal (i.e.). As all the variances
decrease by the same factor, the noise variance of the final image also decreases by this factor f. In the case of
non TOF, nconv= D/d. All volume elements that contain activity along the same Line Of Response (LOR)
contribute to the same projection data. In reconstruction, each detected element is evenly back-projected in alla
image elements along the LOR, not only the image element where it was originated. All n elements contribute to
the noise in each image elements In the case of TOF, nTOF = Δx/d in reconstruction each event is back-projected
only in the position associated to such TOF information and into few elements adiacent to it, with a weight given
by a TOF kernel or probability function of width Δx (localization uncertainty related to Δt. Δx=cΔt/2). Using the
TOF capability not only allows increasing the SNR/NECR and lowering the scanning time(5) but also improves
the angular coverage of the detector system. Simulations indicate that TOF information can provide accurate
reconstructed PET images for a limited angle coverage geometry without detector rotation. As angular coverage
decreases, better timing resolution is needed to produce almost artifact-free images. A 2/3 ring scanner with
timing resolution of ≤ 600ps gives lesion CRC values similar to a conventional Full Ring Non-TOF scanner.
With a timing resolution of 300 ps or better a 2/3 ring scanner has similar SNR values to a conventional Full
Ring Non-TOF scanner (6)
6
Preliminary studies
Simulations
We have used Monte Carlo simulations to predict the performance of an initial prostate probe design in
conjunction with an external ring PET system. Simulations were performed using an anthropomorphic phantom
of the pelvic region and the GEANT4 Monte Carlo code. The internal probe was modeled as two stacked 10 x
40 element arrays of 1mm x 1mm x 3mm thick LSO crystals for a total thickness of 6mm, transverse resolution
of 1mm, and depth resolution of 3mm. The LSO was assumed to be coupled to an appropriate photodetector
such as silicon photomultipliers. The external ring scanner modeled for this study consisted of 31 rings each
80cm in diameter with 500 4.2mm x 4.2mm x 30mm thick BGO crystals. Penelope models were used in
GEANT4 to model the effects of positron range for F-18. A Zubal phantom was incorporated as the
anthropomorphic phantom.
14
Hi-Res Noise Advantage
12
10
8
6
4
2
Figure 6b . The
Reconstructed Resolution (mm FWHM)
conventional PET
scanner embedding
Fig. 6 c. Noise advantage of PET with ~12%
1.5mm intrinsic resolution data added to 4mm
the pelvic region
resolution data over 4mm resolution data alone. Curve
containing the PET
is ratio of standard deviations at central point in
imaging probe shown
reconstructed images for large disk source.
in Fig. 6.
Figure 6a shows a slice of the pelvic region and a stack of two 3mm-thick LSO-APD modules. The front side of
the LSO module is 20mm from the center of the prostate. The pelvic region containing the LSO-APD modules
is embedded in a conventional PET scanner. Figure 6b shows the pelvis region (containing “probe”) within an
external ring PET scanner. (Fig. 6c shows the advantage of the probe fie 12% 1.5 mm intrinsic resolution). A
1mm diameter sphere was placed near the center of the prostate. F-18 positrons having the appropriate energy
distribution (endpoint energy: 633.5 keV) were emitted from the sphere isotropically. Coincidences between the
imaging probe and the conventional PET scanner were collected and lines-of-response (LORs) were
backprojected through a source image plane parallel to and 20 mm in front of the face of the imaging probe. The
FWHM for this simple “backprojection” is about 1mm—considerably better than the resolution attainable with
the external ring alone. The detection efficiency at the same point in the prostate was estimated for the combined
PET/prostate probe imaging system and for the conventional PET scanner alone. Estimated efficiency for this
6mm thick detector is shown in Fig. 4 b The center line (stars) shows efficiency of conventional PET without
the imaging probe. The bottom line shows the detection efficiency of the PET scanner when the imaging probe is
inserted in the phantom. Efficiency of external ring coincidences decreases slightly because some are captured
by the probe. The red line shows the total detection efficiency of the PET scanner and the imaging probe.
Overall efficiency increases by ~12% for the 31-ring (~13cm axial extent) PET ring. While this seems modest,
one must remember that the prostate detector in this case is only 6 mm thick x 10 mm wide. It is likely that
detectors at least 10 mm thick could be used which would correspondingly increase efficiency. Also, it is
important to note that the 12% additional data has high intrinsic resolution. Resolution, efficiency, and
reconstructed image noise trade off in an extremely non-linear manner. The key question is whether such a small
fraction of high resolution information will significantly improve images. It is known that with appropriate
scanner modeling, “resolution recovery” is possible. Let’s consider the following scenario: assume that a PET
scanner with intrinsic resolution of 4 mm FWHM is augmented by higher resolution data having 1.5mm FWHM
Figure 6. a pelvic
region. LSO-APD
modules and the
prostate are shown.
0
1
2
3
4
5
6
7
8
9
10
7
intrinsic resolution. Suppose further that the relative fractions of coincidence events are the same as those
shown in Fig. 4b for the “31-ring” external PET and that these include the fact that the prostate probe reduces to
an extent the sensitivity of the external ring. Noise in the reconstructed images at any desired spatial resolution
can be quantified and used as a reference for performance. Results of that calculation are presented in Fig. 6c
where the ratio of the standard deviations for a point at the center of the reconstructed images for the two
systems are plotted as a function of desired spatial resolution in the reconstruction. The relative advantage of the
small amount of high resolution information (~12%) is marginal when the desired resolution is worse than the
intrinsic scanner resolution of 4mm. The ratio is slightly more than unity reflecting only the small increase in
efficiency. However, as the desired spatial resolution approaches and gets better than the intrinsic 4mm
resolution of the external ring, the performance advantage—even with the addition of just a small amount of
high resolution data—increases rapidly.
SiPM choice and characterization
The choice of photodetectors is mandatory. Silicon Photomultipliers (SiPMs) are compact, insensitive to
magnetic field and good for timing. They are a new type of photon counting device made up of multiple
avalanche photodiodes (APD) pixels operated in Geiger mode (14). They have much higher gain (10 5-106) than
standard APDs. The advantages of SiPMs are their very small size, and ability to operate in strong magnetic
fields. Thus, permitting their use in multimodality (PET/MRI, SPECT/MRI) imagers, supplanting the use of
APDs in this application. The application of SiPMs in PET imposes the additional requirement of excellent
timing performance. This is especially true in the light of recent research showing the potentially tremendous
improvement in image quality when time of flight (TOF) information can be incorporated in the image
reconstruction (8). On one hand, the single photon transit time spread of SiPMs is reported to be relatively small
(in the order of 100 ps – 200 ps) , yet, on the other hand, the timing performance is hampered by the large
capacitance of these sensors which results in a smaller rise time of the electronic signals making them more
susceptible to electronic noise. All these parameters have to be taken into account for the choice of SiPM and for
the crucial front end part of the electron ics. It has been shown that very good timing resolution can be achieved
(5). Characterization of the SiPM from different Companies started. Three typical spectra were acquired using
the same reverse bias voltage (32.0V), but the detector thermostated at different temperatures: a) 13.5C, b)
24.3C, c) 41.5C. The Gain reduction with temperature is due to the variation of breakdown voltage with
temperature (reference IEEE2011). The necessity of good timing resolution requires low and stable temperature
to lower the noise and triggering on few photoelectrons. For this reason a cooling system, with feedback on
SiPM power supply will be needed.
The scintillators
LYSO (or LSO) are the reference choice for the scintillators for PET, possibly doped with Ca to optimize the
timing performance. LaBr3(Ce) will be considered because of its very good light yield and timing
characteristics. Simulations (GEANT4) performed show the energy resolution that can be obtained with LSO
(compared to the measured one) and LaBr3(Ce) advantage in energy resolution obtained using LaBr3(Ce). The
drawback with this scintillator is its hygroscopicity, so the need to be encapsulated with glass and the difficulty
in building pixellated arrays. The stopping power and the photon fractio are the other issues. In practice one
would loose on one hand what would get on the other hand (see reference Moses Dresden, Lyon, Baia, etc). Full
simulation and careful measurements will have to be done for the final choice.
8
QuickTim eﾪ e un
decom pr essor e
sono necessar i per visualizzar e quest'im m agine.
Fig. 8 a. Simulated energy resolution LaBr3(Ce) (15%). b. measured energy resolution LYSO (16%).
c. Simulated energy resolution (LYSO, 17%)
Readout electronics
The readout electronics is very demanding both because it has to be very compact and because of the very good
timing resolution needed. The research work will proceed in three steps:
a. A discrete electronics system with VME modules and dedicated preamplifiers have been used for preliminary
timing measurements with finger scintillators (LYSO 3 x 3 x 5 mm3). It was showed that the design resolution is
obtainable. The LYSO finger scintillator were coupled (1 to 1 to optimize the light collection) to Hamamatsu 3 x
3 mm2 pixel SiPM. Dedicated preamplifiers were used. A timing resolution as good as 350 ps was obtained. with
25 mm microcells. Due to the higher PDE obtainable, this projects out a timing resolution of ~ 250 ps once the
SiPM
with
50

microcells
will
be
used.
QuickTim eﾪ e un
decom pr essor e
sono necessar i per visualizzar e quest'im m agine.
Fig.9 a . Measured timing resolution of finger counters; LYSO 3 x 3 x 10 mm3 coupled (1 to 1) to Hamamatsu
SiPM (microcells of 25 )
b. A dedicated compact electronic system using off-the-shelf components is under design, focusing on existing
components. This work already started in the framework of the Italian collaboration that focused on main
components, yet designed and produced for CERN LHC experiments: a very fast preamplifier-discriminator
9
with low input impedance and Time Over Threshold (TOT) capability, a Time to Digital Converter (TDC) with
dual edge measurement capability and resolution ≤ 100 ps. For the preamplifier-discriminator the choice is
NINO: an 8 channels wide bandwidth fully differential preamplifier-discriminator with 40 Ω input impedance.
For the TDC the HPTDC chip is available. It measures the timing of both edges of the input pulses with
resolution selectable down to 25 ps. The system will be composed by a stack of three boards: the detector board
hosts the SiPM array; the front-end board hosts the NINO chips (128 channels, 16 NINO); the TDC board hosts
the dedicated chips (4 HPTDC, 32 channels each) the Control Board which implements the coincidence logic,
readout control and the link with data acquisition computer.
c. The system described e in point 2 will be upgraded with the insertion of a dedicated ASIC with a dedicated
front end, crucial for the characteristics of the SiPM and the timing resolution needed and a fats TDC derived
form the expertise gained recently for detectors for CERN experiments and focused on this particular
application. The ASIC will be designed in the framework of the most advanced technologies (130 nm).
The ASIC (Ranieri, Loddo, Rivetti)
To be able to handle many channels an ASIC has to be built. We will use CMOS (130 nm) mixed analog-digital
technology. Parameters to be considered are high density integration, 64-128 channels, low power/channel (few
mW, flexibility (gain and threshold tuning), energy resolution (8 bit ADC ), digital output (LVDS), minimizing
time jitter and time walk. The front end part of this ASIC has been designed and built. A 8-channel ASIC for
SiPM readout has been developed for TOF-PET applications. This release contains only the analog chain, in
order to fully characterize the detector and the very front-end part, while next development will be the
integration of a Time-to_Digital converter and serial readout to fully exploit the excellent timing features of
SiPM required in TOF applications. Each channel is based on input current buffer followed by a current
discriminator for fast event tagging and LVDS driver. Each channel is also equipped with a 8-bit voltage DAC
for the fine tuning of the SiPM bias voltage and 6-bit current DAC for offset compensation, while a single 8-bit
current DAC provides the thresholds to all discriminators. A simple standard cell digital section implements a
SPI interface to remotely control the different DACs. Finally, the logical "OR" of the outputs is provided for
- Technology: 0.13um standard CMOS
- Power supply: 1.2 V - 3.3V
- Power consumption:  1.0mA/channel
- Package: CQFP 120 pins; Dynamic range: about 1000 pC
- Parallel output
- Advantages: Low impedance, large bandwidth input
stage, fast current discriminator, internal bias reference,
parallel readout.
Limitations: no internal ADC, noi internal TDC, 50fC
equivalent input noise charge
triggering purpose.
Next step will be the integration of a TDC with the expertise coming from the high energy CERN experiments.
PET – MRI compatibility
10
The value of PET lies in its high-sensitivity tracking of biomarkers in vivo but it lacks resolving morphology.
MRI has lower sensitivity, but produces high soft-tissue contrast and provides spectroscopic information and
functional MRI (fMRI). It has been shown that both modalities preserve their functionality, even when operated
isochronously (reference Pichler Nature). In this reference paper the average magnetic resonance image SNR
and homogeneity showed small degradations, on the order of 6% and o1%, respectively, when the PET insert
was built into the magnet. This degradation is mainly caused by conductive PET detector materials and is
therefore more predominant at the end of the PET FOV, where the connectors for the PET signals are arranged
on the detector board. These degradations, however, are not fundamentally relevant for MRI applications in
living subjects, where the SNR can easily change in the range of this magnitude depending on coil-load
variations. MRI geometrical uniformity and spatial resolution of the investigated magnetic resonance sequences
have not shown substantial degradations resulting from the presence of the PET insert. The paper refers to a
small animal imaging PET MRI scanner with 7 Tesla. The high filed scanners used in prostate MRI imaging
range from 1 Tesla to 3 Tesla so the mutual interference of the two scanners will be certainly lower. Moreover in
this project we will compare the performances of low field MRI with high filed ones. The mutual interference
issue, if any, will be even less important
MRI imaging of the prostate (Silvia Capuani, Gemelli..)
Conventional MRI of the prostate relies on morphologic changes within the prostate to define the presence and
extent of cancer. Currently, the prostate is imaged by MRI using an endorectal coil in combination with four
external coils (pelvic phased-array). Endorectal coil MRI provides higher spatial and contrast resolution on
prostate zonal anatomy than TRUS or CT. T2-weighted MRI has shown high sensitivity in prostate cancer
localization (97%), although performance varies with the patient population studied. It is also not sensitive in
detecting cancer in regions other than the peripheral zone of the prostate. Functional MRI imaging techniques,
such as MR spectroscopy (MRS), diffusion-weighted MRI (DWI), and dynamic contrast-enhanced MRI
(DCEMRI), have been investigated for potential to complement T2-weighted MRI in improving prostate cancer
localization. Endorectal-coil T2- weighted MRI shows decreased signal intensity for prostate cancer relative to
normal peripheral zone tissue but is less sensitive at detecting cancer in other prostatic zones. Also, low signal
intensity is not specific for prostate cancer because benign conditions such as prostatitis, hemorrhage, and
therapeutic effects also have a similar appearance at MRI. Functional MRI techniques, such as MRS, DWI, and
DCE-MRI, have been investigated for their potential to complement morphologic T2-weighted MRI in
improving prostate cancer localization. MRS is an FDA-cleared technology for noninvasively measuring
metabolic activity on the basis of relative concentrations of metabolites in tissues. In the prostate, cancer tissue
shows a decreased concentration of citrate but an elevated concentration of choline relative to normal prostate
tissue. Studies showed that adding metabolic information obtained from MRS to morphologic inforation
obtained by MRI improved cancer localization and predicted prostate cancer aggressiveness. But a recently
completed multiinstitutional study concluded that there was no incremental benefit for MRI–MRS compared
with MRI alone in tumor.
The low field option (Silvia Capuani, Gemelli..)
An inherent limitation of low-field magnets for diagnostic MR imaging, the reduced signal-to-noise ratio
attainable compared with that attainable at higher field strengths can be overcome with the use of MR contrast
agents that possess higher r1 relaxivity compared with conventional MR contrast agents. Low field MRI scanner
with contrast enhanced technique showed to perform as well as high field one’s and even better in some cases
(references breast). Relaxivity in human blood plasma at 0.2 T (10.9 l mmol–1 s–1) is roughly twice that of
gadopentetate dimeglumine (Gd- DTPA) (5.7 l mmol–1 s–1). Numerous studies have shown not only that
gadobenate dimeglumine is effective for breast MR imaging but also that it offers improve diagnostic
performance compared with gadopentetate dimeglumine administered at equivalent doses. Calabrese et al
(reference Sardanelli) showed that all cancers in their study were correctli diagnosed on low field MR imaging
11
using high relaxivity contrast angent. Similar results were obtained by Paarko et al. with high dose od
conventional gadolinum agent.
In our case, prostate cancer……….
Imaging agents: present situation and prospectives (M. Pomper)
Animal studies (Hildeshein University)
12
Clinical studies (Hildeshein University)
13
Clinical studies (University Agostino Gemelli – Rome)
14
Summary and conclusions (Franco + everybody)
15
The scope MPROPECAM project is building a multimodal detector (PET TOF and SPECT MRI) aimed at
solving the prostate cancer diagnosis problem. The probe, while challenging has unprecedented performances
putting together anatomic and metabolic information that will be of fundamental importance in the diagnosis
and management of prostate ca.
References
1.G. Kellof et al. Challenges in Clincal Prostate Cancer: Role of Imagin. AJR:192, June 2009, pg. 1455
M. Pomper et al. New Agents and Techniques for Imaging Prostate Cancer. Focus on molecular Imaging. To be
published in JNM
2. S.S Huh, N. Clinthorne, W.L. Rogers. Investigation of an internal PET probe for prostate cancer Imaging,
NIMA 579 (2007 339-343
3. Tay YC et al. Initial study of an asymmetric PET System dedicated to breast cancer imaging. IEEE TNS 53:
121-126
4. W Moses, Dedicated Nuclera Medical Instrumentation for Prostate cancer. Proceedings of the Rome
workshop….
5. N. Clinthorne et al. Multi-resolution image reconstruction for high resolution small anima PET device. IEEE
2003, Nucl Symp. Conf. Rec. 3: 1997-2001
6. Surti S. et al, Breast Radiology workshop, Dresdeen 2008)
7. J.S Karp et al. Benefit of Time –of-Flight in PET: Experimental and Clinical Resukts, JNM, February 20,
2008/ nume.107.044834
8. W.W. Moses. Time of flight revisited, IEEE TNS Vol 50, N. 5 October 2003, 1325
9. M. Conti et al. Comparison of fast scintillators with TOF PET potential. IEEE TNS Vol 56., N. 3, June 2009
10. Szczęśniak tet al“Timing Resolution and Decay Time of LSO Crystals Co-doped with Calcium”, Transaction
on Nuclear Science, to appear
16