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: MUPROBEPCa
State of art and objectives
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 (1). Tools
for diagnosis and follow up are not very good. The gold standard is PSA -> DRE - > Biopsy that is blind,
because of seeing only the organ, not the lesions. MRI and MRS have a role but lack sensitivity and or
specificity. Four main clinical challenges exist in prostate cancer treatment and management—diagnostic
accuracy; risk stratification, initial staging, active surveillance, and focal therapy (1). So better understanding
and managing this disease is needed. The issue is how new anatomic, functional, and molecular imaging
techniques might help in early diagnosis and add to more accurate characterization of disease, better staging
and evaluation of response to therapy. Imaging has played a relatively minor role in the management of
prostate cancer up to now. 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 (ultrasound, MRI, CT, nuclear medicine), cannot detect early disease. They provide limited
information for disease staging. Several emerging techniques are under investigation, 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. MUPROPESCAM project propose a new approach: 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 (and possibly US) scans. MRI is the ideal complement to radionuclide imaging
techniques for soft tissues and can have adjunct value by using dynamics of contrast enhancement. 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. MUPROPEMSCAM project is aimed at developing an endorectal PET-TOF and SPECT MRI (and
possibly US) 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. The probe will
have also an ultrasound probe for tracking purposes and guiding biopsy. Performance is dominated by the
endorectal detector with improvements in both spatial resolution and efficiency (reference Clinthorne). 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. Drawbacks of MRI exams comes from high cost and lack of
appropriate MRI biopsy guiding methods. A low-field open MRI unit in multimodality with PET and SPECT
and US 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 images with a low
field one’s will be performed. Using low field scanner would decrease the cost dramatically, facilitating the
diffusion of the proposed multimodal detector in the clinical practice. 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 ones. Low field MRI scanner with contrast enhanced technique
showed to perform as well as high field one’s and even better in some cases (9,10) minimizing 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, characteristic of the project is the search for new radiotracers
SPECT and PET with high target-no target ratio, not existing yet. Tests on big animals and clinical trials will
be performed. 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
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
geometries (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 better spatial resolution (12). Full simulation will be needed to
optimized the design this setup. From Fig. 1.a and d 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). Tapered pixel
scintillators would be used in this case (see later on and refernce 20). 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 paragraph for the rationale)
d
Fig. 1. a . 2 external detectors layout. B. the standard single photon detector layout . c. our proposal for the
single photon layout. d. the need to make at least 2 scans for complete prostate sampling.
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 is 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, 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 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 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. . In the first a high-resolution endorectal detector works
in conjunction with a planar external detector (or two (b), see later). The second (c) employs a similar highresolution 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.
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.
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 can be mitigated by using 2 detectors profiting of the TOF capability (see later and (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 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 detector 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 highresolution, 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,
vedi Cherry)). 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 coworkers for high resolution breast imaging among other application (3,5). 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 line-of-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
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:
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).
66
20000
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Events Detected
Resolution (mm FWHM)
Resolution (mm FWHM)
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3 mm
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BGO+LSO Probe 1
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Distance
Distancefrom
fromprobe
probeface
face(cm)
(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 other members of the team (references).
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=2R1R2DT, 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 DT 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 DT. 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 = T 2/ (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 Dx=c/DT where Dx is the position error, c is the speed of
light and DT 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/Dx =
2D/cDT 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 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/Dx = 2D/cDt) 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 NonTOF scanner (6).
Optimization of the detector for the timing resolution depends on many parameters like light yield of the
scintillators, light collection, PDE of SiPM, time jitter, front end electronics, ASIC etc as shown in the
formula (reference M. Conti, Baia delle Zagare)
Using TOF will result in a sort of amplification factors for sensitivity that will depend strongly on the
charachtristics of the scintillator (reference M. Conti, Baia delle Zagare)
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. Some result is shown in Fig. 6. Fig. 6c shows the
advantage of the probe fie 12% 1.5 mm intrinsic resolution. The front side of the LSO module is 20mm from
the center of the prostate. The pelvic region containing the LSO-SiPM modules is embedded in a
conventional PET scanner.
Fig. 6 a .The pelvis region containing the probe within
an external PET scanner. b. A slice of the pelvic region
and a stack of 2 3mm-thick LSO-SIPM modules. c.
Noise advantage of PET with ~12% 1.5mm intrinsic
resolution data added to 4mm resolution data over
4mm resolution data alone. Curve is ratio of standard
deviations at central point in reconstructed images for
large disk source.
14
Hi-Res Noise Advantage
12
10
8
6
4
2
0
1
2
3
4
5
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8
9
10
Reconstructed Resolution (mm FWHM)
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. This seems modest, but the prostate detector in this case is only 6 mm thick x 10 mm
wide. Detectors at least 10 mm thick will be used which would correspondingly increase efficiency (see Fig.
4b. yellow point). 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 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 (SiPM) 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 scintillator arrays
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 (17). Full simulation and careful
measurements will have to be done for the final choice. Scintillator arrays will be designed and built with
coupling one – to one (perfect matching) to SiPM arrays, in order to optimize the light collection, crucial
parameter for timing resolution. Scintillators for the single photon probe will be chosen between CsI(Na),
CsI(Tl) and NaI(Tl) for their characteristics of light yiel, efficiency and photon fractio. As mentioned before
the scintillator pixels will be inserted in the collimator holes for ~ 1/3 of the thickness. Full simulation of the
detector will be needed to optimize this layout. The proposed fan beam collimator will require trapezoidal
(tapered) scintillators. The SiPM arrays will have to match perfectly the collimator design.
Fig. 8 a. Simulated energy resolution LaBr3(Ce) (15%). b. measured energy resolution LYSO (16%). c.
Simulated energy resolution (LYSO, 17%)
Readout electronics. Preliminary results on timing resolution
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, already started, will proceed in three steps:
a. Use of a discrete electronics system with VME modules and dedicated preamplifiers 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 scintillators 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 has been measured with 25 mm microcells devices. Due to the higher PDE obtainable, this extrapolates to
a timing resolution of ~ 250 ps once the SiPM with 50 mm microcells will be used. Detailed measurements
are now in progress.
Fig.9 Measured timing resolution of finger counters; LYSO 3 x 3 x 10 mm3 coupled (1 to 1) to Hamamatsu
SiPM (microcells of 25 mm). System setup on the left and results on the right.
b. A dedicated compact electronic system using off-the-shelf components has been designed and now is in
fabrication. This modular system uses existing components developed in the framework of CERN-LHC
experiments: a very fast front-end preamplifier-discriminator (reference Jarron) with low input impedance
and Time Over Threshold (TOT) capability (a chip called NINO), and a Time to Digital Converter (TDC)
with dual edge measurement capability and resolution ≤ 100 ps (a chip called HPTDC). NINO is an 8
channels device, with low input impedance (down to 40 Ω) wide bandwidth fully differential preamplifierdiscriminator. HPTDC can measures the timing of both edges of the input pulses with resolution selectable
down to 25 ps.
This compact system is composed by a stack of 4 boards:
- the detector board will host an array of 128 SiPM, its bias voltage generator and a temperature monitor;
- the front-end board houses 16 NINO each one with programmable threshold;
- the TDC board has 4 HPTDC chips assembled on it;
- the Control board handles configuration, coincidence trigger, data acquisition and communication with PC
using USB 2.0 protocol.
c.The system described in point b) will be upgraded with the insertion of a dedicated ASIC whose front end
performances are matched to the peculiar characteristics of the SiPM array. A fast TDC block derived form
the expertise gained recently form detectors for CERN experiments (referenze Rivetti) and focused on this
particular application will be added to obtain the requested timing resolution. The following paragraph will
give more details on this item.
The ASIC
In order to deal with many channels in limited space, a custom multi-channel ASIC has to be built. The chip
will be designed in a standard CMOS 130 nm process and will integrates both analogue and digital sections
of the circuitry required to amplify the SiPM signal and make precise time measurement. The main
constraints to be considered are: high density integration (64-128 channels), very low power/channel (few
mW), flexibility (detector gain and threshold tuning), minimizing time jitter and time walk. The front end
part of this ASIC has been already designed and built: a 8-channel ASIC for SiPM readout has been
developed for TOF-PET applications. This release contains only the input current buffer, 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 with the expertise coming from the high energy CERN experiments 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 triggering purpose. Summarizing we are using 130
nm standard cmos technology, power supply 1.2 V – 3.3 V, power consumption of ~ 1 mA/channel, package
CQFP 120 pins; dynamic range ~ 1000 pC, parallel output. The adavantages low impedance, large bandwidth
input stage, fast current discriminator, internal bias reference, parallel readout., limitations (at the moment)
no internal ADC, no internal TDC, 50fC equivalent input noise charge
5 righe Rivetti
PET – MRI compatibility
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 (18). 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 will be even less important.
MRI imaging of the prostate (Silvia Capuani, Bonomo, Bassi)
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 contrastenhanced 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 information obtained by MRI improved cancer localization
and predicted prostate cancer aggressiveness. But a recently completed multi-institutional study concluded
that there was no incremental benefit for MRI–MRS compared with MRI alone in tumor……….
The low field option (Silvia Capuani, Bonomo, Bassi)
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 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……….
The layout of the proposed detector
A possible layout of the system is shown in Fig. 10. This is valid both for the single photon and the PET
probe. For the single
photon probe a tungsten
shielding will be needed.
Fig. 10. The layout of the system with two option for DOI optimization, double side readout of one
scintillator array 10 mm thick and double side array of 2 5 mm scintillator array 5 mm thick.
The container will be a cylinder with external diameter of ~ 35 mm. The materials will be selected as non
magnetic to be compatible with the magnetic field of the MRI scanner and for the right thermal conductivity
(carbon is a good candidate). Copper shielding will be needed to minimize eddy currents. The detector will
be cooled by air and monitored continuously to be held at ~ 20-22 C. An ultrasound sensor will be installed
in front of the detector for tracking/alignement purposes and also for imaging (in particular for guiding
biopsy). In the case of the PET probe the thickness of the scintillator array will be of the order of 10 mm to
have reasonable detection efficiency. The dimension of the detector will be ~ 25 x 50 mm2. The width will be
dictated by the thickness of the SiPM arrays and on DOI optimization needed. In fact simulation and
measurements will show if the optimization of both the timing resolution and the Depth Of Interaction (DOI)
resolution will need an unique scintillator array (10 mm thick) with double side readout (19) or 2 sheets (5
mm thick each) each one with double side readout of SiPM. An optimization is needed for the treatment of
the surfaces of the scintillators good for DOI and timing. It will be probably an intermediate case between
fine polished surface, that favorize the timing resolution minimizing the path of the light and the rough (as
cut) treatment to minimize the attenuation length for DOI. As for the single photon case the geometry of the
PET probe has to be carefully studied to minimize the scan time, to avoid angular undersampling and to
allow good matching with the MRI scanner. Geometries like the ones shown in Fig. 11 a . b. will be
carefully evaluated. The layout of Fig. 1 a would avoid angular coverage undersampling without sacrifying
the detector efficiency with respect to the rectangular shaped detector. Tapered pixellated scintillators would
be used in this case. They showed good performance (20). Fig. 1 b would have not only complete angular
coverage but also the possibility to use an endorectal coil. The price to pay is a reduced efficiency. Both
geometry will be carefully evaluated for the project.
Fig. 1 a. layout of the PET probe with curved geometry. The whole prostate is sampled with one scan. b. the
layout with 2 curved shaped detectors. An internal MRI coil can be used with advantages in terms of
sensitivity.
Imaging agents: present situation and prospectives (M. Pomper)
Animal studies and clinical trials (M. Hofmann, Hildeshein University)
Clinical trials (University Gemelli, Rome)
A 2-phase clinical study will be carried out in a single academic institution. Thirty patients with abnormal
serumPSA (Prostate Specific Antigen) and abnormal free/total PSA ratio, and after digital rectal examination
will be included in the study. This patient population is routinely submitted to prostate ‘random’ sextant
biopsy to rule out the presence of prostate cancer. All the patients will undergo prior imaging evaluation,
such as transrectal prostate ultrasonography, gadolinium enhanced prostate morphological MRI as well as
functional MRI (MR spectroscopy, diffusion weighted MRI and dynamic contrast –enhanced MRI) and
choline PET/CT scan. Then the patients will undergo high sensitivity tracking PET and MRI comparative
evaluation. Thereafter, the patients will undergo prostate random biopsies, according to the European
Urological Association 2010 guidelines. Any abnormal imaging area will be additionally biopsied according
to the prostate gland map retieved either under the new protocol imaging, or the conventional imaging. Any
side effects related to the imaging procedure will be recorded. The study will be submitted to the local
Ethical Committee and carried out under Good Clinical Practice (GCP) rules. The first interim analysis will
be performed after 15 patients’ evaluation. The second subset of patient will be enrolled only if at lest a 10%
advantage in diagnostic accuracy between conventional and new protocol imaging will be demonstrated.
(Franco + everybody)
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.
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18. B. Pichler, Nature
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