Università di Pisa
Dipartimento di Fisica E. Fermi
Laurea Magistrale in Fisica
Optimization and performance evaluation of the
SiPM based PET detectors of TRIMAGE brain
scanner
Candidato:
Tania Cortopassi
Relatore:
Prof. Nicola Belcari
Academic year 2017-2018
Contents
Introduction 7
1 Multimodality imaging 9
1.1 Positron Emission Tomography (PET) . . . 9
1.2 PET physical theory . . . 11
1.2.1 Beta decay . . . 11
1.2.2 Photon interaction with matter . . . 13
1.3 Coincidence events . . . 15
1.4 Time of Flight PET . . . 17
1.5 Multimodality imaging . . . 19
1.5.1 Possible design of PET/MRI systems . . . 20
1.6 State of the Art . . . 22
1.6.1 Mindview . . . 23
1.6.2 Other detector prototypes . . . 23
1.7 TRIMAGE multimodality imaging . . . 24
2 Scintillation Detectors 27 2.1 Scintillators . . . 27 2.1.1 Scintillator properties . . . 28 2.1.2 Organic scintillators . . . 29 2.1.3 Inorganic scintillators . . . 30 2.2 Depth of Interaction . . . 32 2.3 Photodetectors . . . 33 2.3.1 Photomultiplier Tube . . . 34 2.3.2 Semiconductor photodetector . . . 35
2.4 Read out electronics . . . 40
2.4.1 Leading edge timing . . . 42
2.4.2 Constant fraction discrimination . . . 42
3 TRIMAGE PET detector 45
3.1 TRIMAGE PET system . . . 46
3.1.1 The Tile detector . . . 49
3.2 Tile calibration methods . . . 51
4 Optical coupling and detector assembly 55 4.1 Optical coupling in TRIMAGE PET detector . . . 56
4.2 Optical coupling materials . . . 57
4.3 Bonding procedures . . . 57
4.4 Bonding procedure evaluation . . . 59
4.5 Measurements and Results . . . 60
4.6 Conclusions and Future work . . . 63
5 Performance evaluation 65 5.1 Preliminary analyses . . . 65
5.2 Methods . . . 68
5.2.1 Experimental set up . . . 68
5.2.2 Bias voltage analysis . . . 68
5.2.3 Energy resolution algorithm . . . 69
5.3 Experimental results . . . 70
5.3.1 Flood Map analysis . . . 70
5.3.2 Energy resolution . . . 74
5.3.3 Number of events detected in each layer . . . 76
5.3.4 Coincidence time resolution . . . 77
5.4 Working point summary . . . 78
5.5 Conclusions and future work . . . 79
List of Figures
1.1 Positron spectrum in 18F decay jan˙monte˙2005 . . . 12
1.2 Compton Scattering . . . 14
1.3 Photoelectric effect . . . 14
1.4 electronic collimation . . . 16
1.5 Type of events in PET . . . 17
1.6 Conventional and TOF reconstruction . . . 18
1.7 Three concepts for integration of PET and MRI: a brain PET scanner inserted in a whole body MRI (top), a sequential PET MRI scanner for whole body imaging (middle) and a fully integrated simultaneous PET MRI scanner for patient imaging (bottom). . . 22
1.8 A scheme of the TRIMAGE scanner . . . 26
2.1 Scheme of scintillator mechanism in inorganic crystals. On left a pure scintillator crystal, on right doped crystal . . . 31
2.2 Perceived and actual line of response (LOR) in PET imaging due to not DOI correction. Figure from bolus˙petmri:˙2009 32 2.3 Different approaches to design systems with DOI encoding . . 33
2.4 Scheme of the PMT signal production mechanism when cou-pled with a scintillator crystal . . . 34
2.5 Avalanche-photodiode structure. Image from Hamamatsu pho-tonics . . . 36
2.6 SiPM structure lecomte˙novel˙2009 . . . 36
2.7 Semiconductor photodetector characteristic curve . . . 37
2.8 Effect of walk (up) and jitter (down) on timing discriminators using leading edge method . . . 42
2.9 Comparison of threshold triggering (left) used in leading edge timing and constant fraction triggering (right) . . . 43 3.1 Scheme of full TRIMAGE PET scanner (left) and segment
3.2 Picture of an ASIC board that host 4 TRIROC ASIC. Image
from sportelli˙single-mode˙2015-1 . . . 47
3.3 Schematic diagram of the acquisition pipeline from the ASICs to the host PC. Image from sportelli˙single-mode˙2015-1 . 48 3.4 Scheme of light pattern produced by a photon interacting in top bottom layer (a) and top layer(b). Image from camarlinghi˙evaluation˙2016 49 3.5 (a) LYSO staggered crystal matrix and (b) SiPM matrix. Im-age from sportelli˙single-mode˙2015-1 . . . 50
3.6 The main stages of the calibration software . . . 53
4.1 Refraction of light at the interface between two media of differ-ent refractive indices, with n1 > n2 depending on the incident angle. . . 56
4.2 A step in the glueing procedure (left) and the result (right) . . 58
4.3 Examples of artefacts in flood map of double-side tape tiles . . 61
4.4 An example of a flood map of glued tile with an artefact due to an air bubble . . . 62
4.5 An example of a flood map of glued tile without artefacts . . . 62
5.1 Q-threshold values, evaluated as the minimum of the most energetic channel of an event . . . 66
5.2 Number of activated channels at different T-threshold value. . 67
5.3 The sector used (left) and the experimental set up (right) . . . 68
5.4 Example of light patterns of events from different layers, events from 18F , SiPM bias voltage -30 V . . . . 69
5.5 Example of energy spectra (bottom in red, top in blue) com-puted with the two methods: al left considering all the chan-nels, at right considering only a fixed number . . . 70
5.6 flood maps obtained at different bias voltage . . . 71
5.7 Pixels profiles at -29.5 V . . . 72
5.8 Pixels profiles at -30.7 V . . . 73
5.9 Pixel identification expressed as average of resolvability index for each layer. . . 74
5.10 Energy resolution computed considering all the channels of the event . . . 75
5.11 Energy resolution computed considering a fixed number of channels . . . 75
5.12 Number of events detected from each layer at different voltage 76 5.13 CTR of a pair of tiles as a function of bias voltage . . . 77
List of Tables
1.1 List of common PET tracers . . . 10 1.2 TOF NEC gain for a 30 cm diameter uniform cylinder for
different value of time resolution . . . 19 1.3 Summary of specification and performance of the main brain
PET/MRI scanners . . . 24 2.1 Physical properties of various inorganic scintillators used in
PET detector lecomte˙novel˙2009 . . . 31
2.2 Characteristics of photodetector for PET. Table from lecomte˙novel˙2009. Temperature coefficient express the temperature dependence
of gain. . . 40 3.1 Summary and specification of the whole system . . . 48 3.2 Specification of the component of a Tile . . . 51 4.1 Energy resolution of different tiles bonded with double-side
tape(left) and bi-component glue(right) . . . 63 5.1 Summary of the baselines and optimum bias voltage value for
Introduction
Positron emission tomography (PET) is a very sensitive functional imaging modality, that can provide precious information about functional, metabolic and molecular processes, but it gives almost no anatomical information. To overcome this limitation PET system should be integrated with other iing technique based on different physical principles, like X-ray CT or mag-netic resonance imaging (MRI). The excellent soft-tissue contrast provided by MRI is expected to drastically increase the diagnostic potential of a combined PET/MRI system. Applications of such systems can be found in neuroimag-ing, as well as other fields. Therefore, despite many technological challenges to overcome, the possibility of the development of a PET integrated in a PET/MRI system is a well discuss topic in nuclear imaging instrumentation research. The main challenge in combining PET and MRI is the develop-ment of an MRI-compatible PET detector. Started at the end of 2013 and held by an international consortium, TRIMAGE is a project with the aim to create a trimodal, cost-effective imaging tool consisting of PET/MR/EEG (Positron Emission Tomography/Magnetic Resonance/Electroencephalogra-phy ). The research group at the University of Pisa is responsible of the design, development and optimization of the PET scanner. The detector designed is based on dual layer LYSO staggered scintillator crystals coupled one-to-one to Silicon Photomultiplier (NUV-SiPM) matrices. In this thesis the PET component of the TRIMAGE brain scanner has been characterized. The aim of this work is to define the calibration procedures and evaluate the performance of the TRIMAGE PET detector, measuring different detector features, such as pixel/layer identification capability, energy resolution and coincidence time resolution. This study lead to the optimization of detector final design and to a substantial improvement in early stage detector perfor-mances.
In the first two chapters a theoretical overview is presented. In the first one, PET physical theoretical principles and multimodality clinical motiva-tions as well as the state-of-the-art of PET/MRI detectors are introduced.
In the second chapter the physical mechanism which scintillation detectors are based on are discussed. Scintillation materials, photodetectors and front-end electronics are described.
In the third chapter the TRIMAGE PET brain scanner is presented and the main methods of analysis are introduced.
The fourth and fifth chapters describe the system optimization studies. Dur-ing these studies I analized the assembly procedure and optical couplDur-ing materials, as presented in the fourth chapter. The fifth chapter reports the optimization of ”tile” sub-detector parameters I performed in order to eval-uate the best working point of the detector. This optimization required the implementation of a software to evaluate calibration parameters, such as ASIC board thresholds and SiPM bias voltages.
Chapter 1
Multimodality imaging
The TRIMAGE project started at the end of 2013 with the aim of creating a trimodal, cost-effective imaging tool consisting of PET/MR/EEG. Funded by the EU, the project is run by a consortium of 11 partners, seven from academia and four from the SME environment. It is coordinated by the De-partment of Physics of University of Pisa. A specific task of the research group at the university of Pisa is the design, development and optimization of the PET scannerdel˙guerra˙trimage:˙2018. This chapter is dedicated to the illustration of the physical principles of PET and importance of a multi-modal PET/MR/EEG approach in early diagnosis of different diseases. Fur-thermore, an overview of the technological difficulties related to PET/MRI system is presented as well as the state-of-the-art of brain PET/MRI scan-ners.
1.1
Positron Emission Tomography (PET)
PET imaging is based on the detection of γ photons generated by the an-nihilation of positrons. To produce these photons a radiopharmaceutical labelled with a positron emitter is distributed within a biological target. The emitted positron annihilates with an electron of the surrounding tissue to produce two photons of 511 keV emitted along opposite directions. The detection of most of these photons allows reconstructing the distribution of the activity in the body. This allows to obtain functional information used to study metabolism of an organ and its physiology, as well as pharmacoki-netics of new drugs. Due to this wide range of possible applications, many radionuclides with short half-life have been studied, each one for a specific purpose. The most commons are listed in tab 1.1. The most commonly used is fludeoxyglucose (18F − F DG), an analogue of glucose, used to
de-tect abnormal metabolic activity. This information is widely used in clinical oncology, because metastases and cancer cells are known to express higher uptake of glucose than normal tissue reske˙fdg-pet˙2001. Moreover, since the cellular metabolic function of a tumor changes before it starts growing, PET imaging enables the physician to detect cancer at an early stage, before anatomical changes could be seen on a CT or MRI scan. As a result, FDG-PET is used for diagnosis, staging, and monitoring treatment of cancers. Many other clinical studies of human disease are possible. PET systems are also applied to the study of schizophrenia, addiction, anxiety and depression in the field of psychiatry, blood flow activation and neurochemical processes in the area of cerebral functionality, and dementia, movement disorders and stroke recovery in the field of neurological disease.
Radiotracer Average Ek Half life Effective Range in H2O
(MeV) (min) (mm)
18F 0.242 109.8 0.54
15O 0.735 1.7 2.4
11C 0.385 20.4 0.92
68Ga 0.740 68.3 2.8
Table 1.1: List of common PET tracers
PET systems are usually composed by detector rings, which are arranged cylindrically around the patient, thus outlining the volume that can be in-vestigated (the field of view FOV). Standard PET detectors consist of three main components: a scintillator that converts the 511 keV photons from the positron decay into visible light, a photodetector that transform light to elec-trical signal and a front-end electronic to elaborate it.
The scintillator is usually an inorganic crystal in which the gamma spec-trum photons are converted to visible light (scintillation photons) through photoelectric absorption and Compton effects. The scintillation crystals are coupled to photodetectors: photomultiplier tubes (PMT) in standard PET systems or Silicon photomultiplier (SiPM) in MRI compatible PET like TRIMAGE. SiPMs offer many advantages, including being MRI compatible and really compact. Thus, they’re becoming widely used in multi modality imaging. Their output is a measurable electric signal proportional to the energy deposited in the scintillation crystal. Several materials can be used as scintillators and photodetectors with different properties and advantages that will be investigated in chapter 2.
1.2
PET physical theory
1.2.1
Beta decay
Beta decay is a type of radioactive decay in which a beta particle (an electron or a positron) and a neutrino are emitted from a nucleus. Beta decay is due to weak interaction and is usually referred to as β− if the emitted charged lepton is an electron or β+ if it is a positron. By this process, unstable atoms obtain a more stable ratio of protons to neutrons. This is due to the fact that in β− a neutron is converted in a proton and vice versa in β+ according
to the following processes:
β− : n → p + e−+ ¯νe
β+ : p → n + e++ νe
β+ doesn’t occur with free proton due to the fact that the sum of the masses of the neutron and positron is greater than the mass of the proton. Consequently this reaction is possible only over a threshold of energy. β+
decay can only happen inside nuclei when the daughter nucleus is more stable, thus has a greater binding energy (and therefore a lower total energy) than the mother nucleus. The exponential decay law governs the emission of a radioactive source:
N (t) = N0e−λ(t−t0)
where N0 and N (t) represent the number of nuclei at the reference time (t
= 0 ) and at the instant t (t > 0 ) respectively, while λ is the decay constant of the radioisotope.The number of disintegrations per second is the activity of the radioactive source:
A(t) = −dN
dt = λN (t)
Positrons are emitted with a kinetic energy. Concerning the β spectrum, thus the distribution of energy values for the β particles, is a continuous with a maximum fixed for each parent/daughter nuclide couple. As an example in figure 1.1 is reported the Monte Carlo simulation for positron kinetic energy in18F decay jan˙monte˙2005:
18F →18O + e++ ν
e+ 0.633 MeV
The neutrino rarely interacts with matter and is therefore extremely dif-ficult to detect. The positron instead loses most of its kinetic energy by scat-tering interactions with the surrounding atoms until it reach thermal energy.
Figure 1.1: Positron spectrum in18F decay jan˙monte˙2005
The range of the positron before thermalization depends on its energy and on electron density of the medium. Usually the range of the positrons emit-ted from PET radionuclides is about 1-2 mm. An ideal tomograph should be able to reconstruct the positron emission points. Actually, a tomograph can only detect the annihilation point: the range of the positron separates these two points. This range effect introduces a blurring in the image, re-sulting in a degradation of the spatial resolution guerra˙positron˙nodate. When the positron has reached the thermal equilibrium may either annihilate with an e− or enter into a short-lived bound state with the electron called positronium. The annihilation of the positron emits two γ of 511 keV in (nearly) opposite directions. The co-linearity is fundamental in PET image reconstruction in which the image is reconstructed from the line of response (LOR) of the detector but is only an approximation. Since electron and positron are not at rest in the laboratory frame, the angle of emission is deviated approximately of 0.5◦. This process usually take place in approxi-mately 1.8 ns. The positronium, instead, is a unstable system whom ground state has two possible configuration depending on the relative orientation of the spins:
the singlet state 1S
0 called para-positronioum with antiparallel spins
and a half-life of 0.1 ns. the triplet state 3S
1 called ortho-positronioum in which spins are in
The ortho-positronioum decays by self-annihilation emitting 3 γ, but, having a so long decay time It is more likely to occur a 1.8 ns pick-off process with a free electron. These ortho-positronioum pick-off events are the main responsible of angular deviation rickey˙lifting˙nodate.
1.2.2
Photon interaction with matter
After being emitted in the annihilation, the two 511 keV γ photons interact with atoms of the surrounding tissues and of the detector.The four main pro-cesses through which gamma rays interact with matter are the photoelectric effect, the e+e− pair production, the coherent scattering and the incoherent
scattering (Compton scattering). However, considering PET energies, pair production and coherent scattering can be neglected. Besides the energy of the photon, the type of predominant interaction also depends on the atomic number of the absorbing medium. Thus, in human tissue with Zeq of 7.5
Compton effect is predominant whereas in material with higher Z also Pho-toelectric effect occur.
Compton Scattering
Compton scattering is an example of inelastic scattering. The incident pho-ton interacts with a charged particle of the absorber material, usually an electron, transferring part of its energy. The interaction results in the emis-sion of a recoil, or Compton, electron and a scattered photon with the loss of energy determined by the angle of scattering θ like in figure 1.2
Figure 1.2: Compton Scattering
Compton scattering of photons in tissue results in an inaccurate LOR counting, generating noise in the reconstructed image. Hence, an adequate
correction is necessary to obtain quantitative information del˙guerra˙ionizing˙2004 levin˙monte˙nodate.
Photoelectric effect
Photoelectric effect consist in the emission of an atomic electron, usually called photo-electron, after the material has absorbed a photon (Figure 1.3). In order to this effect to happen it happens the energy of the incident photon must be higher than the binding energy of the electron.
During the process, the photon transfers all of its energy to the photo-electron, so that the energy of the emitted photo-electron Ee is given by:
Ee− = hν − Eb
with hν being the photon energy and Eb the binding energy of the electron.
The vacancy created by the ejected electron is filled by an electron from a higher energy shell if present. The difference of energies between these shells is converted in the emission of an X-ray photon. The probability of pho-toelectric absorption increases rapidly with atomic number of the absorber (Z3 ∼ Z4) and decreases proportionally to E3
γ.
Photon attenuation
The total probability for all the possible interactions constitute the linear attenuation coefficient µ. If I0 is the initial intensity of a parallel beam of
mono energetic photons, then the intensity Ix at a distance x through some
attenuating object will be given by:
I(x) = I0e− Rx
0 µ(x)dx
From this equation it is possible to calculate the linear attenuation coeffi-cientRLORµ (x) dx that is a function only of the thickness of the object along each LOR regardless of the position of the source zanzonico˙positron˙2004. Considering that a photon of 511 keV has a mean free path in water of about 10 cm, raw data must be corrected for attenuation otherwise activity would be underestimated.
1.3
Coincidence events
PET is based on the principle of annihilation coincidence detection of the two collinear 511 keV photons. Where, if both photons are detected by oppo-site detectors, validated in energy, and they arrive within a predefined time window, they can be deemed to be originated from an annihilation. That event has occurred along the line that connects the two detection point, the so called line-of-response (LOR)wrenn˙use˙1951. LORs are thus defined through an electronic collimation, bringing a significant gain in sensitivity
with respect to physical collimation zanzonico˙positron˙2004cherry˙physics˙2003. The set of all successfully detected LORs constitutes the raw data used for
Figure 1.4: electronic collimation
In addiction to true coincidence events a number of other types of events occur that contribute to the deterioration of the image quality. These events can be classified in three groups and the most important types of coincidence events are summarized in figure 1.5.
Scatter coincidences occur when at least one photon from the an-nihilation undergo Compton scattering before being detected. Due to the change in the direction of the scattered photon, it can be easily assigned to the wrong LOR producing mispositioned events. Since Compton interaction is an inelastic scattering these false-coincidence can be minimize introducing an energy window.
Random coincidences occur when two photons not arising from the same annihilation event are detected within the coincidence time win-dow of the system. The number of random coincidences in a given LOR is closely linked to the rate of single events measured by the detectors joined by that LOR and the rate of random coincidences in-creases roughly with the square of the activity. Thus, the random event rate between two detector can be estimated from the coincidence time window width τ and the single rates (NA and NB ) of the detectors as
follows:
NRandom = 2τ NANB
The distribution of random coincidences is fairly uniform across the FOV, and will cause isotope concentrations to be overestimated if not corrected for and tend to degrade contrast and signal-to-noise ratio (SNR) hoffman˙positron˙1986.
Multiple coincidences occur when more than two photons are de-tected in different detectors within the coincidence time window. Con-sequently, It is impossible to assign a LOR to the event that is then rejected.
Figure 1.5: Type of events in PET
1.4
Time of Flight PET
Time-of-flight (TOF)-PET estimates approximately the position of annihila-tion along the line of response using the measured difference in arrival times. The principle of TOF has been proposed in the early days of PET technol-ogy, but due to the the poor limited timing resolution and stopping power of the scintillators was dropped till late 1990s vandenberghe˙recent˙2016 campagnolo˙tomographie˙1979. Afterwards the availability of lutetium oxyorthosilicate (LSO), a scintillator with good timing and energy resolution and excellent stopping power, as well as fast photomultiplier tubes and impor-tant improvement in electronics have encouraged the development of
TOF-PET system moses˙prospects˙1999. The main motivation for TOF-TOF-PET has always been the reduction of statistical noise moses˙advantages˙2003. Conventional PET reconstruction uses the time information only to identify the line along which the annihilation occurred. It is unable, though, to de-termine where, along that line, is the source of the annihilation; therefore all the points along the line are given the same probability of emission. TOF, instead, gives the likelihood that the event occurred at a certain position along the LOR using the measured difference in the photon arrival times (Fig. 1.6).
Figure 1.6: Bottom left : Conventional reconstruction, the annihilation position between the detectors is not known. Bottom right : TOF-PET reconstruction in which each time-bin is weighted accordingly by the probability that the source is located in that pixel.
The accuracy ∆x of this estimation depends on the precision achieved in defining the difference between time of arrival ∆t. It therefore depends on the time resolution of the system as described in the following formula:
∆x = c∆t
2 (1.1)
where c is the speed of light. Therefore, for example, a time resolution of 500 ps would provide a position accuracy of 7.5 cm along the LOR.
It is obvious that confining the point of annihilation within a portion of object size will produce a reduction of statistical noise and thus an im-provement in contrast in reconstructed images. Although the largest gain is achieved in heavy patients and total body reconstruction, a recent paper schug˙initial˙2016 shows gains due to the TOF even in 11-cm-diameter ob-jects, thus in geometries really comparable to the TRIMAGE one. Since in PET detectors spatial resolution is of the order of 4 mm in clinical PET scan-ner noauthor˙positron˙2016, TOF measurements do not improve directly the spatial resolution of a PET reconstruction but contribute to increase the signal-to-noise-ratio (SNR). SNR under the assumption of uniform distribu-tion of activity in an image element of size d is defined as conti˙state˙2009:
SN R =√N ECR = cost · n1/2 · N 2 true N2
true+ Nscatter2 + Nrandom2
2
(1.2)
Where NECR is the noise equivalent count rate and Ntrue is the total
num-ber of true coincidence in the image, Nrandom the total random coincidences,
Nscatter scatter events and n the number of volume elements in the LOR,
thus contributing to form the image. In TOF reconstruction n can be esti-mated from equation 1.1 as nT OF = ∆xd And the improvement due to TOF
reconstruction can be evaluated as: SN RT OF =
r D
c∆t· SN Rno−T OF (1.3) with D the diameter of the scanner. In tab. 1.2 , the estimated TOF NEC gain is reported as a function of the time resolution, given a object of dimen-sion equivalent to a 30 cm diameter cylinder.
Time resolution ns ∆x (cm) TOF NEC gain
0.1 1.5 20
0.3 4.5 6.7
0.5 7.5 4.1
1.2 18.0 1.65
Table 1.2: TOF NEC gain for a 30 cm diameter uniform cylinder for different value of time resolution
1.5
Multimodality imaging
Positron emission tomography (PET) is a very sensitive functional imaging modality, thus can provide precious information about metabolic processes,
but it gives almost no anatomical information. To overcome this limitation PET system should be integrated with other imaging technique based on different physical principles. In the late 1990s first PET/X-ray computed tomography (CT) scanners was tested beyer˙combined˙2000 obtaining ex-cellent results because was able to combine exex-cellent skeletal contrast with molecular imaging and are thus now widely used in clinical application. How-ever, PET/CT exposes the patients to an additional radiation dose, offers only limited soft tissues contrast and does not allow simultaneous imaging. Consequently, hybrid PET/MRI system are believed to overcome these lim-its.
Magnetic Resonance Imaging investigates structure and function through the interaction of strong magnetic fields with protons present in tissues’ water molecules. The excellent soft-tissue contrast provided by the MR imaging (MRI) modality is expected to drastically increase the possibilities in neu-roimaging, as well as other fields. The lower dose delivered in PET/MR examinations compared to PET/CT is of great importance in therapeutic progress monitoring, where multiple examinations are necessary, especially in paediatric oncology pichler˙pet/mri:˙2010. The idea to combine PET and MRI dates back to the mid 1990s even before the advent of PET/CT christensen˙positron˙1995. The integration of a MRI and a PET system involves different technical hardware challenges that have to be resolved, and are still representing frontier challenges. The main issues can be summarize as follow:
the photomultiplier technology must be replaced with magnetic field–insensitive solid-state photodetectors coupled with read-out electronics with
min-imum heat radiation.
PET detectors must be constructed to be invisible to the MRI and to not interfere with the field gradients or MR radiofrequency.
The MRI scanner must be adapted to accommodate the PET detectors and to allow simultaneous data acquisition without mutual interference.
1.5.1
Possible design of PET/MRI systems
There are several ways to combine PET and MRI imaging of the same patient(Fig. 1.7)zaidi˙outlook˙2011vandenberghe˙pet-mri:˙2015. The simplest one is to adopt the same arrangement as for PET/CT, in the so-called “tandem” configuration, where the two examinations are taken se-quentially, moving the patient from one machine to another. However, the lack of simultaneity of the two modalities introduces organ motion effects
and prevents the possibility of functional PET-functional MRI simultaneous investigations. This issues have brought to the “insert” concept, employed in the first prototype of clinical PET/MRI scanner for head only where a small size PET is inserted in the bore of the MRI and inside the PET scanner a RF coil for brain imaging is positioned schlemmer˙simultaneous˙2008. However, these systems have a poor timing resolution that doesn’t make possible TOF reconstruction and are particularly sensitive to degradation of PET performance due to interference between electronics and RF of gradi-ent pulse of MR. Lastly, “fully integrated” scanner, where a dedicated PET scanner is built in a dedicated MRI scanner. This latter solution is the most challenging one, but it is certainly the one that can bring to a real step forward the use of combined PET/MRI systems in diagnosis, therapy, and follow-up. The benefits of fully integrated systems are firstly to im-prove patients comfort, to reduce examination times and overall costs, to overcome any registration problems between modalities, but also to simulta-neously observe several biological functions related to any studied pathology brasse˙instrumentation˙2016.
Figure 1.7: Three concepts for integration of PET and MRI: a brain PET scanner inserted in a whole body MRI (top), a sequen-tial PET MRI scanner for whole body imaging (middle) and a fully integrated simultaneous PET MRI scanner for patient imaging (bottom).
1.6
State of the Art
As of today, several clinical whole body PET/MR systems have been devel-oped and brought to the market by major manufacturers such as Siemens Healthineers and GE Healthcare. In last years, a number of brain-dedicated PET prototypes have been developed to be used in combination with existing MRI scanners. Despite many technical challenges the development of brain
dedicated PET/MRI systems, is an hot topic in nowadays literature and sev-eral attempts have been reported. PET systems designed with this approach are usually called PET inserts. The motivation behind these projects is to provide brain-dedicated tools with better performance in terms of spatial resolution and sensitivity than clinical whole body PET/MR systems.
Regarding the choice of the basic components of the detectors, PET/MRI scanners are based on solid-state photodetectors and fast scintillators. The first developed system kolb˙technical˙2012 is the only one based on APDs, all other prototypes make use of SiPM as photodetectors. As scintillation material all of these systems use LYSO crystals. It follows a brief descrip-tion of the main scanner prototypes present in literature. They represent the state-of-the-art in PET/MRI brain scanners and their main feature are summarize in tab. 1.3 in order to ease a comparison with the results obtained in this thesis.
1.6.1
Mindview
Mindview is a project aiming to achieve 1 mm resolution in brain-PET. This scanner is composed by three rings of 20 detector blocks with a dis-tance of about 330 mm between opposite detectors. Each detector module includes a monolithic LYSO crystal of 50 × 50 × 20 mm3 and a custom
12 × 12 SiPM array (TSV-type, SensL). The system defines an axial and trans-axial fields of view (FoV) of about 46 mm and 240 mm, respectively benlloch˙mindview˙2018. First performance tests of the PET detector found spatial resolution (FWHM) values ranging from 0.7 mm to 1.77 mm. The energy resolution slightly depends on the interaction depth. The overall energy resolution is better than 13 % and average DOI resolution is around 4 mm making it possible to efficiently correct for the parallax error using such thick monolithic crystals.
1.6.2
Other detector prototypes
Korean research institutes are developing a PET/MRI detector for human brain. The PET insert consists of 18 detector blocks, circularly mounted on a custom-made plastic base to form a ring with an inner diameter of 390 mm and axial length of 60 mm. Each detector block is composed of a 4× 4 matrix of detector modules, each consisting of a LYSO scintillator array coupled to a 4× 4 GAPD array. The LYSO array consisted of 3 × 3 × 20 mm3 crystal pixels arranged with a 3.3 mm pitchhong˙prototype˙2013.
The PET gantry is shielded with gold-plated conductive fabric tapes with a thickness of 0.1 mm. The charge signals of PET detector is transferred
via 4 m long flat cables to the position decoder circuit. The flat cables were shielded with a mesh-type aluminum sheet with a thickness of 0.24 mm. With this detector they achieved an average energy resolution of about 20% and a timing resolution of 3.8 ns jung˙development˙2015.
Chiba University, in Japan developed a brain-dedicated prototype that con-sists of a RF-head coil with four-layer depth-of-interaction (DOI) PET de-tectors to be coupled with a 3 T MRI. Each four-layer DOI-PET detector consisted of six multi-pixel photon counter (MPPC) arrays, two scintillator crystal blocks, a readout circuit board and a copper shielding box. The scin-tillator blocks consist of LYSO scinscin-tillators arranged in 19 ×6 ×4 layers with reflectors inserted between them. The size of each crystal element is 2.0 mm ×2.0 mm ×5.0 mm nishikido˙development˙2017.
type of crystal Energy res. CTR DOI information Mindview monolithic block ∼13 % 2-3 ns yes, from light
characterization South Korea
research Institutes
dual layer
staggered matrix ∼20 % 3.8 ns yes Chiba University 4 layer matrix ∼19.4 % n.d. yes
Table 1.3: Summary of specification and performance of the main brain PET/MRI scanners
1.7
TRIMAGE multimodality imaging
TRIMAGE is a fully integrated PET/MR/EEG (Positron Emission Tomog-raphy – Magnetic Resonance – ElectroencephalogTomog-raphy) scanner that coupled a 1.5 T non-cryogenic magnet with a commercially available MR-compatible EEG and PET scanner that makes use SiPM matrices as photodetectors, LYSO scintillators matrices and most advanced digital electronics. The com-bined PET/MR/EEG system is dedicated to brain imaging and has an in-ner diameter of 260 mm and an axial Field-of-View of 160 mm. The three modalities complement each other well: MRI supplies structural–functional imaging, metabolic imaging with specific tracers is the strength of PET, and EEG provides finely grain temporal information where the other two modal-ities are weak. The TRIMAGE project aims to discover and develop suitable biomarker combinations that are based on structural-functional-metabolic
changes in the brain. It is focused on alterations in the dopaminergic and the glutamatergic neurotransmitter systems investigated using appropriate PET tracers, since recent studies has pointed out correlations with schizophrenia lodge˙hippocampal˙2011rotaru˙role˙2012. Additionally, alterations in brain structure in schizophrenic patient are habitually observed with stan-dard MRI but they are difficult to detect in the prodromal phase of the illness and could be common to patients with psychotic features across diag-nostic boundaries strasser˙hippocampal˙2005. MR and EEG scan will be now briefly described whereas in the following chapter the TRIMAGE PET system will be analyzed in detail .
Conversely to the insert approach where existing MR scanners are em-ployed, the TRIMAGE system features a custom designed MR sys-tem. The magnet integrates a very compact design with a cryogen-free (i.e. using no liquid helium or liquid nitrogen) magnet cooled with a pulse tube cryorefrigerator at the most widely available clinical field of 1.5T. The limited axial dimension reduces the claustrophobic effects for the patient and gives the possibility of PET bolus injection under control, the arms outside of the magnet. The selected RF coil is de-signed in two layers: the outer layer is a dedicated transmit-only and the inner part is a receive-only.
The EEG cap used is a commercially-available, 64-channel MR com-patible one. An explorative study investigating and quantifying the ef-fect of the components of an MR-compatible EEG cap during a simulta-neous trimodal study on PET images has just been completed in Juelich showing no significant artefactsrajkumar˙simultaneous˙2017.
Chapter 2
Scintillation Detectors
The scintillation detector is the key component of any PET system. Despite variations in the design, the principle of a scintillation detector for PET applications is to generate an electrical signal proportional to the energy of the particle to be detected and to provide information on time arrival of the particle. Essentially, a scintillation detector consists of three components, a scintillator which absorbs the incoming particle and converts its energy into scintillation light, a photodetector which measures the light intensity and produces proportional electrical signals and the electronics that further processes the signal so that the information can be correctly stored and analysed later.
2.1
Scintillators
Scintillation detectors exploit a property of certain materials called lumines-cence, to emit visible light after being excited by radiation. These materials are known as scintillators and are usually coupled with a Photodetector, such as a photomultiplier tube (PMT), photodiode, or silicon photomulti-plier (SiPM), that converts light to current signal. If the re-emission occurs within 10−8 s after absorption, the process is called fluorescence, otherwise, if the excited state is metastable, the process takes considerably more time (from microsecond to hours depending on materials) and is called phospho-rescenceleo˙techniques˙1994. The time evolution of the number of emitted scintillation photons N can be described as:
N = A exp(−t/τf) + B exp(−t/τs) (2.1)
where τf and τs are the fast (or prompt) and the slow (or delayed) decay
and B. In this model we are omitting the rise time of the signal assuming it to be far lower than τf and τs. Since it is typically in the order of 10−1
ns whereas τf for inorganic scintillators is tenths of ns, rise time is negligible
compared to τf for most of scintillators eijk˙inorganic˙2002. However, for
faster organic scintillators, which have a decay time of 1 or few ns rise time can no longer be ignored and a bi-exponential model that considers both rise and decay time better describes the scintillation mechanism. Consequently, for a very fast scintillator it is possible to omit τs and equation 2.1 can be
written as shao˙new˙2007
N = A [exp(−t/τd) − exp(−t/τr)] (2.2)
where τd is the decay constant and τr is the rise time.
2.1.1
Scintillator properties
Scintillation materials exist in many types and can be divided into two cat-egories: organic and inorganic. This demarcation is important because scin-tillation process is based on different mechanism in organic and inorganic scintillators. In both of them emitted photons have longer wavelengths, and thus are less energetic, than the energy gap of the excitation. This is called Stokes shift. This effect allows photons to propagate through the material because they can’t be self-absorbed by re-exciting the material.
The elective properties in choosing a scintillator for PET applications are: high stopping power at 511 kev. It determines the mean distance that a photon travels before it deposits all of its energy within the crystal and thus determines detection efficiency. High stopping power implies that scintillator can be shorter, minimizing the size of the de-tector and its cost. It depends on density and effective atomic number (Zef f) of the scintillator.
light output. It can be defined as the number of photons emitted per unit of absorbed energy. This parameter is related to geometry, energy resolution and timing performance of the scintillator. Further-more other features like the optical coupling with the photodetector as well as surface finish and the reflector material around the scintillator must be taken into consideration.
fast decay time. Giving that a PET scanner works at high count rates a short decay time is desirable. This is due to the fact that decay time limits the count-rate performance due to pulse pile-up.
linearity. The scintillator response to the exciting energy must be linear in the range of interest.
the scintillator must be transparent to the emitted wavelength in order to prevent possible re-absorption.
optimal coupling to photo-detector. In order to optimize coupling scintillator must have a index of refraction near that of glass (∼1.5) and the emission spectrum must overlap the spectral response of the photo-detector.
reasonable cost for materials and manufacturing.
From this parameters it is possible to estimate efficiency as well as energy and time resolution of the detector. For this purpose is often useful to define the Initial photon emission rate I0, which determines the timing performance
of the detector. It can be easily estimated from the light output Lo and
scintillation decay time τd as I0 = Lo/τd
2.1.2
Organic scintillators
Organic scintillators are aromatic hydrocarbon compounds which contain benzene ring structures interlinked in various ways. Their most distinguished feature is a very rapid decay time on order of few nanosecond or less. Further-more they can be produced in any geometry so that they can be customized for specific applicationsahmed˙physics˙2014. The scintillation mechanism is determined by the chemistry and physics of the benzene ring and thus an organic scintillator will scintillate whether it’s in a crystal form, is a liquid, gas or embedded in a polymer. Scintillation arises from S-state elec-tronic transitions made by free valence electrons of the molecules. Ionization energy from incident radiation excites molecules from S0 ground levels to
vibrational levels in the S1 band, which act like a fine structure. Then the
excited state decays radiatevely to vibrational sub-levels of the ground state in few ns. This is the normal process of fluorescence described by the prompt exponential in 2.1. The most commonly used type of organic scintillator are: organic crystals, like anthracene (C14H12) and naphtalene (C10H8).
They are hard crystals and thus very durable, but their response is anisotropic and shaping is difficult.
liquid scintillators, are liquid solutions of one or more organic scin-tillators in an organic solvent. Liquid scinscin-tillators are easily loaded with wavelength shifters or additives to increase the neutron detection efficiency.
plastic scintillator, are fluorescent emitters suspended in a solid poly-mer matrix. Plastics offer extremely fast signal with decay constant of about 2-3 ns and high light output. Moreover they are easily shaped to desired forms and relatively cheap.
Organic scintillators are used for timing applications or for neutron de-tection, having a good content of H. On the other hand, the relatively low Zef f make them not really suitable for PET system.
2.1.3
Inorganic scintillators
In inorganic scintillators the scintillation mechanism depends on the structure of the crystal lattice. Inorganic scintillators are usually insulators and thus electrons have only discrete bands of energy available. The electrons in the valence band are essentially bound at lattice sites, whereas those in conduc-tion band have sufficient energy to be free to migrate throughout the crystal. Energies in the range between these two bands are forbidden. Electrons in the valence band can absorb energy from the incident photon and can be promoted to the conduction band, leaving a hole in the valence band. Since this is not the ground state, the electron de-excites by releasing scintillation photons. However, in pure crystals this is an inefficient process and normally, the value of the energy gap, Eg is such that the scintillation occurs in the
ultraviolet range. By adding impurities to a pure crystal, such as adding thallium to pure NaI (at a concentration of 1%), the band structure can be modified to produce energy levels in the prior forbidden region. These impu-rities are usually called activators and they create special sites in the lattice at which the band gap structure is modified in order to have an emission in the visible spectra called luminescence centre(between 2 and 4 eV). Scintil-lation process in inorganic crystal is slower than in organics, typical lifetimes for such excited states are of the order of 30-500 ns knoll˙radiation˙2010. One important consequence of luminescence through activator sites is the fact that the crystal can be transparent to the scintillation light. In the pure crystal, indeed, the emitted energy of the electron-hole recombination is roughly the same of the energy gap between valence and conduction band. As a result self-absorption is likely to occur. The emission from a doped crystal, instead, occurs at an activator site where the energy transition is less than the one required for the creation of the electron-hole pair 2.1.
Figure 2.1: Scheme of scintillator mechanism in inorganic crystals. On left a pure scintillator crystal, on right doped crystal
For inorganic crystal the scintillation efficiency η resulting from the ab-sorption of a photon of energy hν can be expressed by the following formula:
η = hν Ee−h
· T · q
where Ee−h is average energy required to produce an electron-hole pair, T is
transfer efficiency of the excited state on the luminescence centre and q repre-sents reprerepre-sents the efficiency for photon emission of the luminescence centre itself. Considering these transferring efficiency factors, T is the least pre-dictable. It depends very much on defects in the scintillator, other than the luminescence centres, that may capture electrons or holes or both, and pro-duce radiation-less transitions or afterglow upon thermal excitation. These defects can arise from the interactions themselves, from crystal growing, or can be due to impurities. Inorganic scintillators are generally denser and have higher effective atomic numbers Zef f than organic scintillators. This makes
them attractive in applications where high stopping power for the incident radiation is desired, like in PET detectors. Another advantage is their higher light output compared to organic scintillators. In tab 2.1 the main properties of various inorganic scintillators used in PET detector are summarized.
NaI BGO LYSO YAP LaBr3
Peak emission wavelength (nm) 410 480 420 350 360 Light yield (103ph/Mev) 41 9 30 17 60
Decay time (ns) 230 300/60 40 30 16 I0 at 511 Kev (ph/ns) 90 21 380 290 1900
Effective Z 50 73 64 33 46
Stopping power at 511 keV (mm) 25.9 11.2 12.6 21.3 22.3
Table 2.1: Physical properties of various inorganic scintillators used in PET detector lecomte˙novel˙2009
2.2
Depth of Interaction
After a photon enters a detector, it travels a short distance before it deposits its energy. This distance is called depth of interaction (DOI) and if not considered, the measured position of energy deposition is assumed to be on the entrance surface of the detector. For photons that enter the detector at oblique angles, this projected position can produce significant deviations from the real position (fig. 2.2), leading to a blurring of the reconstructed image.
Figure 2.2: Perceived and actual line of response (LOR) in PET imaging due to not DOI correction. Figure from bolus˙petmri:˙2009
Moreover, gamma rays and scintillation photons travel at different speeds in the scintillator. As a result, also the time between the entrance of the gamma ray and the detection of the scintillation light depends on the depth of interaction. Furthermore, photons that impinge on the detector ring at an oblique angle can penetrate into adjacent crystals which causes mispo-sitioning errors moses˙trends˙2001. Thus, the depth-of-interaction (DOI) encoding ability of the detectors is critical for high performance imaging, especially in detectors with small FOV. DOI information can be obtained by multiple ways such as: stacked layers, double-side readout, Phoswich design, multiple layer with reflective structures, multiple layers of photosensors and
(a) Stacked layer. (b) Double-sided read out. (c) Phoswich de-sign. (d) Layers with reflective optical structure.
(e) Multiple pho-tosensors.
(f ) Width of the light spot in the monolitical scintillators.
Figure 2.3: Different approaches to design systems with DOI encoding
pulse shape discrimination. In fig. 2.3 different approaches to DOI encoding are reported.
2.3
Photodetectors
In order to provide a useful measurement device, an electrical signal has to be formed from the scintillation light. Two main types of device are used to do that: photomultiplier tubes and semiconductor photodetectors (photodi-odes). The sensitivity of a photodetector is usually defined in terms of the quantum efficiency. The quantum efficiency is defined as the probability of the conversion of light to an electrical signal and is defined as:
Quantum ef f iciency = number of photoelectrons emitted number of incident photons
2.3.1
Photomultiplier Tube
Photomultiplier tubes (2.4) represent the oldest and most reliable technique to measure and detect low levels of scintillation light, thus they are widely used in PET and SPECT detectors. Photomultiplier tubes are vacuum tubes in which a photocathode and a series of charge amplifying elements called dynodes are placed. A light photon may interact in the photocathode to eject a low-energy electron into the vacuum. This process can be thought to occur in four steps:
1. The incident photon is absorbed by an electron of the photocathode to which it transfers all of its energy.
2. In consequence of the applied electric field the electron is accelerated towards the first dynode.
3. Upon impact with the first dynode the electron transfers some of its energy to other electrons. This causes secondary electrons to be emitted and then accelerated to the second dynode where they would release more electrons. The result after several dynodes is an electron cascade. 4. The electron cascade is collected in the anode to give a current signal.
Figure 2.4: Scheme of the PMT signal production mechanism when cou-pled with a scintillator crystal
This structure, based on multiple dynodes set at increasing potentials, allows to reach a large gain, of the order of 107, on the anode. This high gain leads to a very good signal-to-noise ratio (SNR). Unfortunately, PMTs are extremely sensitive to magnetic fields, which renders them unsuitable for use in combination with MR. Even very weak magnetic fields, can have an effect on PMT signals due to the deflection of the electron trajectories between the
photocathode and dynodes and thereby affects its efficiency. Thus, significant variations of the gain and energy resolution of PMT-based detectors are observed as soon as the magnetic field strength is increased above ∼10 mT lecomte˙novel˙2009.
2.3.2
Semiconductor photodetector
Semiconductor photodetectors have high sensitivity for detecting low en-ergy scintillation photons. These detectors typically are in the form of PIN diodes (PIN refers to the three zones of the diode: P-type, Intrinsic, N-type). They have several advantages over PMT: high quantum efficiency, compact and flexible shape that can be adapted to individual crystals, ruggedness, demonstrated insensitivity to magnetic fields up to at least 9.4 T and are less expensive lecomte˙novel˙2009. Incident scintillation photons produce electron–hole pairs in the detector and an applied electric field then results in a flow of charge that can be measured through an external circuit. Unfortu-nately, a significant disadvantage of the photodiodes is the low SNR achieved due to the presence of thermally activated charge flow and very low intrin-sic signal amplification. Since the noise is at the level of several hundred electrons, the smallest detectable light flash needs to consist consequently of even more photons.
Avalanche Photo Diode
APDs are compact semiconductor devices, usually silicon devices, based on a modified p–n junction structure (fig. 2.5). APDs are designed such that when a bias voltage is applied, a depleted region is created. The field is high enough that visible light from a scintillator can create a hole–electron pair by photoelectric effect. Charges produced in this region may be accelerated sufficiently to create further electron-hole pairs in the region by impact ion-ization, thus causing a multiplication or “avalanche” of charges that are then collected to form an output pulse. For PET applications, APDs are operated in proportional mode, just below the breakdown voltage. At this bias voltage the output signal is proportional to the amount of scintillation light inter-acting in the APD. APDs are usually operated with a gain of 50–150 where the noise can be kept at a relatively low value, but this low gain combined with the susceptibility to temperature and bias voltage variations make them unsuitable for time of flight PET.
Figure 2.5: Avalanche-photodiode structure. Image from Hamamatsu photonics
Silicon Photomultiplier
Silicon PhotoMultipliers (SiPMs) consist of a 2D array of micro-cells con-nected in parallel (in fig 2.6 are reported a SiPM matrix and its compo-nents). Each micro-cell, named SPAD (Single Photon Avalanche Diode) is an Avalanche Photo-Diode operating with a bias voltage higher then the breakdown voltage, thus in Geiger-mode(G-APD).
Figure 2.6: SiPM structure lecomte˙novel˙2009
In Geiger-mode the charge multiplication is diverging, while the ampli-tude of the generated signal due to a single photon is no more related to its energy, the gain is comparable to PMTs (105 − 106). The output
to the number of cells in which an avalanche occurred. If the SiPM is il-luminated by a low-intensity light source (number of photons number of microcells) and there is good gain uniformity between the cells, it can be assumed that the output pulses produced have amplitudes proportional to the number of photons that are converted into a Geiger discharge within the device roncali˙application˙2011. For this purpose it is important to insert a quenching circuit to reset the cell once the avalanche is started and make it ready to detect another photon. The output of each microcell is termi-nated by either passive or active quenching. Passive quenching of a cell is commonly obtained via a resistor placed in series with the G-APD whereas active quenching requires circuitry such as comparators to detect the pres-ence of an avalanche and rapidly quench it, resetting the cell. Thus, the response of a cell of a SiPM charged at Vbias > VBD (with passive quenching)
to impinging light is reported in fig. 2.7 and can be summarized as follow:
1. When the optical photon hits the cell an input of charge is produced. The the cell discharges through quenching resistor.
2. Due to passive quenching, the reverse voltage decreases to VBD when
current flows through the quenching resistor. The avalanche stops but the cell is no more in Geiger mode.
3. The voltage returns to Vbias.
SiPM parameters
SiPM efficiency depend on many parameters, the most important are:
Overvoltage
Overvoltage is the difference between the bias voltage and the breakdown voltage. It is the parameter that mainly determines the performances of a SiPM. Therefore, all the following parameters (detection efficiency, gain and noise) are strictly dependent on Vov.
Photon Detection Efficiency
Since SiPMs are constructed as a group of small detectors the Photon De-tection Efficiency (PDE) is a function of Fill Factor (FF), which is defined defined as the ratio between the sensible area and the total area o the device. Thus the PDF can be expressed as:
P DE = QE · εg· F F = QE · εg·
Asens
Atot
where QE is the quantum efficiency and εg is the probability that a charge
carrier activates the Geiger process bisogni˙development˙2016. The PDE is a function of the wavelength of the photon impinging on the detector and the parameter εg is strictly dependent on the overvoltage. Therefore, the
higher the overvoltage, the higher the probability that a photoelectron trig-gers the avalanche in the depletion region. The QE is 80–90% at maximum depending on the wavelength renker˙geiger-mode˙2006. The fraction of scintillation light detected by the SiPM depends also on the transmission probability of the photon from the scintillating crystal to the sensitive area. Most of the scintillators used in PET applications have a high refractive index; hence an optical coupling with epoxy resins, silicon glues or silicon grease is usually used to match the refractive index of the silicon oxide. The optimization of optical coupling of TRIMAGE detector has been studied in this thesis and will be analized in Chapter 4.
Gain
SiPM devices have gains comparable to PMTs, in the order of 105 − 106.
The gain of a SiPM is expressed as the ratio between the charge produced by the avalanche and the primary charge produced by the interaction of the optical photon within the device. Since the avalanche is interrupted when the voltage at the two sides of the micro-cell goes down to the breakdown
voltage, the gain G can be expressed as:
Gain = (Cd+ Cq) · (Vbias− VBD)
q (2.3)
where (Cd+ Cq) is the total micro-cell capacitance, given by the sum of the
diode capacitance and the parasitic capacitance (mainly due to the quenching resistor), Vbias is the bias voltage, VBD is the breakdown voltage and q is the
electron charge. SiPM gain is strongly related to temperature because a higher temperature involves a higher breakdown voltage.
Dynamic Range
For PET applications it is important that SiPMs have a high dynamic range. The dynamic range of a SiPM is dominated by the number of available pixels for detection that represents an upper limit to the number of photons that can be detected at the same time and thus the dynamic range. However, both the bias voltage and temperature affect the dynamic range of the SiPM since the PDE depends on them.
Recovery time
Recovery time is the time needed to recharge a cell after a breakdown has been quenched. It depends mostly on the cell size, due to its capacity and the individual resistor.
SiPM noise sources
SiPMs have a much lower noise rate than APDs, however it still is an impor-tant parameter. Noise is mainly caused by three factors in these devices:
Dark count rate refers to the spontaneous Geiger discharge of a SPAD without any photon detections due to thermally generated electron-hole pairs. Thus the avalanche on a single pixel can be triggered without a scintillator photon impinging on it. Thermally generated free carri-ers can be reduced by cooling the device. There is a factor 2 reduc-tion of the dark counts every 8 ◦C. Another possibility is to lower bias voltage resulting in a smaller electric field and thereby lower gain renker˙geiger-mode˙2006. Dark count rate increases also with the dimension of SPAD.
Afterpulse occurs when a charge carrier is trapped during a Geiger discharge and released after a certain delay to trigger a new avalanche. It is mainly due to silicon defects and depends on the cell recovery time and temperature (increases with low temperature).
Optical cross talk occurs when the breakdown of one pixel in a SiPM trigger an avalanche in a neighbouring pixel. Optical photons can be produced within a SPAD cell during an avalanche breakdown and can potentially move to a neighbouring cell, triggering it just as an incident photon would. To reduce crosstalk, cells need to be optically isolated. These solutions, however,reduce fill factor and adversely affect the pho-todetection efficiency.
PMT APD SiPM
Gain 105− 107 102 105− 106
Dynamic range 106 104 103/mm2
Excess noise factor 0.1 − 0.2 > 2 1.1 − 1.2 Rise time < 1ns 2 − 3ns ∼ 1ns Dark current < 0.1mA/cm2 1 − 10nA/mm2 0.1 − 1M HZ/mm2
Bias voltage ∼ 800 − 2000V ∼ 100 − 1500 V ∼ 30 − 50V Temperature coefficient < 1%/K 2 − 3%/K 3 − 5%/K Magnetic susceptibility Very high(mT) NO NO
Table 2.2: Characteristics of photodetector for PET. Table from lecomte˙novel˙2009. Temperature coefficient express the temperature dependence of gain.
2.4
Read out electronics
As we have seen, photodetectors provide informations on detected radiation in the form of electrical signal. In order to extract this information the signal must be processed by an electronic system. In PET detector information usually comes in the form of a pulse signal. Regarding pulse signals we can define the following features leo˙techniques˙1994
Baseline. The baseline of the signal is the voltage or current level to which the pulse decays. While this is usually zero, it is possible for the baseline to be at some other level.
Amplitude. The amplitude is the height of the pulse as measured from its maximum value to the instantaneous baseline below this peak.
Signal width. This is the full width of the signal usually taken at the half-maximum of the signal (FWHM).
Leading edge. The leading edge is that flank of the signal which comes first in time.
Falling edge. The falling edge or tail is that flank which is last in time.
Rise Time. This is the time it takes for the pulse to rise from 10 to 90% of its full amplitude. The rise time essentially determines the rapidity of the signal and is extremely important for timing applications. Fall time. In analogy to rise time is defined to be the time it takes for
the signal to fall from 90 % to 10 %.
In a PET camera, each detector generates a timed pulse when it registers an incident photon. These pulses are then combined in the coincidence circuitry, thus information on the precise arrival time of a pulse in the detector is of particular interest. If one pulse occurs within one other coincidence time window the two photons are assumed to be from the same annihilation. The width of time window depends on the scintillator, photodetector response and electronics. PET, as well as many others applications require to mea-sure time intervals with a resolution of the order of 1 ns. For this purpose, time pick-off units or triggers are used. These devices generate a logic pulse based on the condition at their input and the input signal. Two methods are commonly used for time pick-off: leading edge timing and constant fraction discrimination. The most important factor in any timing system is its reso-lution, the smallest time interval that can be measured with accuracy. The major source of limitation in time resolution occurs in the generation of the timing logic signal by the discriminator. Two principal effects are walk and jitter (fig. 2.8).
Walk is due to variations in the amplitude or shape of the incoming signals. In fact, coincident pulses of same shape and different amplitude cross the same threshold at different times.
Jitter is due to noise and statistical fluctuations in the original detector signal. Thus, two identical signal will not always trigger at the same point in time. The following time variation will be dependent on the amplitude of fluctuations.
2.4.1
Leading edge timing
Leading edge triggering is the easiest and most direct time pick-off method. It consists in generating a trigger signal when the pulse crosses a fixed dis-crimination level. Leading edge is not often used, mostly because it suffers limitations from walk effect.
Figure 2.8: Effect of walk (up) and jitter (down) on timing discrimina-tors using leading edge method
2.4.2
Constant fraction discrimination
One of the most efficient and versatile method available today is the constant fraction triggering technique. In this method, the logic signal is generated at a constant fraction of the peak height to produce an essentially walk-free timing signal. The scintillator pulses have identical rise times that are much longer than the desired temporal resolution. This forbids simple threshold triggering, which causes a dependence of the trigger time on the signal’s peak height, the time walk.
Figure 2.9: Comparison of threshold triggering (left) used in leading edge timing and constant fraction triggering (right)
2.4.3
Coincidence processing
Once the proper choice of time-pickoff method is made, electronic techniques for measuring the time difference between two signals must be considered. We divide these into analogic and digital techniques. In analogic ones the arrival of the first signal (START) activates a Time-to-Amplitude Converter (TAC), which enables the discharge of a capacitor till the arrival of the STOP signal. The total charge collected from the capacitor thus forms an output signal whose height is proportional to the time difference between the START and STOP signals. The capacitor is then recharged for the next event. In order to obtain a time interval measurement in digital form, one method is to digitize the TAC using an Analog-to-Digital-Converter (ADC). However, it is also possible to construct a Time-to-Digital-Converter (TDC) directly from oscillators and counting techniques. The resolution of the counting TDC depends on the frequency of the clock used: the higher the frequency, the smaller the time interval than can be measured.
Chapter 3
TRIMAGE PET detector
TRIMAGE is a fully integrated PET/MRI/EEG system, thus the TRIMAGE PET detector has been designed to overcome several technical challenges. These challenges can be summarize as:
Physical constraints. The PET system must be positioned between the RF-coil and the gradient coils of the MRI system. This design imposes constraints on the size of detectors and front-end electronics PET detectors should not create artefacts in MR images. MRI
systems require very good uniformity of its magnetic field and the pres-ence of the PET detector, especially power supplies or preamplifier elec-tronics, can causes inhomogeneities in the magnetic field. Furthermore, PET can emit signals interfering with the RF and gradient coils of the MRI.
PET signals should not be degraded by the presence of the MR system. PET detectors are sensitive to noise and may be af-fected by the presence of both the static magnetic field and the ra-diofrequency pulses or magnetic field gradient switching during MR sequences. Moreover, eddy currents may be induced in the conduct-ing structures of the detectors and the MR sequences can generate an increase of temperature that would modify the working point of the SiPMs.
3.1
TRIMAGE PET system
The PET component of the trimodal PET/MRI/EEG TRIMAGE system consists of 18 rectangular segments arranged in a plastic ring. The ring is permanently attached to the MR system through a vibration absorbing sup-port. The distance between two opposite crystals of the PET system is 310 mm and the field of view has a diameter of about 260 mm and an axial extension of about 160 mm. In order to maximize the MRI compatibility several technical aspects have been optimized. First of all, every electronic components is RF-shielded. Each of the 18 segment is composed by the proper detector and a liquid cooling system hosted in a shielding cassette. The frequency-selective shield consists of a solid copper layer without holes or slits, which covers a plastic cassette of dimensions (69 × 64 × 442) mm3 in which the PET components are located. The shield behaves like a Faraday cage providing a high protection for the electrical devices placed inside. In each cassette a PCB board is installed, called TX board on which an FPGA and 3 modules are mounted. A module consists in a board (fig. 3.2), called Asic Board, that host four sub-modules called, tiles, arranged in 2× 2. The tile sub-module is composed of a dual layer LYSO:Ce staggered matrix, 64 SiPMs and a TRIROC ASIC that will be described in detail in the next sec-tion. Furthermore, the PET power supply network has been custom designed for MRI compatibility: the PET power supply is located outside the mag-net bore and it has been designed to allow to compensate for voltage drops. In addition, the low voltages requested by the detectors make possible to avoid switching DC/DC converters in the magnet bore and all the power and ground rails placed near the magnet are design in order to avoid loops. The same approach is followed also in the design of PCBs placed inside the PET cassettes.
Figure 3.1: Scheme of full TRIMAGE PET scanner (left) and segment detector (right). Image from sportelli˙single-mode˙2015-1
