Analysis and development of an high performance preamplifier for positron emission tomography applications

POLITECNICO DI MILANO

Master of Science in Electronic Engineering

Department of Electronics, Information and Bioengineering (DEIB)

Analysis and development of an high

performance preamplifier for Positron Emission

Tomography applications

Polimi Advisor:

Prof. Carlo Ettore Fiorini

Advisor:

Prof. Vicente Herrero Bosch

Master Thesis by:

Riccardo Sarta, 898002

Abstract

Positron emission tomography is a nuclear imaging technique based on the coincident detection of two gamma photons originated by a positron-electron annihilation event. This procedure allows to obtain 3D distributions of the radiopharmaceuticals injected into the patient, through which may be obtained information about the metabolism related to these substances.

Gamma radiation detector technologies have evolved during the last decades, pushing the characteristics of the PET systems to their limits. However, front end electronics has undergone a slower development.

The main purpose of this thesis project is the study and the realization of a preamplifier in CMOS 0.35um technology, which can be integrated to constitute the first stage of the analog channels of an ASIC system.

This preamplifier shall be able to read the signal generated by a photosensor and to retain as much as possible the characteristics of this output signal.

The first step in the realization of this work was a careful analysis of the existing architectures, presenting a review of the latter ones and trying to meet the greatest limitations which we would have faced with.

Afterwards, three similar structures were proposed, coming from an already existing architecture (Flipped Voltage Follower), by which it could be possible to obtain very high performances in terms of signal bandwidth.

At the end, a current comparator has been proposed, which can be combined with such preamplifier so as to detect any incoming 'valid' signals. This is a very important operation in TOF- PET diagnostic systems, in which the speed of the signal detection is essential.

Sommario

La tomografia ad emissione di positroni (PET) è una tecnica di diagnostica per immagini basata sul rilevamento di due raggi gamma generati dall’annichilazione di un positrone ed un elettrone. Questo procedimento permette di elaborare distribuzioni tridimensionali del radiofarmaco che viene somministrato al paziente, mediante le quali è possibile ricavare informazioni del processo metabolico relazionato con tale sostanza.

I rilevatori di radiazioni gamma hanno subito un notevole sviluppo negli ultimi anni, spingendo le caratteristiche dei sistemi PET fino ai loro limiti. Sfortunatamente non si può dire lo stesso della elettronica relazionata a questi sistemi, che ha avuto uno sviluppo più lento.

Lo scopo principale di questo progetto di tesi è lo studio e la realizzazione di un preamplificatore in tecnologia CMOS 0.35um, che possa essere integrato per costituire il primo stadio dei canali analogici di un sistema ASIC.

Tale preamplificatore deve essere in grado di leggere il segnale generato da un fotosensore e conservarne il più possibile le caratteristiche di tale segnale in uscita.

Il primo passo per la realizzazione di questo lavoro è stato un’attenta analisi delle architetture già esistenti, presentando una rassegna di queste ultime e cercando di incontrare quali fossero le maggiori limitazioni con le quali ci saremmo dovuti affrontare.

Successivamente sono state proposte tre simili strutture, procedenti da una già esistente architettura (Flipped Voltage Follower), con le quali si fosse in grado di ottenere altissime prestazioni in termini di banda del segnale.

Per concludere, è stato proposto un comparatore in corrente, che possa essere combinato con tale preamplificatore in modo da individuare qualsiasi segnale “valido” in arrivo, operazione assai importante in sistemi di diagnostica TOF-PET in cui la velocità di rilevazione del segnale è essenziale.

Contents

Sommario ... II

Introduction ... 1

1.

Time-Of-Flight PET ... 3

1.1.

Biophysics in PET ... 3

1.1.1.

Radiation origins and interaction with matter ... 4

1.2.

Gamma ray detectors ... 7

1.2.1.

Imaging Detectors... 9

1.2.2.

SPECT Detectors ... 10

1.2.3.

PET Detectors ... 12

1.3.

PET structure and principles ... 13

1.3.1.

TOF PET... 14

1.3.2.

Timing Resolution ... 17

2.

SiPM Detectors and integrated Front-End .... 21

2.1.

SiPM detectors ... 21

2.1.1.

SiPM structure ... 22

2.1.2.

SiPM electrical model ... 25

2.1.3.

S13370 (VUV4 generation) SiPM ... 27

2.2.

Application Specific Integrated Circuit (ASIC) ... 28

2.2.1.

Typical ASIC structure ... 29

2.3.

Analog channel structure ... 32

2.3.1.

Slow output path ... 36

2.3.2.

Fast output path ... 39

3.

Preamplifier Design ... 43

3.1.

Motivations and goals ... 43

3.2.

Current buffers analysis ... 44

3.2.1.

Regulated Common Gate (RCG) topology ... 51

3.3.

Proposed low input impedance solutions ... 56

3.3.1.

OTA possible solution ... 58

3.3.2.

FVF adaptations ... 62

3.3.3.

FVF pros and cons ... 68

3.4.

FVF topologies ... 69

3.4.1.

FVF-Half OTA combination ... 73

3.4.2.

Half OTA design choices ... 74

3.5.

Results ... 76

4.

Trigger circuit for fast path ... 82

4.1.

Comparator specifications ... 82

4.1.1.

Current comparators solution ... 82

4.2.

Current Steering Comparator choice ... 85

4.2.1.

Current Steering Comparator working principles ... 86

4.2.2.

Current Steering proposal and justifications ... 88

4.2.3.

Internal OTA choice and design ... 90

4.3.

Current threshold generator ... 92

4.3.1.

Linear MOS divider ... 94

4.3.2.

Fully Differential Current Collector ... 95

5.

Conclusions ... 98

List of Figures ... 100

List of tables ... 103

Introduction

Modern medicine has made huge strides in the field of the imaging diagnostic. The various imaging systems that have been developed over the years are oriented to acquire two kind of information: morphologic or anatomic on one hand, metabolic or functional on the other.

Among these, magnetic resonance imaging (MRI), computed tomography equipment (CT) and ultrasound systems are designed to collect information about morphologic alteration. The main characteristic of these systems is their excellent spatial resolution, that in MRI case can be of few hundred micrometres.

Even if the imaging is very precise, this kind of systems are sometimes unsuitable to provide information about changes in the functioning system of the human organism.

On the other side, the nuclear imaging devices are based on the detection of a radiation due to the decay of a radioactive tracer that is injected into the organism. Such systems provide information about functional alterations: indeed, the molecules that form human organs and tissues are altered before the organ shape is changed. In this way, the diagnosis is made before the tumoral mass appears.

These systems identify those areas where the tracer is concentrated and construct an image which maps the local position of the tracer distributions.

PET (positron emission tomography) and SPECT (Single Photon Emission Computed Tomography) are examples of the systems we have just described. They have in common the same working principle: to detect gamma rays () generated by radioactive decay of the injected radiopharmaceutical (a molecule bound to a radioactive isotope). At the same time, these systems have a different detection method:

PET systems employ radiopharmaceuticals that after the decay (annihilation event) generate a pair of gamma rays that travel in opposite directions (180° apart from each other), using several arrays of detectors. On the other hand, SPECT systems use a detection technique that states the acquisition of a single  photon, from which the name Single Photon Emission Computed Tomography, employing up to three detectors to cover the full solid angle.

In the nuclear imaging field PET systems provide a better quality in the detection process, because of their highest sensitivity associated to the lack of the collimator. Moreover, these systems can count on temporal information, thanks to the detection of the gamma ray pair, and as consequence they have a very good spatial resolution. Thanks to these characteristics, PET systems are able to represent very detailed images of the human organs.

Nevertheless, modern PET systems have a limit in spatial resolution. Even if the research and development in the detectors field are making significant improvements, the spatial resolution limit has to be attribute to the detector itself.

However, the only improvement of the detection systems is not sufficient to obtain high performances in PET systems.

In the last generation of PET systems, the electronic that takes and processes the information has to follow this development in order not to ideally degrade the incoming signal characteristics.

The front-end has to be designed to be the fastest as possible, so as to reduce: the time jitter in triggering the events and the dead time between two consecutive events. In this way the distortion will be minimum, assuring the best image reconstruction.

An improving in the electronics does not allow only a better and higher data quantity acquisition, but it also allows to reduce the doses of radiopharmaceuticals that are injected in the subject, without excessively hurt the patient health.

Another important point in the electronic design consists in front-end capability to implement additional functions that are able to generate more information about the events. For instance, in recent years, TOF technique has been developed aiming at providing an accurate time information about the detected event. Such an upgrade made front-end systems capable to reduce the errors localizing the corrected place where the event has been generated.

In conclusion we can affirm that, to exploit the improvements that have been made in the detectors field, is mandatory to push as much as possible the development of the frontend electronics.

1. Time-Of-Flight PET

In this first chapter we will introduce the reasons that led to use time of flight (TOF) information in PET diagnostic imaging and its working principles. We will start describing physical principles of PET systems and we will introduce  rays and their interaction with matter. Conversion methodologies are presented to introduce scintillator crystals, that are the bridge between the nuclear events and our electronics. In conclusion a description of TOF PET working principle and its advantage over traditional PET scan.

1.1. Biophysics in PET

The principle in PET imaging consists of injecting intravenously a radiopharmaceutical that travels in patient body moving through the proximity of the pathology.

Radiopharmaceuticals are a group of pharmaceutical drugs containing radioactive isotopes. Isotopes distribute within different tissues according to the carrier molecule and emit a positron. The reason why positrons are important is that, when they are injected, they move in a full electron environment that favours the annihilation event. This is an event that takes place when an electron and a positron collide. At low energies the result of the collision is the annihilation of the electron and the positron and the creation of two, or more,  photons. The most common case is the creation of two  rays that, according to the energy conservation principle and considering negligible positron and electron kinetic energies, would have each the same energy of electron or positron equal to 511keV.

For what concerns the conservation principle of the momentum, under the correct hypothesis that positron and electron collide with a very low momentum, the two photons will be emitted at opposite directions, almost 180° to each other. Fig 1.1 helps us to clarify this phenomenon effect. This directionality provides the mechanism to localize the origin of the photon.

1.1.1. Radiation origins and interaction with matter

Some medical imaging techniques, more precisely X radiography, X tomography, SPECT and PET, have in common the use of ionized radiation, X rays for the first two and  rays for the last.

Although generally X rays have lower energy than  rays, an energy interval exists in which both radiations coexist, therefore they can not be distinguished looking at their energy level but rather at their origin. Both rays are generated by atomic transactions, but while X rays are generated at electron level,  rays are generated by nuclear emission.

X rays can be generated by three phenomena, which are fluorescence, Bremsstrahlung and synchrotron light. However, we are interested in  radiation, so, leaving aside the description of X ray origins, having already described positron annihilation, which is at the origin of  rays, we will now describe the three main processes governing the radiation interactions with matter.

Actually, it should be mentioned that another phenomenon generating  radiation exists, and it is properly used in SPECT systems. This phenomenon is called radioactive decay, but it is not a topic of this thesis project so we will not enter in details.

As previously said, there are three main processes, fig 1.3, through which radiation interacts with matter, that is to say photoelectric absorption, Compton scattering and pair-production [1].

All these processes involve the transfer of the partial, or the complete photon energy to electron energy. After this transfer the photon completely disappears (photoelectronic absorption, pair-production) or is scattered.

In photoelectric absorption the photon disappears, all photon energy is absorbed, and a photoelectron is ejected from the atom.

Figure 1.1: Electron-Positron annihilation effect. The collision between a positron + and an electron e- results in an

1.1 Biophysics in PET

For  ray energies or little higher, few hundred of keV, the photoelectron, with high probability, is originated by most internal shells transferring the energy in according to the relation:

𝐸𝑃ℎ𝑒− = ℎ𝜈 − 𝐸𝑏 (1.1)

Where Eb is the binding energy of the electron, that can be considered negligible compared with the photon one, so we can assume that all energy is transferred to the generated photoelectron.

However, the generation of a photoelectron would ionize the atom positively, leaving an empty space in the shell that can be filled capturing another electron from the medium or through the rearrangement of the electrons from most external shells. This can occur through a radiating process (Fluorescence) or no-radiating process (Auger Effect).

During radiating processes, one or more X rays can be generated. Normally they may not be a big issue because mostly they are reabsorbed by photoelectric absorption close to their emission site, but if they ever escape from the radiation detector it would influence the response. Photoelectric absorption is the characteristic way through which X rays an  rays with low energy interact with matter.

Figure 1.2: Compton scattering event illustration from [1]

The second way through which the  photons interact with matter is the Compton scattering phenomenon fig. 1.2. This is the commonest interaction with matter mechanism of the energy radiation typical of radioisotope sources. In this mechanism the incident photon is scattered by one of the electrons in the most external shells of the atom, the photon is deflected of an angle  compared to his initial direction. During this interaction a portion of the photon energy is transferred to the electron, that takes the name recoil electron, that has been set free.

The resulting energy of the recoil electron is equal to the difference between the energy that the photon had before the collision and the corresponding scattered photon energy, in according to the relation:

𝐸𝑃ℎ𝑒− = ℎ − ℎ

1 +_𝑚ℎ

0𝑐2(1 − cos)

(1.2)

Where the second term of the equation corresponds to the photon scattered energy and, as we can see, it depends on the deflection angle .

As said in [1] the probability of Compton scattering event depends on the number of the electrons in the atoms, that means this probability grows up with the atomic number Z. One can derive the angular distribution of the scattering event using the Klein-Nishina formula (1.3) for the differential scattering cross-section (𝑑𝜎

𝑑Ω), or better the scattering rate at angle  per incident

radiation into a differential solid angle d:

𝑑𝜎 𝑑Ω= 𝑍𝑟0 2₍ 1 1 + 𝛼(1 − cos )) 2 (1 + 𝑐𝑜𝑠 2_ 2 ) (1 + 𝛼 2_{(1 − cos )}2 1 + 𝑐𝑜𝑠2[1 + 𝛼(1 − cos )]) (1.3) Where  = ℎ

𝑚₀𝑐2 and 𝑟0 is the electron radius.

The final kind of interaction with matter is the Pair Production.

This mechanism is energetically possible only if the photon energy exceeds twice the rest-mass energy of the electron (1.02MeV), but the probability of occurrence remains very low until the energy approaches several MeV.

1.1 Biophysics in PET

The result of this interaction is the disappearance of the photon and formation of an electron-positron pair. All the energy, that exceeds the required amount to ensure that the pair production occurs (1.02MeV), is transformed in kinetic energy and shared by the electron-positron pair.

It is important to pay attention that the created positron can annihilate with an electron, thus leading to the creation of two annihilated photons that can have important effects on the detector response.

The following figure 1.3 represents the likelihood that an event occurs rather than another depending on the radiation energy and the atomic number Z.

Figure 1.3: Interaction process as function of incoming radiation energy. The marked lines represent the energy in which the close effects have the equal probability.

1.2. Gamma ray detectors

We can distinguish two broad classes in which we can collocate different typologies of  ray detectors, the first one is the spectrometer class, the second one includes all those detectors that perform the  ray imaging [2].

Spectrometers are detectors that can measure the interactions of photons with matter as a function of the photon energy. They involve energy loss due to the absorption of the radiation by the medium. These detectors are used to reconstruct the spectrum of the  ray source just plotting a histogram of the sampled data collected event by event.

To understand the result of the analysis, it is important to realize that only a fraction of the gamma rays interacts with the medium, and all kinds of interactions presented in 1.1.1 occur. The whole incoming radiation is absorbed mostly by photoelectric and Compton scattering, and it composes the effective readout of the detector, constituting the main peak called photopeak.

Unfortunately, in the medium there are a lot of secondary sources of energy loss, so in our spectrum we can see different peaks, even if with less amplitude compared to the photopeak ones.

The results of the information that can be recovered by the energy spectrum analysis are showed in figure 1.4.

Since each Compton interaction involves an energy loss, which depends on the scattering angle, and in normal conditions all the scattering angles will occur in the detector, an energy continuum distribution is usually visible in the spectrum, constituted by Compton scattered photons not absorbed by the medium.

The lower energy peaks are caused by X rays that manage to escape, the Compton scattered photons that are not absorbed by the medium (called backscatter peak), or are caused by the electron positron annihilation that generate a  rays 511keV pair that results in a peak of the energy equal to the difference between the energy of the original incident photon 𝐸 and 511keV or

1.022MeV if one or both annihilated generated photons manage to escape [3].

The second class of detectors, as said above, includes all those typologies of detectors that perform the task of  rays imaging. This kind of detectors also relay on the interaction of the  ray with matter to identify from which direction the radiation arrives.

We can split this class of detectors in two subclasses on the technology they use to detect the radiation.

They can be:

Pixel detectors, in which the image is reconstructed with a resolution equal to the pixel resolution.

Continuous detectors, in which every point in the revelation area can be taken as a coordinate of interactions.

Figure 1.4: An example of the energy spectrum shows the different peaks that can be observed.

1.2 Gamma ray detectors

There are different parameters that we have to take into account when we speak about radiation detectors. These parameters will be presented with a description of their meaning.

The first important parameter is the Sensibility, that consists on the capability of the detector to produce the electric charge per energy of incident photon. The energy absorption of a photon, through one of the mechanisms previously seen (photoelectric absorption, Compton scattering, pair-production), produces an electron-positron pair. This pair generation produces a chain of ionizing interactions, the absorption of possible fluorescence radiation (X ray) and scattered photons produces these ionizing interactions too. The quantity of charge so generated is proportional to the total energy absorbed.

The second parameter of merit is Revelation Efficiency. This is the result of the product of three factors, geometrical efficiency, absorption efficiency and photopeak efficiency:

𝜂𝑅 = 𝜂𝐺𝜂𝑎𝑠𝑠𝜂𝑓𝑝 (1.4)

Where 𝜂_𝐺 is the geometrical efficiency. When the radiation hits the surface of the detectors, even if a source radiates along all directions, only a portion of that radiation will fall in our detector. The bigger is the surface of our detector the bigger is the geometrical efficiency.

It is identified as: 𝜂𝐺= Ω 4𝜋 with Ω = Α 𝑑2 (1.5)

Where  is the solid angle with which the source of radiation sees the detector surface. The absorption efficiency counts the quantity of photons that enter in the detector and are effectively absorbed. The probability of a photon to be absorbed depends on the parameter , that is a feature of the material and it is inversely proportional to the incident energy. This means that if the energy increases  decreases and also decreases the probability of absorption. So, we can say that more energetic is the radiation harder is its absorption.

Finally, the photopeak efficiency corresponds to the fraction of photons that effectively have reached the detector and that have lost all their energy (have been completely absorbed).

1.2.1. Imaging Detectors

Depending on the physical phenomena that have been used to obtain the imaging data, there are different modalities to measure and visualize different characteristics of the investigated object.

As already said, in X-rays and computer tomography (CT), which include PET a d SPECT, it is observed the attenuation of the radiation in a medium.

the patient; instead, in SPECT systems, one, two or three large area detectors are used and rotated around the patient to cover the full solid angle.

Currently, almost all commercially imaging scanner systems employ scintillation crystals with photomultipliers (PMT) to read the nuclear radiations.

The most common material used in Anger Camera technology is sodium iodine NaI, a scintillator that converts  rays in visible light, while modern PET systems employ lutetium oxyorthosilcate (LSO), gadolinium oxyorthosilicate (GSO), or bismuth germanate (BGO) [4].

Behind the scintillator crystal, to read the just converted signal, proper photodetectors have to be placed, followed by good electronics with an adequate analog to digital conversion.

The electronic stage must guarantee different results, first of all the correct identification of the signal. It has to be capable to identify the “good event”, this means that a pulse analysis is required to quantify the photon energy. In this way it is possible to eliminate all the undesired events, such as backscattering, or secondary positron annihilation, already mentioned.

Another task of the electronics is to keep track of the number of “good” pulses and their coordinates. This allows us to have recorded information about the amount of radiation and the directions from which the photons arrive. These are fundamental data for the image reconstruction.

Additionally, in TOF PET systems, it is important to have a timing stage to assign coincidence pairs, fit to better locate the point where the pair emission takes place.

1.2.2. SPECT Detectors

Speaking about Nuclear Medicine imaging, we can spend a few words about Single Photon Emission Computed Tomography (SPECT) and Anger Camera, the detector used in this imaging system.

As already said, SPECT systems use from one to three detectors, which are rotated around the patient, to acquire the information of the distribution of the detected photons from a number of angles [4].

Since the purpose of the system is to detect from which direction the photons arrive, a thick lead collimator is used; in this way all non-perpendicular photons are absorbed by the collimators, in figure 1.5A is showed a gamma camera scheme.

It is obvious that, even if ideally, only the normal photons are admitted, in reality all those photons that are contained in a cone corresponding to the acceptance angle for each collimator hole would pass.

1.2 Gamma ray detectors

Is important here to refer to spatial resolution, that parameter which, as said in 1.3.1, identifies the origin point of the photon.

For the reasons above, the spatial resolution cannot be precise, but instead it has to be taken as a spatial distribution that can be described by a gaussian function. This is due to the fact that, not only the photons coming from the same point are red, but also the ones from near spots that have a radiation angle directed to the same hole collimator. This means that the resolution of the camera is distance-dependent, indeed it deteriorates as the distance between the source and the collimator increases [4]. It contributes to lower the sensitivity too, owing to the reduced number of photons that can reach the crystal detector.

To improve the spatial distribution, and then the resolution, it is possible to increase the collimator length. However, there is a marked trade-off between resolution and sensitivity, in fact the larger are the collimators the lower their sensitivity will be, or rather, they would hardly let photons that are directed to the scintillator pass. Undoubtedly the larger is the collimator the higher will be the probability of a photon to be absorbed by it instead of the scintillator, figure 1.5B shows two different kinds of collimator septa.

Beside the limit related to the acceptance angle of the collimators, the sensitivity of the detectors also depends on the efficiency of the detector material, that is connected to the radiation energy. In fact NaI crystals, used in SPECT, provide high sensitivity (80-90%) (depends on the thickness of the crystal) only for photons with relatively low energy (140keV), for energy around 511keV the sensitivity is lowered (25-28%).

Figure 1.5: (A) A simplified scheme of a gamma camera shows how only photons parallel to the collimator septa can reach the scintillator. (B) Two different

Even if the crystal efficiency is high enough, in any case the amount of the light produced in the visible range, by the scintillator, is still quite small.

The purpose of the next stages, constituted by photodetectors, is to convert the visible light into an electric signal and amplify it.

1.2.3. PET Detectors

As already known positron annihilation in Positron Emission Tomography (PET) generates a  photon pair with 511keV each, emitted in almost exactly opposite directions. To detect these photons, the acquisition chain of a PET system has to start with several rings of detector that surround an imaging volume.

PET systems use the coincident detection of two  rays to collect projection data from which tomographic images are reconstructed. For this reason, in PET detectors no lead collimators are necessary, because the source of the photons is supposed to be along the line that connects the two detectors that have collected the two opposite rays.

When a photon hits one of the detectors a time window (5-10ns) [5] is opened, and all the other detectors are checked for the detection of the second photon of the pair. If this detection is taken inside the time window a coincidence event is counted, otherwise the event is rejected.

If a good event is recorded, a line connecting the two detectors is create, this is called line of response (LOR).

In 1.3, it will be explained how this information is used to reconstruct the image.

In modern PET systems, since the number of detectors composing the ring has grown, checking each event for coincidence with all other detectors would be impossible. The solution is to give to each event a time-stamp and to send it to a digital coincidence processor that, during a fixed time, tests for coincidence every other event, comparing the differences in detection time in a coincidence window, and only then the LORs coordinates of each detected coincidence will be calculated.

Although currently, the resolution of PET imaging systems is better than SPECT ones, the physics of the positron annihilation and the detection of the radiation results in a lower accuracy in identifying the precise site where the emission occurs. For this reason, to respond to the weakness in PET scanning, Time of Flight PET systems have been developed.

Since the injected positrons have a certain kinetic energy, the relative position between the annihilation site and the location from which the photons are emitted may not be negligible, resulting in an uncertainty in source origin. In addition, it was experimentally seen that the two emitted photons are not emitted perfectly in two opposite directions, but they have a phase shift of more or less 0.3° [4].

1.2 Gamma ray detectors

The signal coming from the detectors can be processed in two different ways.

The first way is the acquisition of the image in 2D mode. To allow this, a thin annular lead, to shield adjacent rings, is placed axially between two rings. Its purpose is to prevent radiations incoming at a large oblique angle, thus enabling only the detection for the detectors of the same ring, or at least between two adjacent rings, isolating all the other rings. The acquisition in 2D is usually adopted in brain scanning.

The second way, the most modern one, is the 3D mode acquisition, where no interslice septa is used, and the interactions are allowed for all detectors of all rings.

This acquisition method increases the sensitivity of the imaging system, boosting the coincidence count and so improving the statistical quality of the acquired data.

3D mode has its own pros and cons, the price of this improvement is high, in fact stricter specifications in electronics are required.

1.3.

PET structure and principles

Positron Emission Tomography (PET) is a type of nuclear medicine procedure that measures metabolic activity of the cells of body tissues. PET is actually a combination of nuclear medicine and biochemical analysis. Used mostly in patients with brain or heart bad conditions and cancer, PET helps to visualize the biochemical changes taking place in the body, such as the metabolism [6].

As said in 1.2.3, PET works by using a scanning device (a machine with a large hole at its centre, see figure 1.6) made by several rings. Each ring is formed by a circular array of photodetectors whose purpose is to detect the photons pair emitted by the positron and electron annihilation.

To read correctly the emitted photons, PET scanner consists of several rings of detectors placed around the patient. According to what was said previously, opposite detectors have to read opposite  photons, as described in fig 1.2. If the time difference between two detected photons is included in the set time window (whose typical values are between 5-10ns) it means that the two events can be considered originated by the same annihilation event. A time window larger than 10ns would promote wrong acquisition of the photons, owing to the presence of different  rays emitted from different annihilation events.

What the traditional PET does with this kind of information is to identify the line, called Line-Of-Response (LOR), already mentioned in 1.2.3, where the annihilation occurs, but it can not predict the correct position on this line because all points along the line have the same probability of being emitted [7]. To correctly identify the infected area, it would be necessary to reconstruct the image. To do this, two main mathematical algorithms exist: an analytic one and an iterative one.

[7]The most common analytic reconstruction technic is called Filtered Back-Projection (FBP). This technic uses virtual LORs between every detectors pair, taken at each angle from 0° up to 90° or 180°, to generate a sinogram. During the diagnostic phase, every real annihilation event will be mapped and stored on its corresponding place in the sinogram. These data will be next filtered and then processed to obtain the image reconstruction [8].

1.3.1. TOF PET

TOF information is used to better locate the annihilation event position on the LOR. Ideally, if we would have a perfect detection of the flight time of two emitted  rays, we could identify the exact point where the annihilation event takes place, by means of the difference between the two arrival times. All position information, thus obtained are stored into a matrix, which will be later used as a reference for the image reconstruction.

Figure 1.7: Single ring of radiation detectors. Two  rays are detected by two opposite

1.3 PET structure y principles

The time of flight theory is based on the concept of perfect location, in fact, if we have an ideal TOF system that provides a perfect information, it is possible to identify the location of each annihilation event, based only on TOF information, without the need to provide any image reconstruction.

Unfortunately, due to imperfections in timing information and consequently positional uncertainty, this is an impossible result [9] and an imaging operation would be necessary in order to identify the infected zone.

We have already described in 1.2 the parameters of merit that we have to take into account for a correct tomography:

Sensibility: capability of the imaging system to identify the parameter of interests. Sensitivity: ability of the imaging system to correctly distinguish the parameter of

interest from other possible signals.

Resolution: capability of the system to detect very close details in the distribution.

o Spatial Resolution: capability to correctly identify the origin of the event.

It is defined by a parameter: Point-Spread-Function (PSF), that is the pulse response of the system, that means that the response of the system has to be as similar as possible to a pulse for a given input pulse stimulus. Spatial resolution is defined as the Full-Width-Half-Maximum (FWHM) of the PSF.

The TOF PET concept started to be developed during the 1980s, but unfortunately, because of the limited stopping power of the scintillators used in those years (𝐵𝑎𝐹₂ and 𝐶𝑠𝐹), TOF systems had limited spatial resolution and sensitivity and so they were not competitive with nonTOF systems that used high density scintillator bismuth germanate (BGO), which was developed in the same period.

With the development of the new scintillator crystal, lutetium oxyorthosilicate (LSO), TOF systems began to gain increasing importance and the interest of the researchers. Such scintillator has a very good timing resolution, excellent specifications for stopping power and a good energy resolution [10]. Beside the LSO, in the early 2000s, another very performing scintillator was introduced, the lutetium-yttrium oxyorthosilicate (LYSO). Thanks to such development in scintillators field it brought to the first commercially available TOF PET scanner in 2006.

In most modern years, the most current clinical TOF PET scanner system has a time of flight resolution of about 400ps. The system time resolution is determined by the different components that are involved during the detection process which are: the scintillator crystal, the photodetector and the electronics (figure 1.8).

An important factor that affects the correct timing operation is the noise. In literature it has been pointed out that TOF PET images, compared to equivalent no-TOF images, show lower levels of noise and better resolution.

In fact, as we can see in [5], using back-projection technics to reconstruct the image, the TOF results in a reduction of the noise propagation along the LOR. Considering the reconstruction technic, mentioned above, we can take the signal-to-noise ratio modelled by Strother in [11]. Here, in a cylinder of diameter D, with uniform distribution activity, one can estimate the SNR in an image element, where an image element corresponds to a slice of the LOR, of size d starting from the data stored in the projection space:

𝑆𝑁𝑅 = 𝐶 ∗ 𝑛−12_{∗ [} 𝑇 2 (𝑇 + 𝑆 + 𝑅)] 1 2 (1.2)

Where C is a constant, T represents the total true event in the image element, S the scatter events, R the random coincidences, and n the number of slices (volumes) in the LOR influencing the noise.

In order to reconstruct the image, the line of response is divided in many small cubic volumes having d size. Each volume contains the information about a position along the LOR. In no-TOF PET all the slices along the same LOR contribute to the same data, so it would be back-projected in all image elements of the LOR. This means that all the volumes of size d, are contributing to the noise (figure 1.9a). In this case we can estimate:

𝑛 =𝐷

𝑑 (1.3)

Figure 1.8: A TOF PET is a combination of fast scintillator, good readout hardware and accurate reconstruction and corrections functions.

1.3 PET structure y principles

In TOF PET each data is back-projected only in the element position corresponding to the event and into adjacent volumes (figure 1.9b) in according to the spatial resolution x. x corresponds to the spatial uncertainty, related to the time resolution t by the equation:

𝑥 = 𝑡 ∗ 𝑐/2 (1.4)

In this case we can estimate n = x/d resulting in SNR terms:

𝑆𝑁𝑅 = 𝑆𝑁𝑅𝑛𝑜−𝑇𝑂𝐹 ∗ √

𝐷

𝑥 (1.5)

If an iterative reconstruction method would be used, the above equation could only roughly estimate the SNR gain [5].

1.3.2. Timing Resolution

The system time resolution is the key parameter that defines a TOF PET scanner. Considering two photodetectors of the circular array, the time resolution can be measured by placing a point source between these detectors and measuring the Full Width Half Maximum (FWHM) of the distribution of the time of flight difference between the two detectors.

The system time resolution results from the average of the individual time resolution associated with each detectors pair.

Figure 9: (a) In non-TOF reconstruction, all volume elements n found in the object along the line of response contribute to the noise in each image element. (b) In TOF

reconstruction only the volume elements n adjacent to the position identified by the measured TOF contribute to the local noise

The effectiveness of the time resolution can be easily understood by looking at the limited uncertainty with respect to the line of response (LOR) (figure 1.10). In fact, while in non-TOF systems the uncertainty is distributed along the entire LOR, in TOF system the uncertainty is limited to the single section corresponding to the time distribution.

Timing resolution in PET systems is important to correctly determine the spatial resolution, significant in TOF PET information, that is used to locate the emission point (figure 1.11). This is determined by different factors: the scintillator, the electronics and the kind of sensor that has been used, which can be photomultiplier tube (PMT) or silicon photomultiplier (SiPM), the latter being the one actually more used because it allows them to reduce timing resolution below 400ps [10]. Since time resolution represents the variability in the signal arrival times for different events, it needs to be properly accounted for when detecting coincidence events.

Figure 1.11: Real annihilation event location. Calculating spatial resolution is possible to identify the position from the centre in which the

Figure 1.10: in non-TOF system on the left the uncertainty is distributed along the entire LOR, in TOF system on the right the uncertainty is adjacent only to the infected

1.3 PET structure y principles

In this thesis project we will adopt a SiPM detector, so as to focus on SiPMs based system characteristics. The working principles and characteristics of PMT will not be taken into account. To derive the time resolution t, as already said, it is necessary to calculate the full-width-half-maximum (FWHM) of the distribution of the time of flight difference, generated from a point source between detectors A and B. Actually, as said above, timing resolution depends on different factors, so it can not be calculated looking at a single distribution, but it has to be an average of the individual time resolutions for each detectors pair surrounding the patient. Moreover, each t can be affected by variation in electronics and detector quality.

Figure 1.12 shows a block diagram that exemplifies the entire timing process [12]. The signal generated by the SiPM is readout by the preamplifier, which feeds a fast path and a slow path. In the fast path a low noise leading edge discriminator delivers the time information to a logic block. In the slow path there is an integrator for the energy discrimination, which delivers the information to the same logic block. The logic block will process the information, selecting only those events representing the whole energy of 511keV of a gamma ray. Those signals that do not respect such characteristic will be neglected.

We can describe now all those factors that influence the time resolution as Conti does in [5].

First: a great limitation in time resolution is given by the material by which the photons are absorbed, those materials are called scintillators.

What is important to know is that, the time resolution is controlled by the rise time, the decay time and the absolute light output. Generally, to have a short rise time is not a problem, so, to improve time resolution, it is important to use materials that have a fast decay time and a high light output.

Another important cause of the reduction in time resolution is due to the assembly of scintillator crystals in arrays. These are mounted on a light guide coupled with a group of light sensors with which they share the light. The time resolution deteriorates as a result of the light sharing between different crystals, the light guide and the sensors.

Another element worsening the time resolution is the electronics by which the signal is read. Time jitter in the comparator due to the electronic noise, time slewing due to small amplitude signal and no optimal time to digital converter (TDC) design, could be the major limitations in time resolution. Thanks to proper design it is possible to minimize the electronics effect.

2. SiPM Detectors and integrated Front-End

This chapter will be divided in two main topics. First, we will deal with Silicon Photomultiplier (SiPM), a class of photodetectors, that have recently been introduced in TOF PET to overcome the shortcoming of PMTs and APDs. We will describe their structure and how they can be modelled electrically. Further we will present the SiPM that our preamplifier development uses as a reference. The second part focuses on the presentation of the Front-End electronic, giving more emphasis to the analog channel part.

2.1. SiPM detectors

Silicon photomultiplier devices (SiPM) consist of an array of microscopic parallel connected avalanche-photodiodes (also called Single-Photon-Avalanche-Diode, SAPDs) working in Geiger-mode (G-APD). This mode takes its name from the Geiger-Mueller counter principle [1], whose structure consists of a central cathode and a gas surrounded by an anode. Anode and cathode are polarized at a high voltage; this solution aims at exploiting the gas multiplication to greatly increase the charge generated by ionizing radiation.

In a similar way, the micro-cells composing the SiPM are biased at high voltages above the breakdown, resulting in a proportional avalanche charge-generation within the diode depletion region, and thus in a great current generation, every time a visible photon from the scintillator hits the depleted region.

The fundamental differences between Geiger-mode APDs (SPADs) and classical APDs is that, while in usual APDs only the electron carriers can trigger the avalanche and the current is proportional to the triggered number, in Geiger-mode both electrons and holes can provoke a self-sustain avalanche [13].

Thanks to their advantages, SiPM have acquired more relevance in PET scanning systems, arriving to replace, in many applications, the already consolidated PMTs.

In the following list we will present the main features that show principal advantages of the SiPMs compared to PMTs [14] [15]:

•

Similar to the PMTs, SiPMs present a high internal gain, this is translated in a strong reduction of the impact of the electronic noise from the analog front-end.

•

A fast response when detecting radiation makes the SiPMs comparable with PMTs and allows to use them in timing application like TOF-PET.

•

Unlike the PTMs, SiPMs are more compact and can use much lower bias voltage, 25V to 28V, depending on the breakdown of the detector, instead of the thousands of volts required by a PMT.

•

SiPMs have higher photon detection efficiency than PMTs for scintillator light. The bests PMTs have a detection efficiency of about 35% vs SiPMs ones that is about 50-60%, this larger signals with less stochastic and Poisson noise.

•

SiPM are more robust and, contrary the PMTs, can operate into environments with a strong magnetic field, such as in magnetic resonance.

•

Their mechanical robustness and compact size allow to design very small and tough modules.

A disadvantage of SiPMs compared to PMTs is a higher dark current rate per unit area. Anyway, thanks to all those advantages, SiPMs are efficiently contesting with PMTs detectors.

The fig 2.1 shows the differences between PMT and SiPM sizes.

2.1.1. SiPM structure

As already said at the beginning, a silicon photomultiplier is constituted by an array of parallel photon counting microcells. In order, for the large amount of current, not to destroy the diode, an appropriate quenching circuit is integrated to each microcell.

There are two kinds of quenching circuits that can be implemented: passive quenching or active quenching [13].

2.1 SiPM detectors

Typically, a quenching resistor with several hundreds of K is integrated and connected in series with each microcell, to prevent that a large generated current, due to the avalanche, destroys the diode (fig. 2.2). In fact, every time a large current will be generated during the avalanche process, the voltage drop between the resistor terminal increases reducing the breakdown voltage across the microcells and stopping the avalanche action.

Due to the intrinsic microcell capacitance and the parasitic capacitance of the overall summing network in a SiPM, each microcell needs a large amount of time to recover the work condition, after that the breakdown voltage has been restored. The time magnitude is typically equal to the product between the quenching resistor value 𝑅_𝑞 and the overall SiPM capacitance 𝐶𝑆𝑖𝑃𝑀.

𝜏𝑟𝑒𝑠𝑡= 𝑅𝑞𝐶𝑆𝑖𝑃𝑀 (2.1)

Even if the most of the commercial SiPMs use passive quenching due to its easy implementation, there are some shortcomings that limit the performances. The shape of the pulse is affected by all those changes that can be subjected to the passive element, such as the temperature. SiPM, as said above, can also be affected by a long recovery time, due to the quenching resistor value.

To overcome these defects, active quenching circuits can be implemented. Fig 2.3 from [13] can help to understand active quenching structure and its operation.

Figure 2.2: On the right the microcell size equals to the pitch between microcells and is either 15 µm, 25 µm or 50 µm. On the left a SiPM equivalent circuit with quenching resistor

Figure 2.3: An example of the structure of passive quenching circuit on the left and of active quenching circuit on the right.

A generic active quenching circuit consists of a comparator-based feedback circuit, with a sensing and a quenching network. Whenever a current is generated and sensed at the output of the SPAD, a feedback pulse is produced and it forces the SPAD to quench the avalanche by means of a controlled voltage source. A plus of this source is its configuration to reset the SPAD at its active state by restoring the voltage drop at its terminal. This active quenching circuit results advantageous in all those applications where a very high counting rate is required.

A single SPAD can grant information only about light detection, but not about any data regarding the light intensity. To find a proportionality between the incident light and the total charge, all micro-cells are connected to the same readout line to sum every individual signal generated by each SPAD.

A big limitation, that we have to take into account, is the trade-off between dynamic range and light detection efficiency. In [13] this dependence is explained.

The SiPM photon detection efficiency (PDE) is defined empirically by the expression:

𝑃𝐷𝐸 = 𝑄𝐸 ∙ 𝐹𝐹 ∙ 𝐺𝑃 ∙ 𝑅𝑇 (2.2)

Where FF, the Fill-Factor, is the ratio of the active to total area of the SiPM, GP is the probability for Geiger discharge in any micro-cell and RT is the factor related to the recovery time of each micro cell.

Considering a SiPM with a fixed dead area (area dedicated to the interconnections) for each SPAD, a large number of micro-cells implies a small micro-cell sensitive area. This reduced area decreases the PDE for each micro-cell and thus it reduces the overall device detection efficiency.

2.1 SiPM detectors

At the same time, a large number of micro-cells are needed by SiPM to respond linearly to the incident light stimulus, that would increase the dynamic range.

In SiPM, as we already said, an avalanche phenomenon is triggered by an electro-hole pair. This pair can be generated not only by a hitting photon, but also by thermal agitation and tunneling effect. This pair generation produces an undesired current, that is called dark current, and corresponds to that current that is always generated, even in absence of light. Undesired dark current results in an increased noise level.

2.1.2. SiPM electrical model

As said in 2.1.1 each SPAD provides only information on where a photon is received, without giving any one about the light intensity. Since the SiPM is built by connecting all micro-cells that share the same anode and cathode, the total amount of current at the output of the detector is the sum of each current contribution from each SPAD. All these micro-cells composing the SiPM are therefore connected to a unique readout circuit, so that the current pulse of each SPAD overlaps creating the detector current signal (fig 2.4).

Figure 2.4: On the left the output pulses from each SPAD. On the right is shown the resulting SiPM output signal.

The electrical model associated to the single SPAD, and the one associated to the SiPM are shown in fig. 2.5. For the sake of simplicity, having aready said that the overall current signal corresponds to the sum of each SPAD pulse, we will use the simplest SPAD model to characterize, with a good accuracy, the output pulse of the SiPM [16].

If we consider a single SPAD, that, in the most common case is implemented with a passive quenching circuit, it is modelled as a parallel connection between the internal resistance of the diode space-charge 𝑅_𝑑 and the inner depletion layer capacitance 𝐶_𝑑. The quenching circuit is represented by the parallel between the quenching resistor 𝑅𝑞 and its parasitic quenching

The hitting photon is modelled by a close switch when it triggers the avalanche, and it is open during the quench effect. A further parasitic capacitance 𝐶_𝑚, between the two terminals of the cell, has to be considered as the model to take into account the metal grind path spanning over the entire surface of the semiconductor device. The difference between the anode-cathode voltage 𝑉𝐾 and the breakdown 𝑉𝐵𝐷 is called overvoltage and corresponds basically to the voltage drop of

the internal node, due to the relation 𝑅_𝑞 ≫ 𝑅𝑑.

To characterize the SiPM current pulse we have to analyse what happens when a photon triggers an avalanche. When the switch is open (so we do not still have radiation) 𝐶_𝑑 is charged to the SiPM bias voltage 𝑉𝐾. When a photon detection occurs the switch in fig 2.5 is closed,

leading to a discharge of the 𝐶_𝑑 capacitance through the resistor 𝑅_𝑑. The resulting current has the expression:

𝐼𝑑𝑒𝑡,𝑝𝑒𝑎𝑘 =

𝑉𝐾− 𝑉𝐵𝐷

𝑅𝑞+ 𝑅𝑑

(2.3)

The presence of the parasitic capacitance is an important factor for the fast-rising edge of the output current. It works as fast path for the charge delivered by the avalanche generation. In fact, the charge collection is dominated by this small capacitance that is discharged exponentially. The overvoltage drops with the same time constant that the detector current rises, and can be estimated as:

𝜏𝑟𝑖𝑠𝑒 = (𝑅𝑞//𝑅𝑑)𝐶𝑞 (2.4)

Figure 2.5: On the left it is presented the SPAD model. On the right the SiPM electrical model that takes into account the contribution of all micro-cells that constitute it (N_fare called fired

2.1 SiPM detectors

When the avalanche is quenched the switch in fig 2.5 is open again and the reset time can be calculated in according to the equation (2.1) considering 𝐶_{𝑆𝑖𝑃𝑀} equivalent to (𝐶_𝑞+ 𝐶𝑑):

𝜏𝑟𝑒𝑠𝑡 = 𝑅𝑞(𝐶𝑞+ 𝐶𝑑) (2.5)

A more accurate result can be obtained analysing the model on the right in figure 2.5. This simple analysis provides an accurate result.

It must be specified that, as we can see in figure 2.5, considering the entire SiPM, and not the single SPAD, the situation becomes more complex. As widely known, SiPM is the result of the parallel interconnections of many SPADs, so the resulting capacitance corresponds to the sum of the all parallel capacitances of each SPAD that composes the detector. The resulting capacitance can be really big.

In recent years, the increasing SiPM’s dimensions have been accompanied by an increasing in the capacitance size, pushing the electronics design towards even more challenging targets.

2.1.3. S13370 (VUV4 generation) SiPM

The current market is constantly investing on the development of even more performing SiPM to manage with the new challenges in TOF PET. In order to improve the photon detection efficiency (PDE) it is necessary to increase the active area of each micro-cell [13]. Actually, SiPMs are getting bigger and bigger to meet the ever-increasing request performances, but a large area is translated in a growing detector capacitance, and so it can be necessary to improve the front-end electronics, that is the purpose of this thesis project.

To choose the SiPM that best meet the specifications is an important task. Nowadays there are many SiPM manufacturers involved in researching ever new and more performing devices, such as Hamamatsu, SensL, Fondazione Bruno Kesler, and so on.

Even if the purpose of our research is to develop a first step preamplifier, able to manage with the ever growing SiPM output capacitance, in our lab the SiPM S13370 𝑉𝑈𝑉₄-MPPC (Vacuum Ultraviolet Multi-Pixel Photon Counter), by Hamamatsu, has been adopted.

𝑉𝑈𝑉4-MPPC are detectors able to detect light down to 120nm, covering scintillation

wavelength of liquid xenon (LXe) and argon (LAr) with cryogenically compatible, ultralow-RI packing options. The most interesting features of this SiPM are listed below:

• Can be operated at low voltages (<60V) in LXe • Single photon counting capability

• PDE close to 25% at 178nm • Gain larger than 2 × 106

When SiPMs are employed, to grant the same equivalent area of a PMT while keeping, for the signal readout, the same number of analog channels, the grouping of several of this SiPM is needed. This requirement makes the analog readout more challenging.

2.2. Application Specific Integrated Circuit (ASIC)

More the performances of the silicon photomultiplier have increased, more the electronics has to improve to follow this progress.

In PET systems, research is aiming to implement compact and performing acquisition chains that can manage with the output signal from photodetector rings. For this reason, it is aiming to develop ICs that are able to implement all necessary readout functions in a single chip.

This kind of ICs, that are capable of performing all operations required to a specific purpose, are called Application Specific Integrated Circuits (ASICs).

An Application Specific Integrated Circuit is an integrated circuit that has been realized for a specific use, in our case the SiPM output signal readout.

The maximum complexity in an ASIC has grown over the years, arriving to count over 100million of logic gates on a single integrated circuit. Modern ASICs can often include microprocessors and memory blocks; such an ASIC is called SoC (System on Chip). To programming an ASIC, it is usually used a hardware description language (HDL), like Verilog or VHDL.

The two primary ASIC design methods are gate-array and full custom design [17]. In the first one the costs are much lower, due to the minimal design work needed to make a working chip. These are often larger in size than full custom design ASICs, that requires a larger power dissipation. This ASIC typology can be employed in more general applications, and this is the main difference with the second one. Gate-array design is a manufacturer method in which the diffused layers, transistors, and other active devices are predefined. The job of the engineer is so limited to open and to close switches to impose the chip to work in the desired way.

To be more precise, the gate-array technology is based on partially prefabricated wafers with simple gate cells. These wafers are not customized and “held in stock”, so the task of the engineer is to place the final metallization layer with the purpose of implementing the logic function using the gate-cells. Besides the possibility for the designer to implement his own functions, gate-array technology presents library functions predefining the constructions of more complex logic functions, such as adders, register, etc [18].

However, to implement such complex functions the designer must respect the predefined positions and type of gate cells, leading to lower efficiency and performances compared to full-custom ASIC.

2.2 Application Specific Integrated Circuit (ASIC)

A full-custom ASIC has a more complex design than the gate-array. The design starts at transistor’s level up to create the overall function. This increased complexity means that this ASIC can do much more than the other one. Full-custom ASICs are designed to fulfil precise specifications, and so can be employed in very precise applications.

A full-custom chip, designed using only predefined logic cells from a specific library and physical design process provided by standard EDA/CAD tools, is called standard cell ASIC and is one of the primary ASIC technologies [18]. The reason why, standard cell ASICs performances are not so high, is due to the quality of the standard gate cells provided by the its manufactures.

Even if some limitations exist, there are different reasons for which it is advantageous to use an ASIC chip.

First of all, it has to be taken into account the small size of the chip. This is directly translated in a reduction of the electrical power, compared to the power dissipation of an acquisition chain made by standard components.

In addition, an ASIC contains only the circuitry necessary for the application, so the chip is more efficient than a classic PCB with multiple interconnected standard devices.

Another advantage is that using this chip requires fewer electronic components and it is much cheaper to assemble. This is possible because an ASIC can hold inside it many different systems.

In summary, to use an ASIC seems a convenient choice, because it can shrink the system size and the costs. It is a chip that can be implemented specifically for a particular application.

In conclusion it could be reasonable to use an ASIC to implement a proper front-end to acquire the output signal from a time of flight PET system.

2.2.1. Typical ASIC structure

Conventional front-end electronics were usually realized by combining discrete circuits on one or more PCB boards. Many readout channels are needed to properly read the signal incoming from all photodetectors in most modern PET imaging systems. In according to this, the data acquisition for each channel could occupy a big area and could have a large power dissipation.

As seen in 2.2, using an ASIC chip, it is possible to overcome these problems.

Actually, many ASIC circuits have been implemented for PET imaging only. We can name some such as NINO [19], Triroc [20], DIET [21] etc, and all of them are different from each other. Among these different ASICs the only points that they have in common are their application and their general structure.

To show how an ASIC circuit work we will describe the structure of the one that has been used in our laboratory, that is one of the most modern ASIC circuits adopted in PET imaging, the TOFPET2 by PETsys Electornics, that uses CMOS technology 0.11𝜇𝑚.

Even if the technology employed in this circuit is different from the one that we used in our research, it is of our interest to know how an ASIC chip is composed.

TOFPET2 ASIC is a new 64 channel chip produced to readout and digitalize fast signals where a high data rate and fast timing is required.

Each of the 64 channels have front-end current conveyor with two outputs, these are connected to two post amplifiers, respectively one for the timing branch and one for the energy branch.

The general architecture is based on the previous version, the TOFPET chip, but it has significant differences. In this new variant a pulse amplitude measurement obtained by charge integration is implemented.

A simplified block diagram of TOFPET ASIC is showed in figure 2.6. It represents the main blocks that compose the chip [22].

The ASIC provides the time and the energy digitization of the signals that are read by 64 analogic channels [23]. Each channel is constituted by two selectable front-end preamplifiers, to allow the correct readout either with positive or negative polarity signals, that provide the signal to two post-amplifiers, that are fundamental for timing and pulse triggering, and a charge integrator.

A scheme of the analogic part is showed in figure 2.7, the two preamplifiers are feeding two discriminators, schematized as a single block in figure 2.6, and a charge integrator.

One discriminator is used for timing measurement and has a programmable threshold. The second discriminator is used to trigger the incoming signal, it has a high threshold value and its purpose is to identify the “good signal”. This means that the discriminator rejects all those signals whose amplitude results too low, allowing to avoid dark current signals, but only those events

2.2 Application Specific Integrated Circuit (ASIC)

The signals from the discriminators and the charge integrator are read by a digital logic block, that works between 160MHz-200MHz and controls two time to digital converters (TDCs). The first TDC measures the phase of the rising edge of the timing discriminator output compared with the reference clock. The second TDC measures the falling edge of the trigger discriminator output.

The mixed mode TDC block is composed by four time to amplitude converters (TAC), for signal de-randomization a 10bit Wilkinson analog to digital converter (ADC).

The digital logic block also controls the charge integration time interval. Actually, this path has four charge integrators (CI) that allow a high event rate with negligible dead time [22].

The output of the energy branch is also digitized by another Wilkinson ADC.

The digital data corresponding to each event are stored in a register to be later sent to the global output.

There are also two other main blocks, the bias block that provides to generate necessary bias current and threshold voltages, and a global calibration block to set and tune the application.