PET Project

Goal

The MEDUSA project focuses on R&D for high energy physics instrumentation with two important and dependant goals. One is to contribute to the research for future particle detectors and develop new improved detectors for the LHC upgrade as well as the planned international linear collider. The other is to provide new technologies for medical imaging devices such as PET. With this, we hope to contribute to bridging the gap between the particle physics research and the medical technology to fully take advantage of the latest development.

Two complementary detector technologies are highly interesting for medical applications. First, the compact calorimeter is a new technology for detection of photons and hadrons, based on a new type of silicon photomultipliers. These detectors form the base of modern medical imaging technology where precise localisation of radioactive tracers in the body. Aquisition speed and sensitivity are two central challenges for high energy physics. In addition, these detectors can be used to develop Time-of-Flight measurements.

The 3D semiconductor devices are based on another new technology, aiming to provide particle and radiation detection by the use of 3 dimensional silicon pixels. The advantage of this method is that these sensors have improved radiation hardness as well as a better to-the-edge detection. A substancial challenge is to provide thin devices and 3D integration, one of the requirement for linear accelerators. Semiconductor detectors are widely used in imaging spectroscopy and particle tracking of ionising radiation, both for charged particles and photons.

This project is set up with the collaboration of the new PET center at Haukeland University Hospital and we will closely collaborate with their researchers. Other research partners are the University of Oslo as well as the CLIC, ALICE and the ATLAS collaboration at CERN and the ILC project.

General PET technology

Positron Emission Tomography (PET) is recognized as a great medical imaging devices thanks to its non invasive technology. 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 conditions and cancer, its big advantage is to identify the onset of a disease process before anatomical changes (that can be seen with other imaging processes such as computed tomography (CT) or MRI) related to the disease.

Radiotracers

Fluorodeoxyglucose-FDG

Positron - Electron Annihilation

The PET technology is based on radioactive emission. Radioactive substances are combined to molecules that the studied cells use particularly in their metabolism. These tracers are radioactive substances. The first step in PET imaging is the production of radionuclides by a cyclotron. These radionuclides will be attached to molecules used by the body before being injected to the patient by intravenous way. The molecule and the adionuclide form the radiotracer. The tracer is injected to the patient and, following the half life of the radionuclide, it will become stable by emitting a positron and a neutrino (the proton which stays in excess will become a neutron). Then, the emitted positron travels a short distance before encountering an electron. When they meet each other, the two particles combine and annihilate each other resulting in the emission of two 511 keV gamma rays in opposite directions.

Scintillators

Electronic transition in a crystal

A scintillator is a substance that absorbs high energy and then, in response, emits photons. Scintillators are defined by their light output (number of emitted photons per unit absorbed energy), short fluorescence decay times, and optical transparency at wavelengths of their own specific emission energy. The high Z-value of the constituents and high density of inorganic crystals favour their choice for gamma-rays spectroscopy (rather than organic crystal) because heavy nucleuses accept better gammas than light nucleus. The scintillation mechanism in inorganic materials depends on energy states determined by the crystal lattice of the material. Absorption of energy can result in the elevation of an electron from its normal position in the valence band across the gap in the conduction ban, leaving a hole in the valence band. A charged particle passing through the detection medium will form a large number of electron-hole pairs, created by the elevation of electrons from the valence to the conduction band. The positive hole will quickly drift to the location of an impurity and ionize it, because the ionization energy of this impurity will be less than that of a typical lattice site. Meanwhile, the electron is free to migrate through the crystal and will do so until it encounters an ionized activator. At this point, the electron can drop into the impurity site, creating a neutral impurity configuration which can have its own set of excited energy states. If the activator state that is formed is an excited configuration with an allowed transition to the ground state, its deexcitation will occur very quickly and with high probability for the emission of the corresponding photon. The migration time for the electronics is shorter than the drop-out time: therefore, the decay time of these states determines the time characteristics of the emitted scintillation light. In order to fully utilize the scintillation light, the spectrum should fall near the wavelength region of maximum sensitivity for the device used to detect the light.

	NaI(Ti)	BGO	LSO	LYSO
ZE	50	74,2	65	65
Density	3,67	7,13	7,35	7,1
Attenuation coeff (cm-1)	0,34	0,95	0,8	0,83
Decay time (ns)	230	300	40	42

We see, through this chart, that the discovery of the LSO and LYSO crystals have helped to make some progresses. LSO and LYSO crystal are the best compromise for a high attenuation coefficient and a short decay time, two useful properties to improve time resolution in PET scanner.

Coincidence detection

Coincidence detection principle

In a PET camera, each detector generates a time pulse when it registers an incident photon. These pulses are then combined in coincidence circuitry, and if the pulses fall within a short time-window, they are deemed to be coincident. A diagram of this process is shown below: A coincidence event is assigned to a line of response (LOR) joining the two relevant detectors. Coincidence events in PET fall into 3 categories: true, scattered and random.

3 different coincidence detection events

• True coincidences occur when both photons from an annihilation event are detected by detectors in coincidence, neither photon undergoes any form of interaction prior to detection, and no other event is detected within the coincidence time-window.

• A scattered coincidence is one in which one of the detected photons (sometimes both) has undergone at least one Compton scattering event prior to detection. Since the direction of the photon is changed during the Compton scattering process, it is highly likely that the resulting coincidence event will be assigned to the wrong LOR. Scattered coincidences add a background to the true coincidence distribution which changes slowly with position, decreasing contrast and causing the isotope concentrations to be overestimated. They also add statistical noise to the signal. The number of scattered events detected depends on the volume and attenuation characteristics of the object being imaged.

• Random coincidences occur when two photons, not arising from the same annihilation event, are incident on the detectors within the coincidence time window of the system. As with scattered events, the number of random coincidences detected also depends on the volume and attenuation characteristics of the object being imaged, and on the geometry of the camera. The distribution of random coincidences is fairly uniform across the field of view and will cause isotope concentrations to be overestimated if not corrected for. Random coincidences also add statistical noise to the data.

Time of Flight

Time-of-flight PET takes advantage of the difference in arrival times of two photons from the same annihilation event to infer spatial information of this event. A detected coincidence between two crystals will have a time difference for any annihilation event that does not occur at the midpoint between the detectors, this time difference is used to place the position of the event.

Time of flight principle

Time of flight principle

Technology

Avalanche Photodiodes

Structure of MAPD

An APD is basically a p-n junction diode operated at large reverse bias voltage. The physical mechanism which avalanche gain depends, is the impact ionization. It occurs when the electric field in the depletion region is strong enough: an electron colliding with a bound valence electron transfers enough energy to this electron to ionize it. The additional carriers can gain sufficient energy from the electric field to cause further impact ionization, creating an avalanche of carriers.

• proportional mode

In a proportional counter, each original electron leads to an avalanche which is basically independent of all other avalanches formed from the other electrons associated with the original ionizing event. The collected charge remains proportional to the number of original electrons.

proportional mode

• Geiger mode

With a higher electric field, a situation is created, in which one avalanche trigger itself a second avalanche at a different position.

Geiger mode

passive quenching

The difference between both modes relies on the holes: in Geiger mode, they trigger avalanches, whereas in proportional mode they have not enough energy to do so. From the critical value of the electric field (corresponding to the breakdown voltage), a self propagating chain reaction occurs. In principle, an exponentially growing number of avalanches could be created.

• b- Passive quenching

The avalanche photodiode (i. e. pixel for the silicon photomultiplier) is connected to the power supply through a large series resistor Rs. If the current through the diode tends to zero, the voltage across the diode equals Vbias, which will be larger than the breakdown voltage. So, when the diode breaks down, the series resistor reduces the voltage across the APD, what quenches the avalanche. After the breakdown is quenched, the diode is recharged through the resistor. The APD is now ready to receive another photon.

different MAPDs

The MAPDs are 3 mm *3 mm, composed of 104 pixels /mm2 (9.104 pixels in total). They should be operated in inverse direction: anode should be grounded, while positive voltage in range 132-136V (it depends on the MAPDs and it is reported by the manufacturer). Exceeding the voltage 137V leads to unstable operation and even to the destruction of MAPD.

The resistance of each pixel allows the passive quenching. Pixels are electrically decoupled from each other by polysilicon resistors and are connected by common Al strips, in order to readout the MAPD signal. Each MAPD pixel operates as a binary device but MAPD in whole is an analogue detector. The output signal allows us to determine the number of fired pixels: in fact, the output signal is proportional to the number of fired pixels. The MAPD is intrinsically very fast due to the very small width of the depletion region and the extremely short Geiger type discharge. We must keep in mind that the name of the device depends on the manufacturer. MPPC (for Multi-Pixel Photon Counter) and SiPM (Silicon PhotoMultiplier) are two other usual names.

Properties of the devices

• Time resolution: even if photons simultaneously enter different pixels at the same time, the output pulse from each pixel will not necessarily be the same time so that a fluctuation or time jitter occurs. When two photons enter APD pixels at a certain time difference which is shorter than this jitter, then that time difference is impossible to detect. Time resolution is the minimum time difference that can be detected by APD pixels and is defined as FWHM of the distribution of the time jitter.

• Photon Detection Efficiency (PDE): this is a measure of what percent of the incident photons were detected.

• Dark count: output pulses are produced not only by photon-generated carriers but also by thermally-generated dark current carriers. The dark current pulses are measured as dark count which then causes detection errors. Although increasing the reverse voltage improves photon detection efficiency, it also increases the dark count. The dark count can be reduced by lowering the temperature.

• Absolute gain: the absolute gain is the number of charges which have been created at the output of the MAPD when one photon has hit this device.

• Quantum efficiency (QE): quantum efficiency is a value showing the number of electrons or holes created as photocurrent divided by the number of incident photons, and is usually expressed as a percent.

• Afterpulse: afterpulses are spurious pulses following the true signal, which occur when the generated carriers are trapped by crystals defects and then release at a certain delay time. A fterpulses cause detection errors. The lower the temperature, the higher the probability that carriers may be trapped by crystal defects, so afterpulses will increase.

• Crosstalk: in an avalanche multiplication process, photons might be generated which are different from photons initially incident on an APD pixel. If those generated photons are detected by other APD pixels, then the MAPD output shows a value higher than the number of photons that were actually input and detected by the MAPD. This phenomenon is thought to be one of the causes of crosstalk in the MAPD.

Documentation

Dominik 09:06, 12 June 2009 (UTC)