Readout system for solid-state detector arrays

Details Readout system for solid-state detector arrays Readout system for solid-state detector arrays

2001-07-30

2003-12-02

Hannaher, Constantine (Department: 2878)

Radiant energy

Invisible radiant energy responsive electric signalling

Semiconductor system

C250S370080, C250S2140RC

Reexamination Certificate

active

06657200

ABSTRACT:

FIELD OF THE INVENTION
The present invention relates to the field of electronic circuitry for the coincidence readout of solid-state detector signals, especially for application in the field of nuclear medical imaging.
BACKGROUND OF THE INVENTION
In the field of nuclear medical imaging, solid state detectors are frequently used in order to directly detect the incidence of high energy photons, either those transmitted through the body from a high energy source, such as an X-ray source, or those emitted from radioactive isotopes injected or ingested into the body of a patient. In gamma-ray imaging applications, the detectors must typically be able to determine the energy of the gamma-ray photons emitted from the patient, and the position of incidence on the detector array.
In positron emission tomography (PET), also known as electronic collimation, isotopes that emit positrons are injected or ingested into the body of the examined patient. Each of the emitted positrons annihilates with an electron to produce a pair of 511 Kev photons propagating along the same line but in opposite directions and out of the patient's body. The 511 Kev photons are detected by a camera which has two separate detector heads, which determine the position where the photons interact with the detector heads and the energy of these interacting photons using coincidence detection methods. Photons of the same pair are emitted simultaneously. Accordingly, the detection time of one photon of a pair should differ from the detection time of the second photon, only by a small time interval, &Dgr;t which depends, among other factors, on the time resolution of the system, and on the different time of flight of each photon to its corresponding detector head. The rate of the measured events in the detector heads determines the average time &Dgr;T between two followings events. Two photons are considered as being related to the same pair when they are detected by the two different detector heads of the camera within a time difference &Dgr;t, which satisfies the condition &Dgr;t<&Dgr;T. The coincidence method is based on the detection of the time of the first impacting photon, and the use of that temporal information in deciding whether the following impacting photon is related to that event or not. If the criteria for coincidence are met, a coincidence trigger pulse is generated for informing the signal detection channels to process the signals accordingly.
Another form of gamma ray imaging is known as Single Photon Emission Computerized Tomography, or SPECT. In this method, lower energy photons are detected, such as the 140 keV photons emitted by the decay of Technecium-99 or Thallium-201 previously injected or ingested into the body. In this method, the photons emitted from the body of the patient are typically passed through a lead collimator, in order to ensure that only photons propagating in a straight line are used to produce the image, so that the image is a true representation of the source.
The detectors used in such imaging cameras are constructed of arrays of separate detection modules, each of which itself may have an array of several hundred separate detection areas, typically in the form of pixellated anodes. In a commonly used configuration, each module has 256 individual anodes, and each anode is connected to its own charge sensitive amplifier and signal processor, such that each is an effectively independent detector. Multiple detector readout channels from one module are often integrated into an ASIC (Application Specific Integrated Circuit). Each pixellated detector circuit is able to determine three pieces of data associated with the photons it detects:
(a) the point in time of the photon impact, known as the trigger time, this being important for the coincidence type of measurement;
(b) the energy of the impacting photon, determined from the amount of charge collected; and
(c) the position of impact, determined by means of an address which the detection channel of each anode pixel transmits together with the detection data it has collected.
A threshold level is used to discriminate between random noise in the detector, and a real incident photon. Since the detection channel associated with each detector pixel is only able to handle one detection event at a time, the detection circuitry must be programmed to reject any signals arriving substantially simultaneously in one ASIC from multiple photon impacts in the module. Substantially simultaneously is defined as being within the time taken for the detection circuitry to process and measure the arriving photon. In such a situation, the arrival of a second photon within the coincidence time, (generally of the order of up to tens of nanoseconds) of a first photon causes both signals to be rejected, in order to prevent corruption of the data of either of the photons. This process is known in the art as “pile-up rejection”.
Because of the advisability of using low doses of radiation in the patient, the flux of imaging photons detected is very low. Furthermore, in SPECT imaging, the collimator typically transmits only 10
−4
of the incident flux, such that the detected flux is even further reduced. The importance of utilizing every piece of imaging information in the photon flux is thus of great importance, and every effort should be made not to lose any such information because of detection limitaitons.
As a result of the low flux levels used, simultaneous impact from the patient in a single module or ASIC, whether from direct emission or from Compton scattering within the patient's body, is infrequent. In a single detector channel, this eventuality is even rarer. As a result, the process of pile-up rejection does not generally result in the loss of any significant data from source scattering or simultaneous emission.
However, there are two other processes which can result in the loss of imaging information from the incident flux of photons from the patient's body. Firstly, Compton scattering can occur, not only in the patient's body, but also in the detector bulk itself. This is a much more common phenomenon, and much more serious, as it can affect an appreciable percentage of all photon detection events. When this occurs, the incoming photon makes an initial impact within the detector, and gives up part of its energy in producing a charge of excited electron-hole pairs. The electrons in this cloud of charge then drift towards the anode under the influence of the field present in the detector, and appear on the anode opposite the point of first impact. The secondary scattered photon continues its path within the detector, at a lower energy, until it makes a second impact within the detector bulk, again resulting in another cloud of carriers, the electrons of which are collected at the closest anode. This anode may be in a pixel a considerable distance from the pixel of the initial impact, especially for high energy photons. (The hole motion has been neglected for the purposes of this explanation). As a result of this process, the primary and secondary impacts are detected by different pixels at different times in the detector module. Furthermore, they each have a different energy, as the energy of the incident photon is shared between the two impacts, and neither therefore has the expected energy of the emitted photon.
The second process which can result in the loss of imaging information, occurs when a photon impacts the detector very close to the border between two pixels. This is known as sharing. When this happens, the cloud of charge carriers is approximately divided between two neighboring pixels, and some of the electrons are collected by one anode, and others by the neighboring anode. The incident photon is thus measured as if it were two separate impacts, in neighboring pixels, neither of which has the characteristic photon energy sought for constructing the image, and this event too would be rejected from the imaging process.
Sharing can be even more complex if the impact occurs near the junction of four pixels

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