Readout circuit for a charge detector

Details Readout circuit for a charge detector Readout circuit for a charge detector

2001-04-05

2003-07-08

Hannaher, Constantine (Department: 2878)

Radiant energy

Invisible radiant energy responsive electric signalling

Semiconductor system

C250S363030

Reexamination Certificate

active

06590215

ABSTRACT:

FIELD OF THE INVENTION
This invention relates to a charge detector for reading charge produced by an active pixel in the detector.
BACKGROUND OF THE INVENTION
A known diagnostic technique used in tomography for locating tumors involves injecting into a patient's bloodstream a radioactive isotope which targets the tumor, so that the location of the tumor can be derived by detecting the location of the radioactive isotope. Typically, the radioactive isotope emits &ggr;-rays which are dispersed from the tumor site. In order to achieve the desired detection so as to determine the precise location of the tumor, it is necessary to image the patient's body in such a manner as to detect only those &ggr;-rays which are emitted normally from the body and to ignore those &ggr;-rays which are dispersed in other directions.
U.S. Pat. No. 5,656,818 (Nygard) assigned to the one of present applicants discloses such a radiation imaging system that includes a detector unit and a receiver unit. The detector unit includes a two-dimensional sensor, first and second amplifier channels, first and second multiplicity generators and first and second address generators. The two-dimensional sensor includes first and second sensing segments that sense radiation in a first direction and a second direction, respectively. Each first amplifier channel generates an output signal based on a detection output from a corresponding one of the first sensing segments. The first multiplicity generator generates a first multiplicity signal representing a number of the first amplifier channels generating output signals. The first address generator generates a first analog address of a first amplifier channel associated with a received output signal. The second amplifier channels, second multiplicity generator and second address generator operate in like manner with respect to the second sensing segments. The receiver unit includes converters for converting the first and second analog addresses into first and second digital addresses. The receiver unit also includes a tester for testing whether the first and second digital addresses represent a valid position address in the first and second direction based on the first and second multiplicity signals.
Different types of computer tomography are known in which such a radiation imaging system may be embodied. In Single Photon Emission Computed Tomography (SPECT) more than one detector is rotated around the subject. During the rotation of the detector, the counting of the gamma rays is repeated. Then, the radioisotope's distribution (tomographic image) is reconstructed based on the obtained count values of the &ggr;-rays.
In contrast to SPECT where a radioisotope in the body emits &ggr;-rays produced by a single photon, in Positron Emission Tomography (PET) a patient is administered a radioisotope that emits positrons (i.e. positively charged electrons). When the positrons meet electrons within the body, the positrons and electrons mutually annihilate and produce two &ggr;-rays that propagate away from each other at an angle of 180° and are detected by respective detector segments in the PET scanner. The scanner's readout electronics record the detected &ggr;-rays and map an image of the area where the radioisotope is located. Here also two simultaneous detections are indicative of a positron emission from the tumor site.
Thus, in PET two simultaneous &ggr;-rays must be detected on opposite sides of the patient's body. The positrons are extremely short lived and simultaneity implies that the two &ggr;-rays can both be detected within a short time window, which is 10 ns, for example. The PET scanner surrounds the patient like a CT scanner and typically comprises in the order of 200,000 pixels on two detector segments. Thus, it is necessary to detect two excited pixels within a short time difference and then to read out the energy of these two pixels. Only if the energy of each active pixel equals about 511 keV, (i.e. the energy of the incident &ggr;-radiation) are the two photons the result of positron-electron annihilation and thus indicative of the tumor's location. Also in a PET scanner Compton scattering can occur within the detector, whereby a photon is only partially absorbed by the pixel and partially scattered to another pixel or occasionally to even more than a single pixel. In this case the sum of the energies of simultaneous active pixels will be equal to about 511 keV.
Another apparatus used for nuclear imaging is the Compton Camera. In the Compton Camera in order to determine the location of the &ggr;-ray based on the detection of the single photon thus emitted, it is necessary that Compton scattering occur, so that another photon will be emitted substantially simultaneously, thereby allowing the angle of the incident &ggr;-ray to be calculated.
Here, too, it is necessary to establish simultaneity of a &ggr;-ray striking multiple pixels, although Compton Camera and PET must detect simultaneous &ggr;-rays and so the time difference for establishing simultaneity of two or more pixels becoming active is more critical than SPECT. Known architectures for PET scanners use sequential readout whereby each pixel is read sequentially in order to determine whether or not it is active.
In EP 893 705 published on Jan. 27, 1999 entitled
“Multi-Channel Reading Circuit for Particle Detector
” and assigned to one of the present applicants, there is described a method for reading an array of pixels in a 2-dimensional image sensor so as to reduce the time taken to detect a single “active” pixel. Specifically, it is proposed in EP 893 705 to group pixels into predefined groups or segments (typically associated with a single ASIC) and to read out the pixels of only those groups of pixels that are found to contain an active pixel. Such an approach is based on the principle that determination of an active group may be made quickly providing that, at this initial stage, it is not required to determine which specific pixel within the active group is itself active. Having thus established which groups of pixels are active, each pixel in the active group is then read sequentially.
Whilst this is suitable for SPECT, it is not suitable for PET principally due to the very high rate at which pixels are activated in a PET scanner owing to its absence of a collimator. As a result, there are too many active pixels to make it practical to provide a tradeoff between providing a large number of groups or segments each containing a sufficiently small number of pixels that sequential readout in each group can be achieved in reasonable time. Specifically, whilst it is theoretically possible to reduce the number of pixels in each group so that pixels in each group that become active simultaneously within the 10 ns time window can be determined, the number of A/D converters that would then be required to effect digital processing of the active pixels would, in practice, be prohibitive.
The reading circuit described in EP 893 705 uses two detectors so as to detect Compton scattering, requiring that data be read out in the first detector from a large number of addressable pixels along respective channels in order to detect which pixel is “active”. This is done by first integrating the charge associated with each pixel using an integrator in the form of an operational transconductance amplifier having a feedback capacitor. The integrated charge pulse is then amplified and shaped and the resulting analog signal is sampled and held, allowing its magnitude to be measured. In order to measure the peak magnitude of the shaped signal, the shaped signal must be very accurately sampled at the peak value. This requires an accurate determination of the peak time that occurs a fixed time difference t
p
after the emission of charge by the excited pixel. The fixed time difference t
p
is a function of the RC time constant of the shaper circuit and is therefore known.
Thus, in order to know when to sample the integrated charge signal, the time of occurrence t
o
of each charge emission m

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