Radiant energy – Invisible radiant energy responsive electric signalling – Semiconductor system
Reexamination Certificate
1998-08-17
2001-05-22
Hannaher, Constantine (Department: 2878)
Radiant energy
Invisible radiant energy responsive electric signalling
Semiconductor system
C250S363030
Reexamination Certificate
active
06236050
ABSTRACT:
FIELD OF THE INVENTION
The present invention relates generally to detection systems, and more particularly, to a method and apparatus for imaging gamma rays, x-rays, and positrons.
BACKGROUND OF THE INVENTION
The dominant absorption process for gamma rays of about 0.05 to 30 MeV is the Compton interaction. At present the best method to detect these gamma rays is the Compton double scatter technique since single Compton scattering alone gives neither the direction nor the energy of the incident gamma ray. In contrast, the Compton double scatter technique yields both the direction and the energy of the incoming photon.
Present detectors which use the Compton double scatter technique determine the direction of the incoming photon to a ring in the sky since the direction of the recoil electron at the first scatter cannot be measured. Time-of-flight measurements are normally used to discriminate between gamma rays coming through the field-of-view and those entering through the back of the system. Typically this requires that the first scatterer (ie., the hodoscope) is separated from the second scatterer (i.e., calorimeter) by approximately 1.5 meters.
Emission computed tomography (ECT) and associated technologies are mainly used for the detection and imaging of the radiation produced by radiotracers and radiopharmaceuticals. The primary application for ECT systems is in medical study and diagnosis due to the potential for imaging organ functions in real time. The two major ECT instruments presently used are Single Photon Emission Computed Tomography (SPECT) and Positron Emission Tomography (PET).
Positron scanners for use in locating brain tumors were first developed in the early 1960s with the first PET system completed in 1975. Pet systems have become an essential medical diagnostic tool for a variety of reasons. For one, very high efficiencies utilizing positron emitting labels can be achieved through the coincidence collimation of the annihilation radiation. Another advantage is that the common radiopharmaceuticals used with PET systems typically have very short life times, thus allowing large doses to be administered to a patient as well as the performance of repetitive studies. Recently, the utilization of photon time-of-flight information with fast scintillators has improved the SNR that can be obtained in images of the distribution of positron emitting radionuclides.
Present PET systems use bismuth germanate oxide (BGO) crystals. BGO crystals have the highest effective atomic number and stopping power of any scintillator crystal available today. This translates into a higher photopeak fraction and a lower Compton continuum than other crystals such as NaI and BaF
2
. Gadolinium orthosilicate (GSO) crystal is an alternative which has a slower decay time but larger pulse yield than BGO. PbCO
3
crystals nearly equal the stopping power of BGO but the light output is about 10 times lower.
From the foregoing, it is apparent that an improved imaging system for use with gamma rays, x-rays, and positrons is desired.
SUMMARY OF THE INVENTION
The present invention provides a high sensitivity, low background detector for use in a variety of applications. The detector also has excellent angular resolution for accurate direction measurement and imaging; good energy resolution for the identification of the source material by its energy spectrum; and low power consumption. Both the direction and energy of the incident photons is measured using a Compton double scatter technique with recoil electron tracking.
The Compton double scatter technique involves two detector layers; a silicon microstrip hodoscope and a calorimeter. The incoming photon Compton scatters in the hodoscope. The second scatter layer is the calorimeter where the scattered gamma ray is totally absorbed. The recoil electron in the hodoscope is tracked through several detector planes until it stops. The x and y position signals from the first two planes of the electron track determine the direction of the recoil electron. The energy loss from all planes is summed to determine the energy of the recoil electron.
The hodoscope of the disclosed system is constructed of position sensitive, double-sided silicon microstrip detectors, preferably with a strip pitch of between 0.5 and 1 millimeter and a thickness of between 200 and 1000 micrometers. The pixel size of the microstrip detectors ranges from approximately 1 square millimeter to approximately 1 square centimeter.
The measurement of the direction of the Compton recoil electron track reduces the incident gamma ray event ring to an event arc. The recoil electron direction calculation requires only the x and y coordinates of the first two adjacent planes along the track of the recoil electron. For the measurement of the direction of motion of the recoil electron (i.e., moving forward or backward in the hodoscope) a track which penetrates 2 or more adjacent planes is required. If the energy of the recoil electron is low enough, it may be absorbed in the same detector plane and not produce a track. Thus for a 200 micrometer thick detector, recoil electron tracking is effective for electrons with an energy of greater than 0.25 MeV energy. Therefore gamma rays with energies greater than 1 MeV will produce recoil electron tracks with high probability. For low energy gamma rays which do not produce recoil electron tracks, imaging is carried out using event rings instead of arcs. This technique increases the background, resulting in a decrease in sensitivity for low energy gamma rays. By utilizing thinner detectors the recoil electron tracking threshold can be reduced to even lower energies.
The selection of the detector for use as the second scatterer depends primarily upon the desired detection energy as well as the desired detection energy range. In one embodiment of the invention, the second scatterer uses thallium activated cesium iodide (i.e., CsI(Tl)) detectors viewed by photodiodes. In a second embodiment used to obtain higher energy resolution, a germanium array calorimeter is used. The germanium array requires refrigeration to liquid nitrogen temperatures. Alternatively, a detection system using only a hodoscope with a large number of silicon detector planes can be used for high energy resolution.
The double Compton scatter measurement determines the direction of the incident photon to a cone with a half angle equal to the scatter angle. This type of measurement requires special data analysis software. The data analysis can be carried out by cone interaction, Maximum Likelihood or Maximum Entropy techniques or using a direct data analysis and imaging technique such as Direct Linear Algebraic Deconvolution (DLAD).
In another embodiment of the invention, a Compton scatter PET system is provided. This system can be designed as a cylindrical detector with a length of approximately 30 centimeters or greater. Cylindrical geometry leads to the production of accurate 3-dimensional images. Alternately, a ring detector with a width of about 7 centimeters or less can be used to provide a multi-slice ring type system. Since the PET system according to the present invention does not utilize PMTs, it is substantially less bulky than present systems. As a result, the structural requirements placed on the gantry are less stringent.
Photon attenuation inside the patient can be corrected using the techniques already developed. For example, the boundary method has already been successfully applied to attenuation correction in PET image reconstruction. In this method, the organ boundaries are determined by transmission tomography and each region is enclosed by a boundary and assigned an average attenuation coefficient. Attenuation correction factors for all angular views can be calculated from the quantized image.
The PET embodiment of the present invention preferably uses thin film strip detectors with high stopping power such as position sensitive, double-sided CdZnTe strip detectors with an approximate strip size of at least 0.1 millimeters in both the x and y dimensions. The CdZnTe detecto
Gabor Otilia
Hannaher Constantine
Snider Ronald R.
Snider & Associates
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