A higher accuracy beam hardening correction with a low calculation load is performed with objects whose elements have a wider range of effective atomic numbers Z.sub.eff, thereby contributing to presentation of more quantitative X-ray images. Of two or more X-ray energy bins, two X-ray bins are selected to normalize X-ray attenuation amount μt in those bins such that one or more normalized X-ray attenuation amounts are obtained at each pixel areas. From reference information indicating a theoretical relationship of correspondence between the normalized X-ray attenuation amounts and effective atomic numbers of elements, one ore more effective atomic numbers are estimated every pixel area. Among the one or more effective atomic numbers (Z.sub.High, Z.sub.Low) and an effective atomic number (Zm) preset for the beam hardening correction, two or more atomic numbers are subjected to their equality determination.

A radiation position detection method includes: a first step of calculating a first centroid position in an incident direction regarding positions where scintillation light is detected, on the basis of electrical signals; and a second step of specifying, on the basis of a first table showing first identification regions for identifying the plurality of segments, and the first centroid position, the segment that initially generates the scintillation light. The first identification region includes a first region, a second region, and a third region. In the second step, in a case where the first centroid position is located in the first region or the third region, the first segment is specified as the segment that initially generates the scintillation light, and in a case where the first centroid position is located in the second region, the second segment is specified as the segment that initially generates the scintillation light.

A method for determining an abnormality score map for Single-photon Emission Computed Tomography (SPECT) gamma camera flood analysis includes extracting a plurality of image patches from an input flood image and generating a feature vector for each image patch. A per-patch abnormality score is generated for each feature vector by comparing the feature vector against a normal flood feature dictionary comprising one or more normal flood feature vectors generated using a plurality of normal flood images. Then, an abnormality score map may be generated to depict the per-patch abnormality scores for the input flood image.

Gamma-ray detectors for the detection of one or more gamma-rays are provided. Also provided are methods of using the detectors for the detection of one or more gamma-rays. The detectors can be used in high-spatial resolution PET systems, including time-of-flight (TOF)-PET systems. Some of the gamma-ray detectors utilize fluors and an optical imaging system to determine the time and location of a first scattering event of a gamma-ray in a low atomic number scintillating medium. Some of the gamma-ray detectors determine the time and location of a first scattering event of a gamma-ray in a low-density scintillating medium by imaging scintillation photons from the scattering event as a time-series of photon “rings” using a planar pixilated photodetector as a scintillation photon counter.

During calibration of a SPECT system, system-specific sensitivities and cross-calibration factors for multiple isotopes for correcting for dose are determined for various combinations of options, including the option of which specific well counter with which to measure the dose. The options may include selected energy windows for isotopes with multiple energy windows. This arrangement allows for custom-specified isotopes not included in standard listings. For use with a particular patient, the cross-calibration factor for the well counter used to measure the dosage for the patient is accessed and used for dose correction. More accurate quantitative functional information may result from the corrected dose. The cross-calibration may be more easily implemented despite the options using the sensitivities and cross-calibrations provided for various combinations.

The present invention relates to a calibration method for a gamma ray detector (100) including a pixelated scintillator array (110) for emitting scintillation photons at photo conversion positions (94) in response to incident gamma rays (90), and a pixelated photodetector array (120) for determining a spatial intensity distribution of the scintillation photons. The present invention bases on the idea that using the concept of optical light sharing of scintillation photons, which are emitted in one element, i.e., one scintillator pixel (112) of the scintillator array (110) and distributed over multiple photodetector pixels (122) of the pixelated photodetector army (120), allows obtaining an estimate for the time skew between adjacent photodetector pixels (122). The present invention further relates to a calibration module (200) for a gamma ray detector (100) including a recorder (210) and a processing module (220) for performing the function of the above-explained method. Still further, the present invention relates to a gamma my detector (100) as well as to a medical imaging device (50) comprising this gamma my detector (100).

Systems and methods to determine random coincidence rates include determination of a detector rate for each of a plurality of detectors of a positron emission tomography scanner based on a frame of positron emission tomography data, determination of a sensitivity for each detector crystal of the plurality of detectors, based on the detector rate of the detector including the detector crystal, determination of a singles rate for each detector crystal based on the detector rate of the detector including the detector crystal and the determined sensitivity of the detector crystal, estimation of a mean random coincidence rate for each of a plurality of pairs of the detector crystals based on the singles rate of each detector crystal of each of the plurality of pairs of the detector crystals, correction of the acquired frame of positron emission tomography data based on the estimated mean random coincidence rates, and reconstruction of a positron emission tomography image based on the corrected frame of positron emission tomography data.

Systems and methods to determine random coincidence rates include determination of a detector rate for each of a plurality of detectors of a positron emission tomography scanner based on a frame of positron emission tomography data, determination of a sensitivity for each detector crystal of the plurality of detectors, based on the detector rate of the detector including the detector crystal, determination of a singles rate for each detector crystal based on the detector rate of the detector including the detector crystal and the determined sensitivity of the detector crystal, estimation of a mean random coincidence rate for each of a plurality of pairs of the detector crystals based on the singles rate of each detector crystal of each of the plurality of pairs of the detector crystals, correction of the acquired frame of positron emission tomography data based on the estimated mean random coincidence rates, and reconstruction of a positron emission tomography image based on the corrected frame of positron emission tomography data.

The invention relates to a combined imaging detector for detection of gamma and x-ray quanta comprising an x-ray detector (31) for generating x-ray detection signals in response to detected x-ray quanta and a gamma detector (32) for generating gamma detection signals in response to detected gamma quanta. The x-ray detector (31) and the gamma detector (32) are arranged in a stacked configuration along a radiation-receiving direction (33). The gamma detector (32) comprises a gamma collimator plate (320) comprising a plurality of pinholes (321), and a gamma conversion layer (322, 324) for converting detected gamma quanta into gamma detection signals.

Described herein is multiplexing scintillation blocks, called interblock muxing. Specifically, the start of an annihilation event is recorded and assigned a time stamp while the energy of the entire event is recorded separately. All events occurring at a series of multiplexed scintillation blocks are reported to a processor which distinguishes individual events and assigns the start of each event with its corresponding energy, thereby allowing for cheaper and more efficient processing of events during PET imaging.