Abstract: A time of flight positron emission tomography apparatus (100) includes a detector (106), a data acquisition system (120), a coincidence system (122) and a reconstructor (129). Various elements of an imaging chain influence the temporal resolution of the system (100) so that positron data collected along different lines of response is characterized by different temporal resolutions. The different temporal resolutions are used to estimate the positions of detected events along their respective lines of response.

Abstract: A nuclear imaging system includes a crystal identification system which receives a flood image which includes a plurality of peaks, each peak responsive to radiation detected by a corresponding scintillator crystal. A crystal identification processor partitions the flood image into a plurality of candidate regions with a watershed segmentator implementing a watershed algorithm. The candidate regions are linked in an adjacency graph, and then classified as background or relevant, where relevant regions contain a peak within the watershed lines. The regions are then assigned to a crystal according to an objective function and an assignability score. A calibration processor maps the peaks to a rectangular grid.

Abstract: A nuclear imaging system (10) includes a crystal identification system (40) which receives a flood image (30) which includes a plurality of peaks, each peak responsive to radiation detected by a corresponding scintillator crystal. A crystal identification processor (42) partitions the flood image (30) into a plurality of regions (56), each region being masked to correspond to one of an array of nuclear detectors. A model image (62) is generated in which the at least one Gaussian models represents the identified peaks. Misidentified peaks in the model image (62) in which locations of the peaks in the flood image (30) differ from the corresponding scintillator crystal are determined and the locations of the misidentified peaks in the flood image (30) are corrected. A calibration processor (43) corrects geometric distortions in acquired projection data according to the corrected peaks.

Abstract: A nuclear imaging system includes a crystal identification system which receives a flood image which includes a plurality of peaks, each peak responsive to radiation detected by a corresponding scintillator crystal. A crystal identification processor partitions the flood image into a plurality of candidate regions with a watershed segmentator implementing a watershed algorithm. The candidate regions are linked in an adjacency graph, and then classified as background or relevant, where relevant regions contain a peak within the watershed lines. The regions are then assigned to a crystal according to an objective function and an assignability score. A calibration processor maps the peaks to a rectangular grid.

Abstract: In nuclear imaging, when a gamma ray strikes a scintillator, a burst of visible light is created. That light is detected by a photodetector and processed by downstream electronics. It is desirable to harness as much of the burst of light as possible and get it to the photodetector. In a detector element (18), a first reflective layer (44) partially envelops a scintillation crystal (34). The first reflective layer (44) diffuses the scintillated light. A second reflective layer (46) and a support component reflective layer (48) prevent the light from leaving the scintillation crystal (34) by any route except a light emitting face (36) of the scintillator (34). In another embodiment, a light concentrator (50) is coupled to the scintillator (34) and channels the diffuse light onto a light sensitive portion of a photodetector (38). The reflective layers (44, 46, 48) and the concentrator (50) ensure that all or nearly all of the light emitted by the scintillator (34) is received by the photodetector (38).

Abstract: A nuclear imaging system (10) includes a crystal identification system (40) which receives a flood image (30) which includes a plurality of peaks, each peak responsive to radiation detected by a corresponding scintillator crystal. A crystal identification processor (42) partitions the flood image (30) into a plurality of regions (56), each region being masked to correspond to one of an array of nuclear detectors. A model image (62) is generated in which the at least one Gaussian models represents the identified peaks. Misidentified peaks in the model image (62) in which locations of the peaks in the flood image (30) differ from the corresponding scintillator crystal are determined and the locations of the misidentified peaks in the flood image (30) are corrected. A calibration processor (43) corrects geometric distortions in acquired projection data according to the corrected peaks.

Abstract: In a hybrid PET-MR system, PET detector elements (30) are added in the bore (14), in close proximity to the gradient coils (16). Fluid coolant is supplied to transfer heat from the PET detector elements (30). Thermal insulation (80) insulates the fluid coolant and the PET detector elements (30) from the gradient coils (16). In some embodiments, a first coolant path (90) is in thermal communication with the electronics, a second coolant path (92) is in thermal communication with the light detectors, and a thermal barrier (94, 96) is arranged between the first and second coolant paths such that the first and second coolant paths can be at different temperatures (Te, Td). In some embodiments a sealed heat pipe (110) is in thermal communication with a heat sink such that working fluid in the heat pipe undergoes vaporization/condensation cycling to transfer heat from the detector elements to the heat sink.

Abstract: A positron emission tomography (PET) detector ring comprising: a radiation detector ring comprising scintillators (74) viewed by photomultiplier tubes (72); and a magnetic field shielding enclosure (83, 84) surrounding sides and a back side of the annular radiation detector ring so as to shield the photomultiplier tubes of the radiation detector ring. Secondary magnetic field shielding (76?) may also be provided, comprising a ferromagnetic material having higher magnetic permeability and lower magnetic saturation characteristics as compared with the magnetic field shielding enclosure, the second magnetic field shielding also arranged to shield the photomultiplier tubes of the radiation detector ring. The secondary magnetic field shielding may comprise a mu-metal.

Abstract: In nuclear imaging, when a gamma ray strikes a scintillator, a burst of visible light is created. That light is detected by a photodetector and processed by downstream electronics. It is desirable to harness as much of the burst of light as possible and get it to the photodetector. In a detector element (18), a first reflective layer (44) partially envelops a scintillation crystal (34). The first reflective layer (44) diffuses the scintillated light. A second reflective layer (46) and a support component reflective layer (48) prevent the light from leaving the scintillation crystal (34) by any route except a light emitting face (36) of the scintillator (34). In another embodiment, a light concentrator (50) is coupled to the scintillator (34) and channels the diffuse light onto a light sensitive portion of a photodetector (38). The reflective layers (44, 46, 48) and the concentrator (50) ensure that all or nearly all of the light emitted by the scintillator (34) is received by the photodetector (38).

Abstract: A method for calibrating an imaging system includes coincident detecting scatter radiation events from a calibration source located within a bore of the imaging system. The scatter radiation events are subsequently used to compute calibration time offsets for each detector channel in the imaging system. Each detector channel is then calibrated with respective calibration time adjustments.

Abstract: In a hybrid PET-MR system, PET detector elements (30) are added in the bore (14), in close proximity to the gradient coils (16). Fluid coolant is supplied to transfer heat from the PET detector elements (30). Thermal insulation (80) insulates the fluid coolant and the PET detector elements (30) from the gradient coils (16). In some embodiments, a first coolant path (90) is in thermal communication with the electronics, a second coolant path (92) is in thermal communication with the light detectors, and a thermal barrier (94, 96) is arranged between the first and second coolant paths such that the first and second coolant paths can be at different temperatures (Te, Td). In some embodiments a sealed heat pipe (110) is in thermal communication with a heat sink such that working fluid in the heat pipe undergoes vaporization/condensation cycling to transfer heat from the detector elements to the heat sink.

Abstract: A time of flight positron emission tomography apparatus (100) includes a detector (106), a data acquisition system (120), a coincidence system (122) and a reconstructor (129). Various elements of an imaging chain influence the temporal resolution of the system (100) so that positron data collected along different lines of response is characterized by different temporal resolutions. The different temporal resolutions are used to estimate the positions of detected events along their respective lines of response.

Abstract: A positron emission tomography (PET) detector ring comprising: a radiation detector ring comprising scintillators (74) viewed by photomultiplier tubes (72); and a magnetic field shielding enclosure (83, 84) surrounding sides and a back side of the annular radiation detector ring so as to shield the photomultiplier tubes of the radiation detector ring. Secondary magnetic field shielding (76?) may also be provided, comprising a ferromagnetic material having higher magnetic permeability and lower magnetic saturation characteristics as compared with the magnetic field shielding enclosure, the second magnetic field shielding also arranged to shield the photomultiplier tubes of the radiation detector ring. The secondary magnetic field shielding may comprise a mu-metal.

Abstract: A method for calibrating an imaging system includes coincident detecting scatter radiation events from a calibration source located within a bore of the imaging system. The scatter radiation events are subsequently used to compute calibration time offsets for each detector channel in the imaging system. Each detector channel is then calibrated with respective calibration time adjustments.

Abstract: A method for calibrating an imaging system includes coincident detecting scatter radiation events from a calibration source located within a bore of the imaging system. The scatter radiation events are subsequently used to compute calibration time offsets for each detector channel in the imaging system. Each detector channel is then calibrated with respective calibration time adjustments.

Abstract: A system reverses degraded energy resolution of semiconductor radiation detection elements (44) which are used in a radiation detector assembly. A means (38) identifies semiconductor elements which exhibit degraded energy resolution as compared to an initial level of energy resolution after application of the forward bias. A means (40) restores the degraded semiconductor elements to the initial level of energy resolution by applying the reverse bias. A heater (74) accelerates the restoration process by supplying an elevated ambient temperature. A screening means (48) screens new semiconductor elements to identify the elements which are susceptible to degradation. A forward bias is applied by a forward bias means (50) to induce the degradation. A heater (52) increases an ambient temperature to accelerate the performance degradation in the new semiconductor elements. The identified degradable elements are treated with a reverse bias prior to installation in the detector.

Abstract: A nuclear imaging apparatus includes a radiation detector comprising an array (18) of solid state detector elements (22) responsive to incident gamma radiation by emitting a current spike. A pixel correction processor (44) detects defective detector elements in the array and a flood correction circuit (66) corrects detected radiation events (70) based on sensitivity differences between a plurality of groupings of detector elements in the array. A reconstruction processor (76) reconstructs an image representation from the corrected radiation events (74).

Abstract: A subject (10) is disposed adjacent a detector array (18) for the purposes of nuclear imaging. The subject (10) is injected with a radioactive isotope (14) and &ggr;-ray emissions indicative of nuclear decay are detected at the detector array (18). P-ASIC (60) preamplifier circuits are complex low-noise integrated circuits which dissipate a considerable amount of power (300-500 mW each). These components account for most of the dissipated power on the daughter cards (62). In order to facilitate the cooling of these electrical components, they are mounted on circuit boards (62) that are arranged parallel to each other extending perpendicularly away from the detector array (18). This provides channels between the boards through which cooling air is drawn by an array of fans (84).

Abstract: A subject (10) is disposed adjacent a detector array (18). The subject (10) is injected with a radioactive isotope (14) and &ggr;-ray emissions indicative of nuclear decay are detected at the detector array (18). The detector array generates electrical signals in response to each &ggr;-ray which signals are processed (64) and reconstructed (46) into an image representation of the anatomy of the subject (10). A high voltage bias is applied across the detector array. The bias is applied by a set of bias strips (80) and an electrically isolated common busbar (82) built onto a sheet of flexible circuit material. This flexible circuit (81) is highly transmissive to gamma radiation in the energy range 60-180 keV which is typically used in diagnostic nuclear medicine. Connections between the common busbar (82) and the bias strips (80) are made by resistors (92) on individual detector cards.

Abstract: A subject (10) is disposed adjacent a detector array (18) for the purposes of nuclear imaging. The subject (10) is injected with a radioactive isotope (14) and &ggr;-ray emissions indicative of nuclear decay are detected at the detector array (18) these signals are processed and reconstructed into an image representation of the anatomy of the subject (10). A dual level arbitration system orders detected signals for ease of processing and efficiency of reconstruction. The first level of the arbiter monitors a group of individual detectors (22). It locks out any signal that arrives from its group of detectors if a previous signal is still being analyzed. This avoids paralyzation of the system. The second level of the arbiter consists of a plurality of memories, one for each group of individual detectors (22) that store an address and energy of each processed signal.