Abstract: When detecting scintillation events in a nuclear imaging system, time-stamping and energy-gating processing is incorporated into autonomous detection modules (ADM) (14) to reduce downstream processing. Each ADM (14) is removably coupled to a detector fixture (13), and comprises a scintillation crystal array (66) and associated light detect or (s) (64), such as a silicon photomultiplier or the like. The light detector(s) (64) is coupled to a processing module (62) in or on the ADM (14), which performs the energy gating and time-stamping.

Abstract: A preclinical positron emission tomography (PET) imaging method includes acquiring time-of-flight localized PET imaging data from one or more non-human animal subjects and reconstructing the acquired data to form an image. In an illustrative PET scanner includes: radiation detectors (12) viewing an examination region; a subject support assembly (14) supporting a plurality of preclinical subjects in the examination region for simultaneous PET imaging; coincidence electronics (20) acquiring time-of-flight localized PET imaging data from the preclinical subjects using the radiation detectors; and reconstruction electronics (22) that (i) perform a filtering operation based at least in part on the time-of flight information, the filtering operation including at least one of discarding non-probative time-of-flight localized PET imaging data and associating time-of-flight localized PET imaging data with individual preclinical subjects and (ii) reconstruct the filtered data to form images of the preclinical subjects.

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 color imaging device comprises: one or more arrays (10, RA, GA, BA) of color selective photodetectors (R, G, B) configured to acquire a color image of a subject; a set of avalanche photodiode photodetectors (APD) arranged to acquire a luminance image of the subject; and digital image processing circuitry (30) configured to process the acquired color image and the acquired luminance image to generate an output image of the subject. In some embodiments the avalanche photodiode photodetectors are configured to perform photon counting. In some embodiments, the one or more arrays comprise an imaging array (10) including the color-selective photodetectors (R, G, B) distributed across the imaging array with the set of avalanche photodiode photodetectors (APD) interspersed amongst the color-selective photodetectors.

Abstract: A nuclear medical imaging system employing radiation detection modules with pixelated scintillator crystals includes a scatter detector (46) configured to detect and label scattered and non-scattered detected radiation events stored in a list mode memory (44). Coincident pairs of both scattered and non-scattered radiation events are detected and the corresponding lines of response (LOR) are determined. A first image representation of the examination region can be reconstructed using the LORs corresponding to both scattered and non-scattered detected radiation events to generate a lower resolution image (60) with good noise statistics. A second higher resolution image (62) of all or a subvolume of the examination region can be generated using LORs that correspond to non-scattered detected radiation events. A quantification processor is configured to extract at least one metric, e.g.

Abstract: A PET scanner (8) includes a ring of detector modules (10) encircling an imaging region (12). Each of the detector modules includes at least one detector pixel (24,34). Each detector pixel includes a scintillator (20, 30) optically coupled to one or more sensor APDs (54) that are biased in a breakdown region in a Geiger mode. The sensor APDs output a pulse in response to the light from the scintillator corresponding to a single incident radiation photon. A reference APD (26, 36) also biased in a break-down down region in a Geiger mode is optically shielded from light and outputs a temperature dependent signal. At least one temperature compensation circuit (40) adjusts a bias voltage applied to the sensor APDs based on the temperature dependent signal.

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 medical imaging system employing radiation detection modules with pixelated scintillator crystals includes a scatter detector (46) configured to detect and label scattered and non-scattered detected radiation events stored in a list mode memory (44). Coincident pairs of both scattered and non-scattered radiation events are detected and the corresponding lines of response (LOR) are determined. A first image representation of the examination region can be reconstructed using the LORs corresponding to both scattered and non-scattered detected radiation events to generate a lower resolution image (60) with good noise statistics. A second higher resolution image (62) of all or a subvolume of the examination region can be generated using LORs that correspond to non-scattered detected radiation events. A quantification processor is configured to extract at least one metric, e.g.

Abstract: A PET scanner (8) includes a ring of detector modules (10) encircling an imaging region (12). Each of the detector modules includes at least one detector pixel (24,34). Each detector pixel includes a scintillator (20, 30) optically coupled to one or more sensor APDs (54) that are biased in a breakdown region in a Geiger mode. The sensor APDs output a pulse in response to the light from the scintillator corresponding to a single incident radiation photon. A reference APD (26, 36) also biased in a break-down down region in a Geiger mode is optically shielded from light and outputs a temperature dependent signal. At least one temperature compensation circuit (40) adjusts a bias voltage applied to the sensor APDs based on the temperature dependent signal.

Abstract: A color imaging device comprises: one or more arrays (10, RA, GA, BA) of color selective photodetectors (R, G, B) configured to acquire a color image of a subject; a set of avalanche photodiode photodetectors (APD) arranged to acquire a luminance image of the subject; and digital image processing circuitry (30) configured to process the acquired color image and the acquired luminance image to generate an output image of the subject. In some embodiments the avalanche photodiode photodetectors are configured to perform photon counting. In some embodiments, the one or more arrays comprise an imaging array (10) including the color-selective photodetectors (R, G, B) distributed across the imaging array with the set of avalanche photodiode photodetectors (APD) interspersed amongst the color-selective photodetectors.

Abstract: When detecting scintillation events in a nuclear imaging system, time-stamping and energy-gating processing is incorporated into autonomous detection modules (ADM) (14) to reduce downstream processing. Each ADM (14) is removably coupled to a detector fixture (13), and comprises a scintillation crystal array (66) and associated light detect or (s) (64), such as a silicon photomultiplier or the like. The light detector(s) (64) is coupled to a processing module (62) in or on the ADM (14), which performs the energy gating and time-stamping.

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 preclinical positron emission tomography (PET) imaging method includes acquiring time-of-flight localized PET imaging data from one or more non-human animal subjects and reconstructing the acquired data to form an image. In an illustrative PET scanner includes: radiation detectors (12) viewing an examination region; a subject support assembly (14) supporting a plurality of preclinical subjects in the examination region for simultaneous PET imaging; coincidence electronics (20) acquiring time-of-flight localized PET imaging data from the preclinical subjects using the radiation detectors; and reconstruction electronics (22) that (i) perform a filtering operation based at least in part on the time-of flight information, the filtering operation including at least one of discarding non-probative time-of-flight localized PET imaging data and associating time-of-flight localized PET imaging data with individual preclinical subjects and (ii) reconstruct the filtered data to form images of the preclinical subjects.

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 medicine imaging system that includes a plurality of detectors arranged about an imaging region. A transmission source can be provided opposite the detectors and rotating about the imaging region to obtain different imaging angles. The nuclear imaging system provides for the ability to acquire high sensitivity transmission data with high emission data spatial resolution.