Abstract: Disclosed herein are systems and methods for real-time monitoring of patient position and/or location during a radiation treatment session. Images acquired of a patient during a treatment session can be used to calculate the patient's position and/or location with respect to the components of the radiation therapy system. One variation of a radiation therapy system includes a circular gantry with a rotatable ring coupled to a stationary frame, a therapeutic radiation source mounted on the rotatable ring, and a patient-monitoring imaging system mounted on the rotatable ring. The patient-monitoring system may have one or more image sensors or cameras disposed on the rotatable ring within a bore region of the radiation therapy system, and may be configured to acquire image data as the ring rotates.

Abstract: Disclosed herein are systems and methods for real-time monitoring of patient position and/or location during a radiation treatment session. Images acquired of a patient during a treatment session can be used to calculate the patient's position and/or location with respect to the components of the radiation therapy system. One variation of a radiation therapy system includes a circular gantry with a rotatable ring coupled to a stationary frame, a therapeutic radiation source mounted on the rotatable ring, and a patient-monitoring imaging system mounted on the rotatable ring. The patient-monitoring system may have one or more image sensors or cameras disposed on the rotatable ring within a bore region of the radiation therapy system, and may be configured to acquire image data as the ring rotates.

Abstract: Time of flight (TOF) corrections for radiation detector elements of a TOF positron emission tomography (TOF PET) scanner are generated by solving an over-determined set of equations defined by calibration data acquired by the TOF PET scanner from a point source located at an isocenter of the TOF PET scanner, suitably represented as matrix equation at ?t=CS where ?t represents TOF time differences, C is a relational matrix encoding the radiation detector elements, and S represents the TOF corrections. A pseudo-inverse C?1 of relational matrix C may be computed to solve S=C?1 ?t. TOF corrections can be generated for a particular type of detector unit by identifying the radiation detector elements in C by detector unit. Further, multi-photon triggering time stamps can be adjusted to first-photon triggering based on ?{square root over (P1/Pm)} where P1 is average photon count using first-photon triggering and Pm is a photon count using multi-photon triggering.

Abstract: The present application relates generally to positron emission tomography (PET). It finds particular application in conjunction with energy calibration of a digital PET (DPET) detector and will be described with particular reference thereto. In one aspect, a difference spectrum is produced by finding a difference between a background radiation spectrum with no radioactive source loaded and a calibration source radiation spectrum with a radioactive source loaded. The difference spectrum may then be used to identify an energy peak.

Abstract: Time of flight (TOF) corrections for radiation detector elements of a TOF positron emission tomography (TOF PET) scanner are generated by solving an over-determined set of equations defined by calibration data acquired by the TOF PET scanner from a point source located at an isocenter of the TOF PET scanner, suitably represented as matrix equation Formula I=CS where Formula I represents TOF time differences, C is a relational matrix encoding the radiation detector elements, and S represents the TOF corrections. A pseudo-inverse C?1 of relational matrix C may be computed to solve S=C?1 Formula I. TOF corrections can be generated for a particular type of detector unit by identifying the radiation detector elements in C by detector unit. Further, multi-photon triggering time stamps can be adjusted to first-photon triggering based on Formula II where P1 is average photon count using first-photon triggering and Pm is a photon count using multi-photon triggering.

Abstract: A medical nuclear imaging system (10) and method (100) generate smooth energy histograms. Radiation events are detected by a plurality of detectors (14), the radiation events localized to a plurality of pixels of the detectors (14). The energy levels of the detected radiation events are estimated and the estimated energy levels are scaled with scaling parameters that scale the energy centroids of the plurality of pixels to target values differing by offsets around a common target value, the target values differing with spatial location of the plurality of pixels. Target value offsets are removed from the scaled energy levels and the detected radiation events are combined into an energy histogram using the energy levels with the target value offsets removed.

Abstract: A photon detector includes a sensor array of optical sensors disposed in a plane and four substantially identical scintillation crystal bars. Each optical sensor is configured to sense luminescence. Each of the four scintillator crystal bars being a rectangular prism with four side surfaces and first and second end surfaces, each scintillation bar has two side surfaces which each face a side surface of another scintillation bar, and each scintillation crystal bar generating a light scintillation in response to interacting with a received gamma photon. A first layer (80) is disposed in a first plane disposed between and adjacent facing side surfaces of the four substantially identical scintillation crystal bars with a light sharing portion (82) adjacent the first end surface and a reflective portion (84) adjacent the second end surface.

Abstract: A system (10) and method for energy correction of positron emission tomography (PET) event data by at least one processor. Event data for a plurality of strike events corresponding to gamma events is received. Each strike event is detected by a pixel of a detector module (50) and includes an energy and a time. The energy of the strike events is linearized using an energy linearity correction model including one or more parameters. Clusters of the strike events are identified based on the times of the strike events, and sub-clusters of the clusters are identified based on the pixels corresponding to the strike events of the clusters. Energies of the sub-clusters are corrected using a first set of correction factors, and energies of clusters including a plurality of sub-clusters are corrected using a second set of correction factors.

Abstract: A positron emission tomography (PET) apparatus and method employs a plurality of radiation detectors (20) disposed around an imaging region (16) and configured to detect 511 keV radiation events emanating from the imaging region. A calibration phantom is disposed in the imaging region. One or more processors are configured to: acquire and store listmode data of the phantom; measure a random rate for each line of response (LOR) from the listmode data using a coincident 511 keV events detector (34) with a time offset (54); determine a singles rate for each detector pixel from the random event rate, for example via a histogram plotting singles rate for each detector pixel; compute a live time factor of each LOR; compute a dead time correction factor as the reciprocal of the live time factor; and correct images according to the dead time correction factor.

Abstract: A system (10) and a method (100) compensate for one or more dead pixels in positron emission tomography (PET) imaging. A pixel compensation processor receives PET data describing a target volume of a subject. The PET data is missing data for one or more dead pixels. The pixel compensation estimates PET data for the dead pixels from the received PET data.

Abstract: A diagnostic imaging system includes a plurality of radiation detectors (20) configured to detect radiation events emanating from an imaging region. The system includes a calibration phantom (14) configured to be disposed in the imaging region spanning substantially an entire field of view and to generate radiation event pairs that define lines-of-response, wherein the calibration phantom is thin such that each LOR intersects the calibration phantom along its length, the thickness of the phantom being smaller than the length of the LORs. A calibration processor (24) receives input of the radiation detectors and calculates an incidence angle independent crystal delay Ti for each detector. The calibration processor (24) constructs a first look-up table for the timing correction of each LOR and a second look-up table for the angle depth of interaction correction for each crystal by combining Ti and ?i.

Abstract: A system (10) and a method (150) identify non-functioning pixels in positron emission tomography (PET) imaging. Data describing scintillation events localized to a plurality of pixels (22, 32) of a PET scanner (12) is received. A count map histogram is generated from the received data. The count map histogram maps each of the pixels (22, 32) to a count of scintillation events localized to the pixel (22, 32). One or more non-functioning pixels are identified from the count map histogram.

Abstract: A positron emission tomography (PET) system (10) and method (100) classifies gamma events. At least one processor (62, 66, 70) is programmed to receive event data for a plurality of scintillation events corresponding to gamma events. The gamma events are generated by gamma photons from a region of interest (ROI) (14). The gamma events of the event data are classified into a plurality of classifications. The classifications distinguish between single-crystal gamma events and multi-crystal gamma events.

Abstract: A scintillator element (22) includes scintillator blocks (60) arranged to form an array, and transparent or translucent material (62) disposed between adjacent scintillator blocks of the array. The transparent or translucent material may comprise epoxy or glue disposed between adjacent scintillator blocks of the array and adhering the adjacent scintillator blocks together. In some embodiments the scintillator blocks have a refractive index for scintillation light of at least •=1.8, and the transparent or translucent material has a refractive index for the scintillation light of at least •=1.6. An array of light detectors (24), such as silicon photomultipliers (SiPM) detectors formed monolithically on a silicon substrate, may be disposed on a bottom face of the scintillator element to detect scintillation light generated in the scintillator element. For PET applications, the scintillator element and the array of light detectors define a radiation detector (20) configured to detect 511 keV radiation.

Abstract: A medical nuclear imaging system (10) and corresponding method (100) are provided. A plurality of pixels (20, 24) detect radiation events and estimate the energy of the detected radiation events. A memory (58) stores a plurality of energy windows (44), the energy windows corresponding to the pixels. An event verification module (56) windows the radiation event with the energy windows corresponding to the detecting pixels. A reconstruction processor (60) reconstructs the windowed radiation events into an image representation.

Abstract: A detector maintains thermal stability between two different operating modes. The detector includes at least one controller which sets the detection sensitivity of the detector to a level disabling the detection of gamma photons. The controller further controls a heat generator to maintain the temperature of the detector at a predetermined temperature. The predetermined temperature is the steady state temperature of the detector when the detection sensitivity of the detector is set to a level enabling the detection of gamma photons. A method for maintaining thermal stability of a detector between two different operating modes is also provided. Approaches are also disclosed for normalize acquired imaging data during image reconstruction using dark current-dependent normalization factors.

Abstract: A positron emission tomography (PET) system (10) and method (100) classifies gamma events. At least one processor (62, 66, 70) is programmed to receive event data for a plurality of scintillation events corresponding to gamma events. The gamma events are generated by gamma photons from a region of interest (ROI) (14). The gamma events of the event data are classified into a plurality of classifications. The classifications distinguish between single-crystal gamma events and multi-crystal gamma events.

Abstract: A system (10) and method for energy correction of positron emission tomography (PET) event data by at least one processor. Event data for a plurality of strike events corresponding to gamma events is received. Each strike event is detected by a pixel of a detector module (50) and includes an energy and a time. The energy of the strike events is linearized using an energy linearity correction model including one or more parameters. Clusters of the strike events are identified based on the times of the strike events, and sub-clusters of the clusters are identified based on the pixels corresponding to the strike events of the clusters. Energies of the sub-clusters are corrected using a first set of correction factors, and energies of clusters including a plurality of sub-clusters are corrected using a second set of correction factors.

Abstract: When calibrating a positron emission tomography (PET) scanner, a radioactive calibration phantom is scanned over a period of several half lives to acquire a plurality of frames of scan data. Interlaced timing windows are employed to facilitate acquiring coincidence data for a plurality of coincidence timing windows and energy windows during a single calibration scan. Coincident events are binned according to each of a plurality of selected coincidence windows, and the PET scanner is calibrated for each of the plurality of coincidence timing windows using data acquired from the single calibration scan.

Abstract: When constructing a nuclear detector module in a gantry, a plurality of overlapping light guide modules (10) are mounted to the gantry in a spaced-apart fashion, and a plurality of underlapping light guide modules (12) are mounted in between each pair of overlapping light guide modules (10). Each of the underlapping modules and the overlapping modules includes a scintillation crystal array (16) on an interior surface thereof, and a plurality of PMTs on an exterior surface thereof. Overlapping modules (10) have overlapping structures (22) that interface with underlapping structures (18) on the underlapping modules (12) and thereby eliminate a seam directly beneath PMTs that overlap the crystal arrays of both an overlapping module and an underlapping module. Optical grease is used to form a resilient grease coupling and reduce light scatter between the underlapping and overlapping modules.