Koninklijke Philips Corporation N.v.

University Of Washington

Koninkluke Philips N.v.

Thomas Leroy Laurence

Sharon Xiaorong Wang

Jerome John Griesmer

Steven E Cooke

Zhiqiang Hu

David Sowards-Emmerd

Jeffrey Allan Kolthammer

Jinghan Ye

Lingxiong Shao

Wenli Wang

Xiyun Song

Steffen Renisch

Thomas Blaffert

Andreas Thon

Steven J Plummer

Donald R Wellnitz

Robert S Miyaoka

Ralph Brinks

Carsten Degenhardt

Jinling Liu

Sharon X Wang

Sharon Xiaorong Wong

Sharon Xiarong Wang

Adrienne Lehnert

William Hunter

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=CS whererepresents TOF time differences, C is a relational matrix encoding the radiation detector elements, and S represents the TOF corrections. A pseudo-inverse Cof relational matrix C may be computed to solve S=C. 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 &#x221a;{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.

Time of flight calibration in digital positron emission tomography

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.

Energy calibration with Lu spectrum subtraction

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 Cof relational matrix C may be computed to solve S=CFormula 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.

A medical nuclear imaging system () and method () generate smooth energy histograms. Radiation events are detected by a plurality of detectors (), the radiation events localized to a plurality of pixels of the detectors (). 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.

Histogram smoothing in positron emission tomography (PET) energy histograms

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 () 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 () adjacent the first end surface and a reflective portion () adjacent the second end surface. A second layer () is disposed in a second plane orthogonal to the first plane and disposed between and adjacent facing side surfaces of the four substantially identical scintillation crystal bars with a light sharing portion () adjacent the second end surface and a reflective portion () adjacent the first end surface.

Pet detector scintillator arrangement with light sharing and depth of interaction estimation

A system () 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 () 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.

Digital positron emission tomography (DPET) energy calibration method

A system () and a method () 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.

Dead pixel compensation in positron emission tomography (PET)

A positron emission tomography (PET) apparatus and method employs a plurality of radiation detectors () disposed around an imaging region () 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 () with a time offset (); 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.

Pixel based dead time correction

A diagnostic imaging system includes a plurality of radiation detectors () configured to detect radiation events emanating from an imaging region. The system includes a calibration phantom () 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 () receives input of the radiation detectors and calculates an incidence angle independent crystal delay Tfor each detector. The calibration processor () 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 Tand &#x3b7;.

Pet detector timing calibration

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

Dead pixel identification in positron emission tomography (PET)

A positron emission tomography (PET) system () and method () classifies gamma events. At least one processor () 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) (). 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.

High spatial resolution mode solid state positron emission tomography (PET) scanner

A scintillator element () includes scintillator blocks () arranged to form an array, and transparent or translucent material () 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 &#x2022;=1.8, and the transparent or translucent material has a refractive index for the scintillation light of at least &#x2022;=1.6. An array of light detectors (), 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 () configured to detect 511 keV radiation.

Large-area scintillator element and radiation detectors and radiation absorption event locating systems using same

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

Scatter reject method via energy calibration

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.

Temperature stability for a digital positron emission tomography (PET) detector

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.

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North Royalton, OH, United States of America

Thomas Leroy Laurence