Abstract: Systems and methods of dynamic PET imaging are disclosed. A system includes a positron emission tomography (PET) imaging modality configured to execute a first scan to acquire a first PET dataset and a processor. The first PET dataset includes dynamic PET data. The processor is configured to back-project the first PET dataset to generate a plurality of histo-image frames, input each of the plurality of histo-image frames to a trained neural network, and receive a dynamic PET output from the trained neural network. Each of the histo-image frames corresponds to a first axial position of the PET imaging modality.

Abstract: Various systems and computer-implemented methods for background radiation based attenuation correction are disclosed. Nuclear scan data including scan data associated with a first imaging modality and background radiation data are received. An initial background radiation attenuation map is generated and provided to a trained model configured to generate a final background radiation based attenuation map from the initial background radiation attenuation map. Attenuation correction of the scan data associated with the first imaging modality is performed based on the background radiation based attenuation map and a nuclear image is reconstructed from attenuation corrected scan data associated with the first imaging modality.

Abstract: Various systems and computer-implemented methods for background radiation based attenuation correction are disclosed. A first set of nuclear scan data including first scan data associated with a first imaging modality having a long-axial field of view and first background radiation data is received and a first background radiation attenuation map is generated by applying a trained machine-learning model to the first background radiation data. A first set of attenuation corrected scan data is generated by performing attenuation correction of the first scan data based only on the first background radiation attenuation map and a first image is reconstructed from the first set of attenuation corrected scan data. The disclosed background radiation based attenuation correction may be used for longer duration scans, repeat scans, and/or low-dose clinical applications, such as pediatric applications, theranostics, and/or other suitable applications.

Abstract: Apparatuses, computer-readable mediums, and methods are provided. In one embodiment, a positron emission tomography (“PET”) detector array is provided which includes a plurality of crystal elements arranged in a two-dimensional checkerboard configuration. In addition, there are empty spaces in the checkerboard configuration. In various embodiments, the empty spaces are filled with passive shielding, transmission source assemblies, biopsy instruments, surgical instruments, and/or electromagnetic sensors. In various embodiments, the crystal elements and the transmission source assemblies simultaneously perform emission/transmission acquisitions.

Abstract: Apparatuses, computer-readable mediums, and methods are provided. In one embodiment, a positron emission tomography (“PET”) detector array is provided which includes a plurality of crystal elements arranged in a two-dimensional checkerboard configuration. In addition, there are empty spaces in the checkerboard configuration. In various embodiments, the empty spaces are filled with passive shielding, transmission source assemblies, biopsy instruments, surgical instruments, and/or electromagnetic sensors. In various embodiments, the crystal elements and the transmission source assemblies simultaneously perform emission/transmission acquisitions.

Abstract: Apparatuses, computer-readable mediums, and methods are provided. In one embodiment, a positron emission tomography (“PET”) detector array is provided which includes a plurality of crystal elements arranged in a two-dimensional checkerboard configuration. In addition, there are empty spaces in the checkerboard configuration. In various embodiments, the empty spaces are filled with passive shielding, transmission source assemblies, biopsy instruments, surgical instruments, and/or electromagnetic sensors. In various embodiments, the crystal elements and the transmission source assemblies simultaneously perform emission/transmission acquisitions.

Abstract: Determining the position of a radioactive source in a PET system. Detecting a scatter coincidence event characterized by a full-energy photon detected at a first detector and partial-energy photon at a second detector. Measuring the arrival time difference between the partial energy photon and the full energy photon. Measuring the energy of the partial-energy photon. Determining a scattering point as a function of the position of the first detector, the position of the second detector, the energy of the partial-energy photon, the energy of an unscattered photon, the mass of a scattering electron, and the speed of light. Determining the position of a radioactive PET source along a line between the scatter point and the first detector as a function of the distance between scatter point and the first detector, the distance between scatter point and the second detector, and the measured time difference.

Abstract: Apparatuses, computer-readable mediums, and methods are provided. In one embodiment, a positron emission tomography (“PET”) detector array is provided which includes a plurality of crystal elements arranged in a two-dimensional checkerboard configuration. In addition, there are empty spaces in the checkerboard configuration. In various embodiments, the empty spaces are filled with passive shielding, transmission source assemblies, biopsy instruments, surgical instruments, and/or electromagnetic sensors. In various embodiments, the crystal elements and the transmission source assemblies simultaneously perform emission/transmission acquisitions.

Abstract: A method for processing an image which has the steps of a) receiving acquired data necessary to obtain an image and estimating a preliminary image; b) selecting at least one image element within the image; c) performing an iterative algorithm for processing the image at least on the at least one image element; d) computing a difference between the processed at least one image element and the at least one image element; and e) repeating the steps c) and d) until the difference is below a predefined threshold.

Abstract: For patient transmission data acquired simultaneously with patient emission data, blank transmission data are acquired in the absence of the patient emission and therefore under count rate conditions different from the count rate conditions of the patient transmission data. To prevent the different count rate conditions from causing artifacts in reconstructed tomographic images, a correction is made for spatially varying count rate effects on the attenuation correction. For example, the blank scan data are adjusted according to the count rate at which the patient emission data are acquired, and the adjusted blank scan data and the patient transmission data are used for attenuation correction of the patient emission data used for reconstructing a tomographic image.

Abstract: A method for processing an image which has the steps of a) receiving acquired data necessary to obtain an image and estimating a preliminary image; b) selecting at least one image element within the image; c) performing an iterative algorithm for processing the image at least on the at least one image element; d) computing a difference between the processed at least one image element and the at least one image element; and e) repeating the steps c) and d) until the difference is below a predefined threshold.

Abstract: Fast reconstruction methods are provided for 3D time-of-flight (TOF) positron emission tomography (PET), based on 2D data re-binning. Starting from pre-corrected 3D TOF data, a re-binning algorithm estimates for each transaxial slice the 2D TOF sinogram. The re-binned sinograms can then be reconstructed using any algorithm for 2D TOF reconstruction. A TOF-FORE (Fourier re-binning of TOF data) algorithm is provided as an approximate re-binning algorithm obtained by extending the Fourier re-binning method for non-TOF data. In addition, two partial differential equations are identified that must be satisfied by consistent 3D TOF data, and are used to derive exact re-binning algorithms and to characterize the degree of the approximation in TOF-FORE.

Abstract: A method for acquiring PET images with reduced time coincidence window limits includes obtaining a preliminary image of a patient within a radiation field of view (FOV), determining the spatial location of the patient within the FOV based on the preliminary image, calculating a different time coincidence window based on the spatial location of the patient for each possible pair of oppositely disposed detectors, scanning the patient with a PET scanning system to detect a pair of gamma photons produced by an annihilation event, determining whether the detection of the pair of gamma photons occurs within the time coincidence window, accepting the detected event only if the detection of the gamma photons occurs within the time coincidence window, calculating the spatial location of accepted annihilation event, and adding the calculated spatial location of the annihilation event to a stored distribution of calculated annihilation event spatial locations representing the distribution of radioactivity in the patient

Abstract: For patient transmission data acquired simultaneously with patient emission data, blank transmission data are acquired in the absence of the patient emission and therefore under count rate conditions different from the count rate conditions of the patient transmission data. To prevent the different count rate conditions from causing artifacts in reconstructed tomographic images, a correction is made for spatially varying count rate effects on the attenuation correction. For example, the blank scan data are adjusted according to the count rate at which the patient emission data are acquired, and the adjusted blank scan data and the patient transmission data are used for attenuation correction of the patient emission data used for reconstructing a tomographic image.

Abstract: A method for acquiring PET images with reduced time coincidence window limits includes obtaining a preliminary image of a patient within a radiation field of view (FOV), determining the spatial location of the patient within the FOV based on the preliminary image, calculating a different time coincidence window based on the spatial location of the patient for each possible pair of oppositely disposed detectors, scanning the patient with a PET scanning system to detect a pair of gamma photons produced by an annihilation event, determining whether the detection of the pair of gamma photons occurs within the time coincidence window, accepting the detected event only if the detection of the gamma photons occurs within the time coincidence window, calculating the spatial location of accepted annihilation event, and adding the calculated spatial location of the annihilation event to a stored distribution of calculated annihilation event spatial locations representing the distribution of radioactivity in the patient

Abstract: Fast reconstruction methods are provided for 3D time-of-flight (TOF) positron emission tomography (PET), based on 2D data re-binning. Starting from pre-corrected 3D TOF data, a re-binning algorithm estimates for each transaxial slice the 2D TOF sinogram. The re-binned sinograms can then be reconstructed using any algorithm for 2D TOF reconstruction. A TOF-FORE (Fourier re-binning of TOF data) algorithm is provided as an approximate re-binning algorithm obtained by extending the Fourier re-binning method for non-TOF data. In addition, two partial differential equations are identified that must be satisfied by consistent 3D TOF data, and are used to derive exact re-binning algorithms and to characterize the degree of the approximation in TOF-FORE.

Abstract: A normalization apparatus and method for a PET scanner with panel detectors for obtaining an estimate of a normalization array, for correction for count rate effects on the normalization array, and for measurement of the relation between the normalization array and the count rate. The method of the present invention is based on two quasi-independent radial and axial components, which are count rate dependent due to sensitivity changes across the detector blocks. A scatter source is disposed at the center of the FOV and a scatter-free source is disposed at the outer edge of the FOV. The method computes the normalization array through several steps which evaluate the geometric profile, the axial profile, and the correction factor. A count rate correction is introduced to extend the normalization array to any count rate.

Abstract: A normalization apparatus and method for a PET scanner with panel detectors for obtaining an estimate of a normalization array, for correction for count rate effects on the normalization array, and for measurement of the relation between the normalization array and the count rate. The method of the present invention is based on two quasi-independent radial and axial components, which are count rate dependent due to sensitivity changes across the detector blocks. A scatter source is disposed at the center of the FOV and a scatter-free source is disposed at the outer edge of the FOV. The method computes the normalization array through several steps which evaluate the geometric profile, the axial profile, and the correction factor. A count rate correction is introduced to extend the normalization array to any count rate.