The disclosure is related to an apparatus and method for charged hadron therapy verification. The apparatus comprises a collimator comprising a plurality of collimator slabs of a given thickness, spaced apart so as to form an array of mutually slit-shaped openings, configured to be placed at a right angle to the beam line, so as to allow the passage of prompt gammas from the target, the collimator being defined at least by three geometrical parameters being the width and depth of the slit-shaped openings and a fill factor. The disclosure is also related to a method for charged hadron therapy verification with a multi-slit camera.

The disclosure is related to an apparatus and method for charged hadron therapy verification. The apparatus comprises a collimator comprising a plurality of collimator slabs of a given thickness, spaced apart so as to form an array of mutually slit-shaped openings, configured to be placed at a right angle to the beam line, so as to allow the passage of prompt gammas from the target, the collimator being defined at least by three geometrical parameters being the width and depth of the slit-shaped openings and a fill factor. The disclosure is also related to a method for charged hadron therapy verification with a multi-slit camera.

A method for measuring parameters of an analog signal to determine times at which the analog signal (S) crosses predetermined voltage thresholds (V.sub.A, V.sub.B, V.sub.C, V.sub.D), the method comprising the steps of: splitting the analog signal (S) into a number of interim signals (S.sub.A, S.sub.B, S.sub.C, S.sub.D), the number of the interim signals corresponding to the number of the preset voltage thresholds (V.sub.A, V.sub.B, V.sub.C, V.sub.D); providing an FPGA system (10) comprising differential buffers (11 A, 11 B, 11 C, 11 D) with outputs connected to a number of sequences (20A, 20B, 20C, 20D) of delay elements (21, 22, 23), the number of sequences of delay elements corresponding to the number of the preset voltage thresholds (V.sub.A, V.sub.B, V.sub.C, V.sub.D); inputting, to an input of each differential buffer (11 A, 11 B, 11 C, 11 D), one interim signal (S.sub.A, S.sub.B, S.sub.C, S.sub.D) and a reference voltage corresponding to a particular preset voltage threshold (V.sub.A, V.sub.B, V.sub.C, V.sub.D); reading, by means of vector generators (31 A, 31 B, 31 C, 31 D), assigned separately to each of the sequences (20A, 20B, 20C, 20D) and connected to a common clock signal (CLK), current values of output signals of each of the delay elements (21, 22, 23) in the particular sequence (20A, 20B, 20C, 20D) at the same moment for all vector generators and providing these values as sequence output vectors (W.sub.A, W.sub.B, W.sub.C, W.sub.D); and determining times at which the analog signal (S) crosses the predetermined voltage thresholds (V.sub.A, V.sub.B, V.sub.C, V.sub.D) on the basis of the values of the sequence output vectors (W.sub.A, W.sub.B, W.sub.C, W.sub.D) and the delays introduced by the delay elements (21, 22, 23).

A method for determining parameters of a reaction of a gamma quantum within a scintillator of a PET scanner, comprising transforming a signal measured in the scintillator using at least one converter into an electric measurement signal, wherein the method comprises the steps of: obtaining access to a reference parameters memory (10) comprising reference signals represented in a time-voltage (Wt-v) coordinate system and in a time-amplitude fraction (Wt-f) coordinate system and having associated reaction parameters; sampling the electric measurement signal (S) measured in the time-voltage (PT-V) coordinate system and in the time-amplitude fraction (Pt-f) coordinate system; comparing results of the sampling (PT-V, PM) of the electric measurement signal (S) with the reference signals (Wt-V, Wt-f) and selecting reference shape parameters so that the reference (W) is best fitted to the results of the sampling (PT-V, PM) of the electric measurement signal (S); and determining the parameters of the reaction of the gamma quantum within the scintillator (1) for the electric measurement signal (S) based on pre-calibrated functions that determine the values of parameters of signal shape depending on the parameters of the reaction of gamma quantum within the scintillator.

A hybrid TOF-PET/CT tomograph comprising a detection chamber, gamma radiation detectors, X-ray detectors and a movable X-ray source, wherein the gamma radiation detectors (150, 250, 350, 450, 550) and the X-ray detectors (170, 270, 370, 470, 570) surround the detection chamber (102, 202, 302, 402, 502) around the whole perimeter of the detection chamber (102, 202, 302, 402, 502), and wherein the gamma radiation detectors (150, 250, 350, 450, 550) are located closer to the longitudinal axis (115, 215, 315, 415, 515) of the detection chamber (102, 202, 302, 402, 502) than the X-ray detectors (170, 270, 370, 470, 570), and wherein the gamma radiation detectors (150, 250, 350, 450, 550) comprise polymer strips (151, 251, 351, 451, 551) made of a scintillation material having a density lower than the density of the X-ray radiation detectors (171, 271, 371, 471, 571).

A method and apparatus for generating crystal efficiency correction factors by performing a normalization calibration based on delayed data. The method and apparatus obtain delayed data from a scan of a patient using a Positron Emission Tomography (PET) scanner, generate a sinogram from the obtained delayed data, determine, using a processing circuit, mean fan and block line of response sensitivities from the generated sinogram, determine, using the processing circuit, mean detector efficiency based on the determined mean fan and block line of response sensitivities, determine, using the processing circuit, an individual crystal efficiency based on the determined mean fan and block line of response sensitivities and the mean detector efficiency for each module, and calculate the crystal efficiency correction factors based on the determined individual crystal efficiency of each module.

The invention relates to a detection arrangement, a detection system comprising said arrangement, and a method of processing data from said arrangement. The detector arrangement disclosed comprises at least one array of detectors, wherein the detectors are configured to detect photons emitted from an object as a result of positron annihilation due to irradiation of the object with photons of a predetermined energy. Each detector in the array is linked to or associated with one or more other detector in the array to define a region of interest (RoI). The detector arrangement comprises or is communicatively coupled to a coincidence trigger unit which is configured to register or determine a coincidence in response to receiving detection signals from two different detectors forming part of the same RoI and indicating detection of substantially back-to-back co-linear and co-incident photons in the RoI.

There is provided an x-ray detector system including a photon-counting x-ray detector for detecting x-ray radiation from an x-ray source, and a coincidence detection system configured to determine and/or obtain information about the radiation incident on the x-ray detector based on information about the time of photon interactions in the x-ray detector and information about the location of the x-ray source in relation to the x-ray detector. There is also provide an x-ray imaging system including such an x-ray detector system, as well as a corresponding coincidence detection system and a corresponding method.

A method for calibration of TOF-PET detectors comprising polymeric scintillator strips and photoelectric converters, wherein cosmic radiation is used as a source of radiation, the method comprising the steps of: recording times of reactions of particles of cosmic radiation with the scintillator strips (101, 411, 421, 511, 521); determining spectra (301) of distribution of differences in the times at which pulses are recorded at ends of the scintillator strips (101, 411, 421, 511, 521) connected to photoelectric converters (102, 103, 412, 413, 422, 423, 512, 513, 522, 523); using the determined spectra (301) to determine timing synchronization constants of the photoelectric converters (102, 103, 412, 413, 422, 423, 512, 513, 522, 523), the constants being related to: delays within the electronics; speed of light propagation within the scintillator strip of the detection module; and resolution of the difference in times of the signals recorded at the ends of the module.

A novel proton imaging system incorporates positron emission detections to enhance proton therapy treatment preparation and procedural efficiencies while reducing operational costs associated with proton therapy. In one case, the novel proton imaging system incorporating a positron emission tomography (PET) module enables rapid on-the-fly in vivo range verification for proton therapy using position information from short-lived positron emitters produced during treatment. This unique in vivo range verification method produces more streamlined, accurate, and cost-effective results relative to conventional proton imaging systems. In another case, the novel proton imaging system incorporating the PET module provides a unique combinatory PET/pCT (proton computer tomography) scanning that creates more accurate maps for proton therapy planning for metabolically-active tumors. This proton imaging system also utilizes a novel concept of “virtual protons” originating from in vivo range verification measurements that mimic proton particle's characteristics for more accurate proton computer tomography (pCT) or computer tomography (CT).