EP4241117A1 - Système de codeur de tomographie par émission de positrons à résolution temporelle - Google Patents

Système de codeur de tomographie par émission de positrons à résolution temporelle

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Publication number
EP4241117A1
EP4241117A1 EP21889769.2A EP21889769A EP4241117A1 EP 4241117 A1 EP4241117 A1 EP 4241117A1 EP 21889769 A EP21889769 A EP 21889769A EP 4241117 A1 EP4241117 A1 EP 4241117A1
Authority
EP
European Patent Office
Prior art keywords
event
time
diametrically opposed
positron
scintillation detectors
Prior art date
Legal status (The legal status is an assumption and is not a legal conclusion. Google has not performed a legal analysis and makes no representation as to the accuracy of the status listed.)
Pending
Application number
EP21889769.2A
Other languages
German (de)
English (en)
Other versions
EP4241117A4 (fr
Inventor
Ronald Nutt
J. Nutt Lynda
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Individual
Original Assignee
Individual
Priority date (The priority date is an assumption and is not a legal conclusion. Google has not performed a legal analysis and makes no representation as to the accuracy of the date listed.)
Filing date
Publication date
Priority claimed from US17/093,095 external-priority patent/US11054534B1/en
Application filed by Individual filed Critical Individual
Priority claimed from PCT/US2021/036153 external-priority patent/WO2022098394A1/fr
Publication of EP4241117A1 publication Critical patent/EP4241117A1/fr
Publication of EP4241117A4 publication Critical patent/EP4241117A4/fr
Pending legal-status Critical Current

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Classifications

    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01T—MEASUREMENT OF NUCLEAR OR X-RADIATION
    • G01T1/00—Measuring X-radiation, gamma radiation, corpuscular radiation, or cosmic radiation
    • G01T1/29—Measurement performed on radiation beams, e.g. position or section of the beam; Measurement of spatial distribution of radiation
    • G01T1/2914—Measurement of spatial distribution of radiation
    • G01T1/2985—In depth localisation, e.g. using positron emitters; Tomographic imaging (longitudinal and transverse section imaging; apparatus for radiation diagnosis sequentially in different planes, steroscopic radiation diagnosis)
    • G—PHYSICS
    • G01—MEASURING; TESTING
    • G01T—MEASUREMENT OF NUCLEAR OR X-RADIATION
    • G01T1/00—Measuring X-radiation, gamma radiation, corpuscular radiation, or cosmic radiation
    • G01T1/16—Measuring radiation intensity
    • G01T1/20—Measuring radiation intensity with scintillation detectors
    • G01T1/2018—Scintillation-photodiode combinations
    • G01T1/20184—Detector read-out circuitry, e.g. for clearing of traps, compensating for traps or compensating for direct hits
    • G—PHYSICS
    • G04—HOROLOGY
    • G04F—TIME-INTERVAL MEASURING
    • G04F10/00—Apparatus for measuring unknown time intervals by electric means
    • G04F10/005—Time-to-digital converters [TDC]

Definitions

  • the present general inventive concept relates to a new Time-Resolved Positron
  • TPET Emission Tomography system
  • CT computed tomography
  • LOR Line of Response
  • a typical, state-of-the-art, prior art Ring PET tomograph will have several hundred detector elements in the Ring, see e.g., FIG. ID.
  • FIG. ID In the prior art there will be more than 100,000 LORs in a single Ring.
  • the LORS Using this set of LORs to form an image, the LORS must be arranged in an orderly format by a fast computer. The arrangement is called a sinogram and the process is called sorting.
  • FIG. 2A depicts a prior art PET imaging system as illustrated by Henseler ‘063.
  • the two major computers that are typically used in state-of-the-art image reconstruction, as illustrated in FIG. 2A, are represented by the box labeled Computer- Readable Medium 806 where the sorting process takes place and the memory of the sinograms are located; further, the processor 810 is the reconstruction processor and typically will be a powerful computer.
  • the reconstruction process will either be a statistical process or a back-projection process but will require several minutes with a fast computer to form an approximate image.
  • Time-of-flight (TOF) positron emission tomography (“TOF-PET”) is based on the measurement of the difference AT between the detection times of the two gamma photons arising from the positron annihilation event.
  • This measurement allows the annihilation event to be localized along the LOR with a resolution of about 75-120 mm FWHM, assuming a time resolution of 500- 800 ps (picoseconds).
  • this approximate localization is effective in reducing the random coincidence rate and in improving both the stability of the reconstruction and the signal -to-noise ratio (SNR), especially when imaging large objects.
  • SNR signal -to-noise ratio
  • the normal LOR is determined by standard block detectors that measure the gamma ray position, both X and Y coordinates, in two-dimensional space.
  • the third dimensional coordinate, Z in an XYZ coordinate system is then provided by the time measurements of the travel of the coincident gamma rays, and more specifically the difference between the time one gamma ray is detected and the time when the second gamma ray is detected.
  • This time measurement can be made by an analog technique known in the art as Time-to- Amplitude Converter (“TAC”).
  • TAC Time-to- Amplitude Converter
  • the voltage on the capacitor is then proportional to the time difference between the two signals.
  • the problem with this technique is that many events are occurring in a short period of time; and, in order to provide real-time, i.e. gamma event-by- gamma event analysis, the system must be able to measure all of these events, and, in realtime, detect and isolate those events where the time difference falls within the desired field of view.
  • a state-of-the-art analog TAC simply does not have this capability.
  • a gamma ray travels at the speed of light, i.e., 30 centimeters per nanosecond, and gamma rays travel in opposite directions for PET, 15 centimeters from the center corresponds to a time difference of one nanosecond.
  • the reconstructed resolution of conventional PET systems is approximately 5 to 6 millimeters in three dimensions so 6 millimeters along a LOR corresponds to 40 picoseconds. This is the needed time resolution of the TPET imaging system such that Z-coordinate position resolution is equivalent to the existing standard PET imaging systems.
  • the present PET scanners measure and record the LORs and then perform an image reconstruction by statistically determining the intersection of all LORs with each other.
  • the statistical process requires a very large number of events, and thus, a large number of detectors, and, therefore, is very inefficient.
  • This process requires many more events than does the direct measurement of the TPET imaging system of the present general inventive concept and therefore is much more inefficient than the TPET imaging system disclosed herein. This can be best understood for a simple point source of activity in the field of view where a single TPET measurement can result in an image. If the volume being considered has a complex distribution of activity, it is known in the PET art that the number of events needed for a good image may increase to millions of events for the conventional system.
  • the time measuring system currently known in the art uses either standard Photomultipliers or SiPM detectors such as the SiPM detectors supplied by Philips or Hamamatsu.
  • SiPM detectors such as the SiPM detectors supplied by Philips or Hamamatsu.
  • the known measurement systems which utilize an SiPM optical detector and an LSO scintillator, utilize a leading-edge discriminator and a simple frequency counter. It is known that this type of system is not capable of producing event-by-event, i.e., real-time, high resolution performance in the time domain or in the third dimension of the PET system for two additional reasons.
  • the first problem is the leading-edge timing used in the discriminators. This leading-edge technique uses an electronic discriminator that changes states when the leading edge of the pulse exceeds a defined level.
  • the state-of-the art time resolution is 250 picoseconds for the conventional PET system and is used to set the coincidence window for the standard PET system. This time resolution is set primarily by the LSO scintillator being used. It does not appear in the art that there has been any attempt to use this timing information to eliminate the image reconstruction. However, even if this timing information were used in the conventional PET system to eliminate the reconstruction process, the image resolution would be approximately 37 millimeters. Even if the TOF were used to eliminate image reconstruction, the result is an unacceptably low resolution in the diagnostic imaging industry.
  • time walk associated with leading edge discriminators is undesirable; and that, even in cases in which the amplitude of the light signal does not vary appreciably, residual time walk can have a negative impact on time resolution.
  • FIG. 3 if one assumes an input rise time of 9,000 picoseconds, i.e. 9 nanoseconds, the time walk due to pulse height changes can be very significant. In this regard, it is known in the timing area that the input current pulse height can vary V from 30 to 50%. If one assumes the input pulse has a linear rise-time, then the time walk associated with the leading-edge discriminator is given by;
  • the time resolution in this example would be the square root of two times 60ps, or 85ps.
  • the best resolution reported in the prior art is around lOOps. From this example, it is apparent that the time uncertainty due to time walk is significant and will have to be significantly reduced, or eliminated, in any practical time-resolved PET imaging system.
  • a second known time measuring system is the Digital Intervalometer described by the applicant herein in U.S. Patent No. 3,541,448, which is incorporated herein by reference, and which was issued to the applicant herein on Nov. 17, 1970.
  • This time measurement consists of a digital clock which counts the ticks on an oscillator between a start and a stop signal.
  • the time between the start and the oscillator is measured by a digital Time-to- Amplitude Converter, (“TAC”), and the difference between the oscillator and the stop signal is also measured with a digital TAC.
  • TAC Digital Time-to- Amplitude Converter
  • the most significant digits of the clock are provided by the digital TAC and added sequentially to the word containing the ticks of the oscillator. If the desired measurement is between a start and the oscillator, which is the case for TPET operation, only one of the TAC interpolators is used to measure the time of arrival of one of the 51 IKeV gamma rays.
  • a time-resolved PET imaging system offering event-by-event, real-time, high resolution imaging utilizing fewer detectors than traditional PET systems, and not requiring image reconstruction, is unavailable in the known art. Accordingly, it is a feature of the present general inventive concept to provide a Time-resolved PET system having a much simpler design than conventional PET imaging systems, therefore resulting in a much lower cost to construct, that provides high-resolution images event-by-event, in real time, and which eliminates the need for a CT scan for attenuation correction.
  • the primary feature of the present general inventive concept is to provide a TPET imaging system that utilizes time measurement information to determine a position along an LOR in order to provide a measurement of a third dimensional coordinate, resulting in near real-time, high resolution three-dimensional imaging in a TPET imaging system having fewer detectors with significantly fewer measured events, i.e., without detecting and measuring millions of events before forming an image.
  • Smaller probes using the TPET concept of the present general inventive concept can be used to image specific organs of the body without introducing artifacts, unlike the present PET systems that require all views around the patient. This allows very practical breast imaging, cardiac imaging, and prostate imaging, to name a few organ systems that could be imaged using the TPET imaging system of the present general inventive concept.
  • a TPET imaging system capable of providing event-by-event, i.e., real-time, high resolution, three- dimensional imagery is provided without the need for image reconstruction or CT correction of the attenuation of gamma rays.
  • the radiation detectors provide X and Y coordinates
  • the third dimensional coordinate i.e. the Z coordinate
  • TPET imaging system disclosed herein utilizes further electronic circuitry, which both measures the arrival time of each photon in the annihilation process with picosecond resolution and which decreases the time walk to an insignificant contribution of the overall time resolution is utilized.
  • decreasing the time walk is accomplished by use of a Constant Fraction Discriminator, (“CFD”), operating in conjunction with the Digital Intervalometer, for performing the time-measuring function, invented by the applicant herein.
  • CFD Constant Fraction Discriminator
  • Constant Fraction Discriminator and the Digital Intervalometer could be utilized as discrete circuits, or separate chip sets
  • Constant Fraction Discriminator and the Digital Intervalometer are embedded on an Application Specific Integrated Circuit (“ASIC”).
  • ASIC Application Specific Integrated Circuit
  • the arrival time of each photon in the annihilation process is recorded with respect to a clock frequency with picosecond resolution.
  • this TPET approach requires significantly fewer gamma events, thus requiring fewer detectors, thereby resulting in a system that is more efficient and more economical to produce.
  • the TPET imaging system of the present general inventive concept allows real-time image comparisons of multiple organs thereby giving the clinician/ diagnostician greater understanding of the relationships of multiple organ systems. Also, the system allows the observation in real time of cancer lesions that are being treated by radiation or other means.
  • FIG. 1 A illustrates, in a simplified schematic view, the physics of a positron event in which two gamma rays are produced that travel in directions that are 180° opposed to one another;
  • FIG. IB illustrates, in a simplified schematic view, Lines of Response, (“LORs”), from a single point source of positrons;
  • LORs Lines of Response
  • FIG. 1C illustrates, in a simplified schematic view, complications that arise in the process of locating the intersection of LORs to identify the exact location of the positrons if the image has two locations of positrons;
  • FIG. ID is a schematic view of a prior art PET Imaging system having a detector Ring and reliant upon image reconstruction;
  • FIG. 2A illustrates, in schematic view, a conventional PET imaging system as exemplified in U.S. Published Patent Application No. 2013/0009063;
  • FIG. 2B further illustrates, in schematic view, a conventional PET imaging system
  • FIG. 3 is chart illustrating the problem of time walk due to pulse height changes common in many of the standard PET systems
  • FIG. 4 is a chart illustrating the operational process of applicant’s Digital Intervalometer as utilized in an exemplary embodiment of the present general inventive concept
  • FIG. 5 is a schematic view of the TPET imaging system of the present general inventive concept
  • FIGS. 6A, 6B, and 6C are more detailed views of a TPET according to exemplary embodiments of the present general inventive concept
  • FIGS. 7A and 7B illustrate schematic views of the process of capturing a three- dimensional image, and determining the X, Y, and Z coordinates, in 7A, or vector coordinates in 7B, of the respective positron events, with a two-detector system according to an exemplary embodiment of the present general inventive concept;
  • FIG. 8 is a schematic view of an exemplary embodiment of a simpler TPET Imaging detector array system.
  • the normal LOR 20 is determined by standard block detectors 30A and 30B, which, as will be understood, are diametrically opposed to one another, and measure the gamma ray position, in both the X and Y coordinates, 40 and 50 respectively, in two-dimensional space as schematically illustrated in FIG. 7A.
  • FIG. 5 depicts a TPET imaging system of the present general inventive concept.
  • FIGS. 7A and 7B the positron events, such as positron event 75 in FIG. 7A, that is occurring in the organ 70, is illustrated schematically.
  • the arithmetic for a single positron event at Point B, in Fig. 7A can be represented as follows:
  • Zi is the location along the LOR.
  • ZLOR is calculated and obtained from a look-up table.
  • the time difference, AT, between the arrival times of coincident gamma rays detected by block detectors 30A and 30B determines the position of the positron event along the LOR 20.
  • FIG. 7A shows the results of the basic measurement including the X, Y and Z directions. The X and Y positions are determined by the conventional block detectors and the Z direction is determined by the difference in arrival times of the two gamma rays in opposing detectors. With the Z time measurement and the X and Y determination, the point in space where the event originated is determined and can be displayed.
  • the third dimension, Z coordinate 60 in an XYZ coordinate system see, e.g., FIG. 7A, is provided by the time measurement of the travel of the gamma ray and by calculating the difference between the time one gamma ray is detected and when the second coincident gamma ray is detected.
  • positron event 75 While measuring the gamma ray, i.e., positron event, position and the position itself has been discussed in Cartesian terms, regarding X, Y, and Z coordinates, in an exemplary embodiment, vector analysis, i.e., vector algebra, would be employed to determine the location of a positron event, such as positron event 75 in FIG. 7B.
  • processing circuitry will determine the position of the positron event 75 when the detector element of each detector plate, 30A and 30B, is measured and the time difference, AT, between the respective detector elements is measured. Referring to FIG. 7B, each vector will be defined with the X,Y and X '.Y 1 data points.
  • an exemplary array will be composed of approximately 40 detectors in the horizontal, or X, direction and 40 detectors in the vertical, or Y, direction. This produces an array having 1,600 detectors in each detector plate. These two detector plates are diametrically opposed to one another and are, in an exemplary embodiment, as illustrated in FIG. 8, parallel to one another. With that element count, there will be (l,600) 2 Lines of Response, (“LORs”). Each LOR will be defined before system use with the location of each memory pixel along the LOR vector. With such arrangement, a computer, having cubic memory would have the intersection of each LOR with each memory pixel stored in its memory.
  • positron event 75 When the time difference, AT, is measured and the LOR is identified, the location is then identified for each positron event such as positron event 75.
  • the computer uses vector analysis, instead of Cartesian coordinates, for forming the three-dimensional image.
  • the three points of a hypothetical vector would be defined, or expressed, as follows:
  • block detectors 30A and 30B could be either silicon photomultiplier, (SiPM), or conventional photomultipliers.
  • the TPET imaging system 10 of the present general inventive concept utilizes an SiPM for the optical detector and a very high yield photon, fast scintillator such as CeBrs (35,000 photons per 511KeV) or LaBn.
  • CeBrs scintillator the stopping power for gamma rays is worse than the usual PET scintillators; but this loss in stopping power is offset by the high inherent efficiency of the TPET.
  • the position resolution along the LOR 20 is calculated as follows:
  • This resolution in time corresponds to 5.8 mm resolution along the LOR 20.
  • the electronic circuitry 90 functions to have the trigger for time measurement occur at a fixed fraction of the input pulse. This is accomplished by placing an input pulse on two input lines, delaying, inverting, and amplifying the one input line. The resulting negative, delayed and amplified input is added back to the original input pulse. The result is a bipolar pulse with a zero crossing time which will occur at the same time of a fixed fraction of the pulse. As the amplitude of the pulse changes, the zero crossing time will remain.
  • time measuring electronic circuitry 100 in FIG. 6A, is provided which consists of a digital clock which counts the ticks on an oscillator between a start and a stop signal.
  • further electronic circuitry 100 measures the arrival time of each photon in the annihilation process with picosecond resolution.
  • the time measuring electronic circuitry 100 functions to count the ticks of an oscillator between a start and stop signal.
  • the limitation of the conventional, state- of-the-art system is that the accuracy is plus or minus ( ⁇ ) one tick of the oscillator.
  • the time from the start pulse to the oscillator is measured using an analog time-to-pulse height converter and this time is added back to the conventional pulse counter. The same process is used to eliminate the inaccuracy at the stop signal.
  • the phase between the start and the oscillator is measured by a digital TAC and the difference between the oscillator and the stop signal is also measured with a digital TAC.
  • the most significant digits of the clock are provided by the digital TAC and added sequentially to the word containing the ticks of the oscillator. If the desired measurement is between a start and the oscillator, which is the case for TPET operation, only one of the TAC interpolators is used to measure the time of arrival of one of the 51 IKeV gamma rays.
  • the electronic circuitry for decreasing the time walk is provided by a Constant Fraction Discriminator, (“CFD”), 90'.
  • CFD 90' decreases the time walk to an insignificant contribution of the overall time resolution of the TPET imaging system 10.
  • this picosecond resolution time-measuring electronic circuitry is defined by a Digital Intervalometer 100'. With this time measurement system, the arrival time of each photon in the annihilation process is recorded with respect of a clock frequency with picosecond resolution. With the proper design and the use of this time measurement, the time resolution will be insignificant to the overall measurement.
  • the Digital Intervalometer 100' utilizes the TAC to interpolate between the start signal and the clock.
  • the TAC is converted to a digital signal by standard Analog-to-Digital (“ADC”) method and this digital signal is added to the end of the digital clock word to form the time measurement.
  • ADC Analog-to-Digital
  • the stop signal is generated by an arbitrary fixed signal synchronized with the clock.
  • Fig. 4 it is known in the art that if no interpolation is used, the measured time resolution using only the counting of the clock pulses will result in a plus and minus 125ps for a clock frequency of 8 GHz.
  • Prior experimentation and research in the art has attempted to compensate for various issues with state-of-the-art timing discriminators and time encoders by doing the interpolation off-line and in so doing by attempting to measure the input rise-time and extrapolating the phase. This is not a practical approach in a clinical, diagnostic application.
  • the clinical diagnostic environment requires a TPET imaging system 10 that is capable of producing high resolution, three-dimensional imagery gamma event-by -gamma event without image reconstruction.
  • TPET imaging system 10 This is the case for the TPET imaging system 10.
  • the arrival time of each gamma ray event is measured and digitally stored and then digitally compared with all other measured events. Any two measured events that occur within a few picoseconds of each other will be defined as a positron annihilation event and used to form the three-dimensional image.
  • this approach requires significantly fewer gamma events, thus requiring fewer detectors, and gives rise to a gamma event-by -gamma event high resolution three-dimensional imagery. This results in a TPET imaging system 10 that is more efficient and more economical to produce than a conventional PET imaging system.
  • the TPET imaging system 10 of the present general inventive concept includes the use of electronic circuitry for measuring the difference of arrival time in order to determine the position of a positron event along an LOR 20, in a manner that allows for event-by-event, high resolution imagery and that also reduces time walk.
  • this further electronic circuitry is defined by both a Constant Fraction Discriminator 90 and a Digital Intervalometer 100.
  • Constant Fraction Discriminator 90 and the Digital Intervalometer 100 could be used as discrete chip sets or discrete integrated circuits, in an exemplary embodiment, the Constant Fraction Discriminator 90 and the Digital Intervalometer 100 are combined in an Application Specific Integrated Circuit, (“ASIC”), 80. While an exemplary embodiment that utilizes a Constant Fraction Discriminator and a Digital Intervalometer to perform these functions has been described, it will be understood that other electronic circuitry adapted to perform these functions could also be utilized especially as computing efficiency and power continue to increase.
  • ASIC Application Specific Integrated Circuit
  • the TPET imaging system of the present general inventive concept has the capability of meeting or even exceeding the image resolution of existing PET systems and will contribute three additional very important characteristics to the system.
  • the TPET imaging system 10 of the present general inventive concept will be, comparatively, very simple and therefore very inexpensive to manufacture compared to full ring modem PET tomographs.
  • the TPET imaging system 10 of the present general inventive concept can provide artifact free images without covering the entire body as is required by conventional PET imaging systems.
  • This feature makes the TPET imaging system 10 of the present general inventive concept feasible for individual organ imaging such as breast, heart, prostate, and imaging lungs for Coronavirus damage. This feature allows a design with a very small number of detectors.
  • the third main benefit is that the images are formed as each gamma event is detected, i.e., in an event-by-event manner, variously referred to herein as being “in real-time” without the inherent delay required by image reconstruction or attenuation correction.
  • This feature gives rise to event-by-event, real-time image comparisons of multiple organs, thereby giving the clinician/diagnostician greater understanding of the relationships of multiple organ systems, for example of the brain and heart.
  • the real-time feature provides the clinician a view of cancer sites as the site is being treated with radiation or other means such as Proton Therapy.
  • the TPET imaging system 10 of the present general inventive concept has the potential of replacing state-of-the-art PET systems and represents a major breakthrough for medical imaging in general.
  • the Avalanche Photodiode not only improves the time resolution of scintillator systems; but its resistance to magnetic fields allows the TPET imaging system 10 of the present general inventive concept to be used in a combined TPET and Magnetic Resonance Imaging (“MRI”) system for acquiring MRI and TPET images simultaneously in a single device, and will allow the TPET imaging system to operate in relatively high radiation fields.
  • MRI Magnetic Resonance Imaging
  • a fourth feature is that the TPET uses only one view, while in comparison the traditional PET uses many views around the patient. With only one view, calculated attenuation for the emitted gamma-rays can be achieved with good accuracy. This feature eliminates need for a CT scan for attenuation correction. This further reduces the cost and complexity of the TPET imaging system 10.
  • Various exemplary embodiments of the present general inventive concept may provide an event-by-event, high resolution, three-dimensional positron emission tomography encoder system which includes a plurality of cooperating pairs of diametrically opposed scintillation detectors adapted for receiving gamma rays from a positron event, electronic circuitry in electronic communication with each pair of said diametrically opposed scintillation detectors for determining the two-dimensional position of the positron event occurring between each pair of detectors, each event producing said gamma rays which travel along a line of response extending between cooperating pairs of detectors, and further electronic circuitry for measuring the difference between the arrival times of coincident gamma rays from a positron event detected by said diametrically opposed scintillation detectors along said line of response thereby enabling determination of the third dimension along said line of response of said positron event to produce an event-by-event, real-time, high resolution, three dimensional positron emission tom
  • the further electronic circuitry may include a Constant Fraction Discriminator and a Digital Intervalometer.
  • a Constant Fraction Discriminator and a Digital Intervalometer are utilized, such can be utilized as discrete chip sets or, in an exemplary embodiment, embedded on an application specific integrated circuit.
  • the digital intervalometer may utilize a time amplitude converter.
  • the pairs of diametrically opposed scintillation detectors may each include a silicon photomultiplier operating in an avalanche photodiode mode.
  • the pairs of diametrically opposed scintillation detectors may include a channel multiplier.
  • Various exemplary embodiments of the present general inventive concept may provide a process for capturing event-by-event, high resolution, three-dimensional positron emission tomography imagery, in real-time, which includes providing a selected number of cooperating pairs of diametrically opposed scintillation detectors adapted for receiving coincident gamma rays from a positron event, determining a two-dimensional position of the positron event producing said gamma ray photons, determining a line of response for each cooperating pair of detectors of said coincident gamma rays, and detection of and measuring, in real-time, the time of said coincident gamma rays being detected by a cooperating pair of diametrically opposed scintillation detectors along said line of response in order to determine the third dimension along said line of response of said positron event, thereby producing a real-time, high resolution, three dimensional positron emission tomographic image.
  • the process for capturing real-time, high resolution, three-dimensional positron emission tomography imagery may include a Constant Fraction Discriminator and a Digital Intervalometer, and may include these specific integrated circuits utilized as discrete chip sets or, in an exemplary embodiment, embedded on an application specific integrated circuit.
  • the digital intervalometer may utilize a time-to-amplitude converter.
  • the cooperating pairs of diametrically opposed scintillation detectors may each include a silicon photomultiplier and/or a channel multiplier.
  • the cooperating pairs of diametrically opposed scintillation detectors may include a channel multiplier.
  • Various exemplary embodiments of the present general inventive concept may provide an event-by-event, real-time, high resolution, three-dimensional positron emission tomography encoder system which includes a plurality of cooperating pairs of diametrically opposed scintillation detectors adapted for receiving gamma rays from a positron event, electronic circuitry in electronic communication with each pair of said diametrically opposed scintillation detectors for determining the two-dimensional position of the positron event occurring between each pair of detectors, each event producing said gamma rays which travel along a line of response extending between cooperating pairs of detectors.
  • Further electronic circuitry may include a Constant Fraction Discriminator and a Digital Intervalometer utilized either as discrete chip sets or, in an exemplary embodiment, embedded on an application specific integrated circuit for measuring the difference between the arrival times of coincident gamma rays from a positron event detected by said diametrically opposed scintillation detectors along said line of response thereby enabling determination of the third dimension along said line of response of said positron event to produce a real-time, high resolution, three dimensional positron emission tomographic image without image reconstruction.
  • the digital intervalometer may utilize a time amplitude converter.
  • the pairs of diametrically opposed scintillation detectors may each include a silicon photomultiplier operating in an avalanche photodiode mode.
  • the pairs of diametrically opposed scintillation detectors may include a channel multiplier.
  • Various exemplary embodiments of the present general inventive concept may provide a process for capturing event-by-event, high resolution, three-dimensional positron emission tomography imagery, in real-time, which includes providing a selected number of cooperating pairs of diametrically opposed scintillation detectors adapted for receiving coincident gamma rays from a positron event, determining a two-dimensional position of the positron event producing said gamma ray photons, determining a line of response for each cooperating pair of detectors of said coincident gamma rays, and detection of and measuring, in real-time, the time of said coincident gamma rays being detected by a cooperating pair of diametrically opposed scintillation detectors along said line of response in order to determine the third dimension along said line of response of said positron event using a Constant Fraction Discriminator and a Digital Intervalometer utilized as discrete chip sets or, in an exemplary embodiment, embedded on an application
  • the digital intervalometer may utilize a time-to-amplitude converter.
  • the pairs of diametrically opposed scintillation detectors may include a silicon photomultiplier and/or a channel multiplier.
  • the pairs of diametrically opposed scintillation detectors may include a channel multiplier.

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Abstract

Système d'imagerie TEP à résolution temporelle pour produire des images tomographiques tridimensionnelles à haute résolution en temps réel événement par événement par émission de positrons sans réaliser de formation de sinogramme ni de reconstruction d'image. La troisième dimension est fournie par la mesure de ΔT entre les temps d'arrivée de rayons gamma d'un événement de positron qui est détecté par deux détecteurs coopérants. Afin de déterminer l'emplacement d'un événement de positrons le long des lignes de réponse, la mesure comprend un scintillateur rapide, un discriminateur de fraction constante et l'intervallomètre numérique. Le temps d'arrivée de chaque photon dans le processus d'annihilation est enregistré par rapport à une fréquence d'horloge avec une résolution en picosecondes. Cette approche nécessite beaucoup moins d'événements de positrons, ce qui nécessite moins de détecteurs, ce qui permet d'obtenir un système d'imagerie TTEP en temps réel événement gamma par événement gamma qui est plus efficace et plus économique à produire que les systèmes TEP classiques.
EP21889769.2A 2020-11-09 2021-06-07 Système de codeur de tomographie par émission de positrons à résolution temporelle Pending EP4241117A4 (fr)

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US17/093,095 US11054534B1 (en) 2020-04-24 2020-11-09 Time-resolved positron emission tomography encoder system for producing real-time, high resolution, three dimensional positron emission tomographic image without the necessity of performing image reconstruction
PCT/US2021/036153 WO2022098394A1 (fr) 2020-11-09 2021-06-07 Système de codeur de tomographie par émission de positrons à résolution temporelle

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