Time-resolved positron emission tomography encoder system

Abstract

A time-resolved PET imaging system for generating event-by-event, real-time, high-resolution, three-dimensional positron emission tomography images without sinogram tracing or image reconstruction. The third dimension is provided by measuring the ΔT between the arrival times of gamma rays from a positron event detected by two cooperating detectors. To determine the location of the positron event along the line of response, the measurement involves a fast scintillator, a constant ratio discriminator, and a digital intervalometer. The arrival time of each photon in the annihilation process is recorded relative to a clock frequency with picosecond resolution. This approach requires significantly fewer positron events, and therefore fewer detectors, resulting in a gamma event-by-gamma event, real-time TPET imaging system that is more efficient and more economical to produce than conventional PET systems.

Description

[Technical field]

[0001] This application claims the benefit of U.S. Non-provisional Patent Application No. 17 / 340,485, filed June 7, 2021, and U.S. Non-provisional Patent Application No. 17 / 093,095, filed November 9, 2020, each of which is incorporated by reference in its entirety herein.

[0002] The general concept of the present invention relates to a Time-Resolved Positron Emission Tomography ("TPET") system, and more particularly to a much simpler time-resolved PET imaging system that provides event-by-event, or real-time imaging, that produces high-resolution PET images with only a single view, requiring fewer detectors, no image reconstruction, and no additional / simultaneous computed tomography ("CT") for attenuation correction. [Background technology]

[0003] When a positron decays, it generates two gamma rays with 180 degrees opposite direction (see Figure 1A). Although there is a small error in angle due to the momentum of the positron, imaging it for Positron Emission Tomography (PET) is assumed to be negligible. In PET imaging, the line formed by the detection of the two gamma rays is called the "Line of Response" (LOR). In the field of positron emission tomography, it is known that state-of-the-art photomultiplier tubes and relatively slow scintillators have limited the timing resolution of scintillator-based optical detection systems for decades. It is also known that existing PET systems require all views in a plane to form a complete image of the object. These systems utilize software algorithms that attempt to find the intersection of multiple LORs to form an image. In Figure 1B, the LOR for a point source is shown. The circles represent detectors that can detect gamma rays from a point source. A typical state-of-the-art prior art ring PET tomograph would have hundreds of detector elements in a ring, see for example FIG. 1D. In the prior art, there would be more than 100,000 LORs in a single ring. To use this set of LORs to form an image, the LORs must be arranged in a regular format by a high-speed computer. The arrangement is called a sinogram, and the process is called sorting.

[0004] From FIG. 1C, it can be seen that if an image has two locations of the positron, the process of finding the intersection of the LORs to identify the exact location of the positron becomes more complicated. In fact, more than one LOR intersection is always required to identify a single source of positrons. Considering hundreds of locations of positron events in an image, it can be understood that millions of LOR intersections are required to form an approximate image. The process of interpreting millions of LORs is called "image reconstruction". This is well known in the prior art, see for example US Pat. No. 5,399,433, which discloses: "The LORs, defined by the coincident events after being sorted into parallel projections, represent the three-dimensional distribution of the positron-emitting radionuclide within the patient. To reconstruct "It is used." Paragraph 1 of Patent Document 1

[0005] Emphasis added. Henselaer further teaches that "the imaging system 812 may use the onset of gamma interactions to reconstruct the three-dimensional distribution of positron-emitting radionuclides within the patient and display a useful image to the physician or operator." Paragraph

[0051] of the '591 patent. The process of image reconstruction requires the use of powerful computers and typically requires several minutes to complete, which is time consuming from a diagnostic standpoint. Figure 2A shows a prior art PET imaging system as shown by the '591 patent.

[0005] As shown in Figure 2A, the two main computers typically used in state-of-the-art image reconstruction are represented by the box labeled Computer Readable Medium 806, where the sorting process occurs and where the memory of the sinograms is located, and Processor 810, the reconstruction processor, which is typically a powerful computer. The reconstruction process can be either a statistical process or a backprojection process, which can require several minutes with a fast computer to form an approximate image.

[0006] The advent of avalanche photodiodes began to replace photomultiplier tubes and brought major improvements in the time resolution of scintillator systems and in other parameters such as resistance to magnetic fields. Avalanche photodiode technology applied to silicon photomultipliers (SiPMs) made it possible to measure the time it takes for a gamma ray to travel along a particular line of response ("LOR"). In other words, the improved time resolution allows for precise measurement of the location of annihilation events along the LOR. This time measurement is the basis for the time-resolved positron emission tomography system disclosed herein, referred to herein as a TPET imaging system. TPET imaging systems should not be confused with existing time-of-flight systems, such as those shown in FIG. 2A and described in U.S. Pat. No. 5,393,363, currently commercially available. For example, U.S. Pat. No. 5,393,363 discloses the following: Time-of-flight (TOF) Positron Emission Tomography (PET) ("TOF-PET") is based on the measurement of the difference in detection times ΔT of two gamma photons resulting from a 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). Although less accurate than the spatial resolution of the scanner, 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. Thus, in TOF-PET, the "TOF" coordinate ΔT is stored together with s and φ.

[0007] Paragraph

[0010] of Patent Document 1.

[0008] It is known in the art that the use of TOF information in PET has the potential to enable significant improvements in image noise characteristics, which may in turn improve lesion detection. This concept of TOF has been proposed in the literature and is commercially available, but no designs have been proposed or developed that produce acceptably high temporal resolution, single view, and directly form images using temporal information. The present disclosure includes a system that optimizes the temporal resolution of silicon photomultipliers (SiPMs), including the use of very fast scintillators, and has the potential to provide comparable or better image resolution compared to conventional PET images.

[0009] A typical LOR is determined by a standard block detector that measures the gamma ray position, both X and Y coordinates in two-dimensional space. The third-dimensional coordinate Z in the XYZ coordinate system is then provided by a time measurement of the travel of the coincident gamma rays, more specifically, by the difference between the time when one gamma ray is detected and the time when the second gamma ray is detected. This time measurement can be done by an analog technique known in the art as a Time-to-Amplitude Converter ("TAC"). A TAC measurement is achieved by the first gamma ray switching a constant current onto a small capacitor and the second gamma ray switching the current off. As a result, the voltage on the capacitor is proportional to the time difference between the two signals. The problem with this technique is that many events occur within a short period of time, and to provide real-time, i.e., gamma event-by-event, analysis, the system must be able to measure all of these events and detect and separate them in real time such that the time difference falls within the desired field of view. State-of-the-art analog TACs simply do not have this capability.

[0010] Since gamma rays travel at the speed of light, i.e. 30 centimeters per nanosecond, multiple gamma rays travel in opposite directions for PET, 15 centimeters from the center corresponds to a time difference of 1 nanosecond. The reconstructed resolution of a conventional PET system is roughly 5-6 millimeters in three dimensions, so 6 millimeters along the LOR corresponds to 40 picoseconds. This is the required time resolution of the TPET imaging system so that the Z coordinate position resolution is equivalent to existing standard PET imaging systems.

[0011] Current PET scanners perform image reconstruction by measuring and recording the LORs and then statistically determining the intersections of all the LORs with each other. The statistical process requires a very large number of events, and therefore a large number of detectors, and is therefore very inefficient. This process requires many more events than the direct measurement of the TPET imaging system of the general concept of the present invention, and is therefore much less efficient than the TPET imaging system disclosed herein. This can be best understood for a simple point source of activity within the field of view, where a single TPET measurement can result in an image. It is known in the PET art that when the volume under consideration has a complex distribution of activity, the number of events required for a good image can increase to millions of events for conventional systems.

[0012] Time measurement systems currently known in the art use standard photomultiplier tubes or SiPM detectors such as the SiPM detectors supplied by Philips or Hamamatsu. Known measurement systems utilizing SiPM optical detectors and LSO scintillators utilize a leading edge discriminator and a simple frequency counter. This type of system is known to be incapable of producing high resolution performance in the time domain or in the third dimension of a PET system, event by event, i.e. in real time, for two additional reasons. The first problem is the leading edge timing used in the discriminator. This leading edge technique uses an electronic discriminator that changes state when the leading edge of the pulse exceeds a defined level. The state of the art time resolution is 250 picoseconds for conventional PET systems and is used to set the coincidence window for standard PET systems. This time resolution is primarily set by the LSO scintillator used. There does not appear to have been any attempt in the art to use this timing information to eliminate image reconstruction. However, even if this timing information was used in a conventional PET system to eliminate the reconstruction process, the image resolution would be roughly 37 millimeters. Even if TOF were used to eliminate image reconstruction, the result would be unacceptably low resolution for the diagnostic imaging industry. These problems are addressed in the TPET imaging system disclosed herein.

[0013] Furthermore, it is known in the art that the time walk associated with a leading edge discriminator is undesirable and that even if the amplitude of the optical signal does not vary appreciably, the residual time walk can adversely affect the time resolution. As is evident from FIG. 3, assuming an input rise time of 9,000 picoseconds, i.e. 9 nanoseconds, the time walk due to pulse height variations can be very significant. In this regard, it is known that in the timing area, the input current pulse height can vary V by 30 to 50%. Assuming that the input pulse has a linear rise time, the time walk associated with a leading edge discriminator is given by:

number

[0014] Using a 40% variation in the current pulse, a 9,000 picosecond rise time, and assuming a 100:1 ratio from VMAX to VT, this ratio is considered the maximum workable ratio by those working in the analog field. Using these parameters in the formula:

number

[0015] For the two detectors required in a traditional PET system, the time resolution in this example is 60 ps or 85 ps square root of two. The best resolution reported in the prior art is about 100 ps. From this example, it becomes clear that the time uncertainty due to the time walk is significant and must be significantly reduced or eliminated in any practical time-resolved PET imaging system.

[0016] Yet a second known time measurement system is the digital intervalometer described by the applicant hereof in U.S. Pat. No. 5,399,323, issued Nov. 17, 1970 to the applicant hereof, which is incorporated herein by reference. This time measurement consists of a digital clock that counts the ticks of an oscillator between a start and a stop signal. In addition, 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 using the digital TAC. The most significant digit of the clock is supplied by the digital TAC and is added successively to a word containing the oscillator ticks. If the desired measurement is between the 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 511 KeV gamma rays.

[0017] A time resolved PET imaging system that provides event-by-event, real-time, high resolution imaging that utilizes fewer detectors compared to traditional PET systems and does not require image reconstruction is not available in the known art. It is therefore a feature of the general concept of the present invention to provide a time resolved PET system that has a much simpler design, and therefore is consequently less costly to build, compared to traditional PET imaging systems, that provides real-time, event-by-event, high resolution imaging, and eliminates the need for CT scans for attenuation correction. A key feature of the general concept of the present invention is to provide a TPET imaging system that utilizes timing information to determine position along the LOR to provide measurements of third dimensional coordinates, resulting in near real-time, high resolution, three dimensional imaging in a TPET imaging system that has 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 general concept of the present invention can be used to image specific organs of the body without introducing artifacts, unlike current PET systems that require views all around the patient. This allows for very practical imaging of organ systems such as chest imaging, heart imaging, and prostate imaging, to name a few, that can be imaged using the TPET imaging system of the general concept of the present invention. Another very important feature of TPET is that only one view is required, which allows for accurate calculation of the attenuation of gamma rays emitted from the patient. Known prior art systems require CT images to correct for the attenuation of the emitted gamma rays. This results in a much more complicated PET system, and therefore a much more expensive system. Therefore, there is a need for a time-resolved PET system that will generate real-time, event-by-event, high-resolution, three-dimensional positron emission tomography images without the need for image reconstruction or the need to use CT images to correct for attenuation. [Prior art documents] [Patent documents]

[0018] [Patent Document 1] US Patent Application Publication No. 2013 / 0009063 [Patent Document 2] U.S. Pat. No. 3,541,448 Summary of the Invention

[0019] In an exemplary embodiment of the overall concept of the present invention, a TPET imaging system capable of providing event-by-event, i.e., real-time, high-resolution, three-dimensional images is provided without the need for image reconstruction or CT correction of gamma ray attenuation. The radiation detector provides the X and Y coordinates, while the third-dimensional coordinate, i.e., the Z coordinate, is provided by measuring the detection time difference ("ΔT") between gamma rays from a positron event, which are detected by two gamma ray detectors that are much faster than the scintillator detectors used in conventional PET systems. To provide this time measurement in real time and in a manner that requires detection of significantly fewer events than traditional PET, rather than using a leading edge timing discriminator and a traditional time encoder to determine the location of the gamma event along the LOR, in an exemplary embodiment, the TPET imaging system disclosed herein utilizes additional electronic circuitry, which together measure the arrival time of each photon in the annihilation process with picosecond resolution, reducing the time walk to an insignificant contribution to the overall time resolution. In an exemplary embodiment, reducing the time walk is accomplished through the use of a Constant Fraction Discriminator ("CFD") invented by the present applicant that operates in conjunction with a Digital Intervalometer to perform the time measurement function. While the Constant Fraction Discriminator and Digital Intervalometer may be utilized as separate circuits or separate chip sets, in an exemplary embodiment, the Constant Fraction Discriminator and Digital Intervalometer are embedded in an Application Specific Integrated Circuit ("ASIC"). Using this time measurement system, the arrival time of each photon in the annihilation process is recorded relative to a clock frequency having picosecond resolution. In an exemplary embodiment, this TPET technique requires significantly fewer gamma events and therefore fewer detectors, thereby resulting in a system that is more efficient and more economical to manufacture.By providing real-time, high-resolution, artifact-free images, the TPET imaging system of the present invention's general concept allows real-time image comparison of multiple organs, thereby giving the clinician / diagnostician a better understanding of the relationships of multiple organ systems. The system also allows for real-time observation of cancer lesions being treated by radiation or other means.

[0020] The above-described features of the general inventive concept will be more clearly understood from the following detailed description of the general inventive concept read in conjunction with the drawings. [Brief description of the drawings]

[0021] [Figure 1A] FIG. 1 is a simplified schematic showing the physics of a positron event in which two gamma rays are produced, traveling 180° opposite each other. [Figure 1B] FIG. 1 is a simplified schematic diagram showing the line of response ("LOR") from a single point source of positrons. [Figure 1C] FIG. 1 is a simplified schematic diagram illustrating the complications that arise in the process of locating the intersection of the LORs to identify the exact location of the positron when the image has two locations of the positron. [Figure 1D] 1 is a schematic diagram of a prior art PET imaging system having a detector ring and relying on image reconstruction. [Figure 2A] 1 is a schematic diagram of a conventional PET imaging system, as exemplified in US Pat. No. 6,399,323. [Figure 2B] FIG. 1 further illustrates, in a schematic diagram, a conventional PET imaging system. [Diagram 3] Figure 1 illustrates the time walk problem due to pulse height variation, which is common in many standard PET systems. [Figure 4] FIG. 2 illustrates the operational process of Applicant's digital intervalometer as utilized in an exemplary embodiment of the present general concept. [Diagram 5] 1 is a schematic diagram of a TPET imaging system according to the general concept of the present invention; [Figure 6A] FIG. 2 is a more detailed diagram of a TPET according to an exemplary embodiment of the general concept of the present invention. [Figure 6B] FIG. 2 is a more detailed diagram of a TPET according to an exemplary embodiment of the general concept of the present invention. [Figure 6C] FIG. 2 is a more detailed diagram of a TPET according to an exemplary embodiment of the general concept of the present invention. [Figure 7A] FIG. 2 is a schematic diagram illustrating the process of capturing a three-dimensional image and determining the X, Y, and Z coordinates of each positron event using a two-detector system according to an exemplary embodiment of the present general concept. [Figure 7B] FIG. 2 is a schematic diagram illustrating a process for determining vector coordinates of each positron event using a two-detector system according to an exemplary embodiment of the present general concept. [Figure 8] FIG. 1 is a schematic diagram of an exemplary embodiment of a simpler TPET imaging detector array system. DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

[0022] A time resolved positron emission tomography, TPET, system 10, capable of producing real-time, high resolution, three-dimensional images, is shown in Figures 5, 6A, 6B, 6C, 7A, 7B, and 8. In state-of-the-art PET images, it is understood that the normal LOR 20 is determined by standard block detectors 30A and 30B, which, as will be understood, are diametrically opposed to each other and measure the gamma ray position in both X and Y coordinates, 40 and 50 respectively, in two-dimensional space as shown diagrammatically in Figure 7A. Figure 5 shows a TPET imaging system of the general concept of the present invention. Figure 5 shows a TPET imaging system showing two parallel plates 30A and 30B of detectors required, which typically has ten times fewer detectors than the prior art.

[0023] In Figures 7A and 7B, a positron event 75, such as the positron event in Figure 7A, is shown diagrammatically occurring in an organ 70. With reference to Figures 7A and 7B, a single positron event at point B in Figure 7A can be represented as follows: Z LOR =Z 1 +Z 2 ΔT = (Z 1 -Z 2 )k Here, k = 6.66ps / mm (i.e. the reciprocal of the speed of light)

number

number

number

[0024] The time difference ΔT between the arrival times of the coincident gamma rays detected by the block detectors 30A and 30B determines the location of the positron event along the LOR 20. FIG. 7A shows the results of a basic measurement including 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 the arrival times of the two gamma rays at the opposing detectors. By the Z time measurement and the X and Y determination, the point in space where the event occurred can be determined and displayed. In other words, referring to FIG. 7A for example, the third dimension in the XYZ coordinate system, the Z coordinate 60, is provided by time measurement of the travel of the gamma rays and by calculating the difference between the time when one gamma ray is detected and the time when the second coincident gamma ray is detected.

[0025] While measuring gamma ray, i.e., positron event position, and the position itself, has been discussed in Cartesian terms of X, Y, and Z coordinates, in an exemplary embodiment, vector analysis, i.e., vector algebra, will be used to determine the location of a positron event, such as positron event 75 in FIG. 7B. In this regard, the processing circuitry will determine the position of positron event 75 when the detector elements of each detector plate 30A and 30B are measured and the time difference ΔT between the respective detector elements is measured. With reference to FIG. 7B, each vector will be represented by an X, Y, and Z coordinate system. 1 , Y 1 The element count is defined in terms of data points. In an exemplary embodiment, the exemplary array would consist of approximately 40 detectors in the horizontal or X direction and approximately 40 detectors in the vertical or Y direction. This results in an array with 1,600 detectors in each detector plate. The two detector plates are diametrically opposed to each other and, in an exemplary embodiment, are parallel to each other as shown in FIG. 8. With this element count, (1,600) 2There will be lines of response ("LORs"). Each LOR will be defined prior to use of the system with the location of each memory pixel along the LOR vector. With such an arrangement, a computer having a cubic memory will have the intersection of each LOR along with each memory pixel stored in its memory. When the time difference ΔT is measured and the LOR identified, then a location is identified for each positron event, such as positron event 75. The computer then uses vector analysis instead of Cartesian coordinates to form a three dimensional image. The three points of the hypothetical vector are defined or represented as follows: Vector: (X 25 ,Y 35 )(X 15 ,Y 20 ),ΔT=2.33ns

[0026] Thus, referring back to FIG. 7A, the vectors for positron events A, B, and C occurring in organ 70 can be expressed as follows: Point A:(X 21 ,Y 23 ),ΔT=250ps Point B:(X 22 ,Y 21 ),ΔT=350ps Point C:(X 18 ,Y 19 ),ΔT=500ps

[0027] These events are stored in the miniature memory and data acquisition module 120 for analysis and continuous display in the imaging system 110. In the imaging system 110, vector analysis and ΔT are calculated and displayed for each gamma event in real time as it is detected, without the need for either time-consuming image reconstruction or CT scans for attenuation correction. In an exemplary embodiment, the block detectors 30A and 30B can be silicon photomultipliers (SiPMs) or conventional photomultipliers.

[0028] In an exemplary embodiment, the TPET imaging system 10 of the present general concept uses SiPMs for the photodetectors and CeBr 3 (35,000 photons per 511 KeV) or LaBr 3 It utilizes a very high photon yield, fast scintillator such as a CeBr3 scintillator. With a CeBr3 scintillator, the stopping power for gamma rays is worse than a normal PET scintillator, but this loss in stopping power is offset by the high intrinsic efficiency of TPET. For this TPET design, the position resolution along LOR20 is calculated as follows:

[0029] The theoretical resolution of a single exponential scintillator decay is

number

[0030] The time resolution of SiPM is T(fwhm)=25ps

[0031] For one channel, T=27.5ps

[0032] Since there are two channels, T s =1.41×27.5=38.8

[0033] This resolution at time corresponds to a resolution of 5.8 mm along LOR20.

[0034] Referring to FIG. 6C, to provide this time measurement on a gamma event-by-gamma event basis, i.e., in real time, and in a manner that requires detection of significantly fewer events than traditional PET, an exemplary embodiment of the TPET imaging system 10 utilizes additional electronic circuitry from FIG. 6A that reduces the time walk to an insignificant contribution to the overall time resolution.

[0035] The electronic circuit of FIG. 6A for reducing the time walk described above functions to cause a trigger for the time measurement to occur at a fixed ratio of input pulses. This is accomplished by applying an input pulse to two input lines and delaying, inverting, and amplifying one of the input lines. The resulting negative, delayed, and amplified input is added back to the original input pulse. The result is a bipolar pulse with zero crossing times that occur at the same times as a fixed ratio of the pulses. While the amplitude of the pulse changes, the zero crossing times remain the same.

[0036] A time measurement electronic circuit 100 is provided, Fig. 6A, consisting of a digital clock that counts oscillator ticks between the start and stop signals. In this regard, the additional electronic circuit 100 measures, with picosecond resolution, the arrival time of each photon in the annihilation process.

[0037] The time measurement electronics 100 of FIG. 6A functions to count the ticks of an oscillator between the start and stop signals. The limitation of conventional, state-of-the-art systems is that the accuracy is plus or minus (±) one tick of the oscillator. To obtain greater accuracy, an analog time-to-pulse height converter is used to measure the time from the start pulse to the oscillator, and this time is added back to a conventional pulse counter. The same process is used to remove any inaccuracies in the stop signal.

[0038] In addition, the phase between the start and oscillator is measured by the Digital TAC, and the difference between the oscillator and stop signals is also measured using the Digital TAC. The most significant digit of the clock is provided by the Digital TAC and is continuously added to the word containing the oscillator ticks. If the desired measurement is between the start and oscillator, as is the case for TPET operation, then only one of the TAC interpolators is used to measure the time of arrival of one of the 511 KeV gamma rays.

[0039] In an exemplary embodiment, the electronics for reducing the time walk is provided by a constant ratio discriminator ("CFD") 90'. The use of the CFD 90' reduces the time walk to an insignificant contribution to the overall time resolution of the TPET imaging system 10. In an exemplary embodiment of the TPET imaging system 10 of the general concept of the present invention, this picosecond resolution time measurement electronics is defined by a digital intervalometer 100. With this time measurement system, the arrival time of each photon in the annihilation process is recorded relative to a clock frequency having picosecond resolution. With proper design and use of this time measurement, the time resolution becomes insignificant for the overall measurement. The digital intervalometer 100' utilizes the TAC to interpolate between the start signal and the clock. The TAC is then converted to a digital signal by standard analog-to-digital converter ("ADC") methods, and this digital signal is added to the digital clock word to form the time measurement. The stop signal is generated by an arbitrary fixed signal synchronized to the clock. This process is illustrated in FIG. 4. In this regard, it is known in the art that if no interpolation is used, the time resolution measured using only clock pulse counting results in a ±125 ps for a clock frequency of 8 GHz. Previous experiments and work in the art have attempted to compensate for various issues using state-of-the-art timing discriminators and time encoders, by performing the interpolation offline, and by attempting to measure the input rise time and extrapolating the phase to do so. This is not a practical approach in clinical diagnostic applications. In this regard, the clinical diagnostic environment requires a TPET imaging system 10 capable of generating high-resolution, three-dimensional images for every gamma event without image reconstruction.

[0040] This is the case for the TPET imaging system 10. The arrival time of each gamma ray event is measured, digitally stored, and then digitally compared to every other measured event. Any two measured events occurring within a few picoseconds of each other will be defined as a positron annihilation event and used to form a three-dimensional image. In an exemplary embodiment, this approach requires significantly fewer gamma events, thus requiring fewer detectors, and produces a high-resolution, three-dimensional image per gamma event. This results in a TPET imaging system 10 that is more efficient and more economical to manufacture than conventional PET imaging systems.

[0041] In summary, in an exemplary embodiment, the TPET imaging system 10 of the present invention's overall concept includes the use of electronic circuitry to measure arrival time differences to determine the location of positron events along the LOR 20 in a manner that allows for event-by-event, high resolution images and reduces time walk. In an exemplary embodiment of the TPET imaging system, this additional electronic circuitry is defined by both a fixed ratio discriminator 90 and a digital intervalometer 100. While the fixed ratio discriminator 90 and the digital intervalometer 100 may be used as separate chip sets or separate integrated circuits, in an exemplary embodiment, the fixed ratio discriminator 90 and the digital intervalometer 100 are combined within an application specific integrated circuit ("ASIC"). Although an exemplary embodiment has been described that utilizes a fixed ratio discriminator and a digital intervalometer to perform these functions, it will be understood that other electronic circuitry adapted to perform these functions may also be utilized, particularly as computing efficiency and power increases.

[0042] A third dimension can be introduced by measuring the arrival time of each photon in the annihilation process and calculating the difference in arrival times at opposing detector blocks, resulting in the TPET imaging system 10. In an exemplary embodiment, the TPET imaging system of the present invention's general concept will have the ability to meet and even exceed the image resolution of existing PET systems, and contribute three additional very important attributes to the system: The TPET imaging system 10 of the present invention's general concept will be relatively simple and therefore very easy and inexpensive to manufacture compared to full-ring state-of-the-art PET tomographs.

[0043] As seen in FIG. 8, the disclosed TPET system utilizes significantly fewer detector blocks compared to state-of-the-art PET imaging rings (80 pixels compared to 18,000 pixels), so the TPET imaging system 10 of the present invention's general concept can provide artifact-free images without covering the entire body as is required for conventional PET imaging systems. This feature makes the TPET imaging system 10 of the present invention's general concept viable for individual organ imaging, such as chest, heart, prostate, and lung imaging for coronavirus lesions. This feature allows for designs with a very small number of detectors. A third major advantage is that the images are formed as each gamma event is detected, i.e., event-by-event, and are variously referred to herein as "real-time", without the inherent delays required in image reconstruction or attenuation correction. This feature gives the clinician / diagnostician real-time, event-by-event image comparisons of multiple organs, thereby providing a better understanding of the relationship of multiple organ systems, for example, the brain to the heart. The real-time feature also provides the clinician with a view of the cancer site as it is being treated with radiation or other means, such as proton therapy. These features make the TPET imaging system 10 of the present invention's general concept potentially replaces state-of-the-art PET systems and represents a major advancement for medical imaging in general. Furthermore, as will be appreciated by those skilled in the art, the avalanche photodiodes not only improve the time resolution of the scintillator system, but the resistance to magnetic fields allows the TPET imaging system 10 of the present invention's general concept to be used in combined TPET and magnetic resonance imaging ("MRI") systems to simultaneously acquire MRI and TPET images in a single device, and allows the TPET imaging system to operate in relatively high radiation fields. A fourth feature is that TPET uses only one view, compared to traditional PET's 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 the need for a CT scan for attenuation correction.This further reduces the cost and complexity of the TPET imaging system 10.

[0044] Various exemplary embodiments of the general concept of the present invention can result in an event-by-event, high resolution, three-dimensional positron emission tomography encoder system including a plurality of cooperating pairs of diametrically opposed scintillation detectors adapted to receive gamma rays from positron events, electronics in electronic communication with each pair of diametrically opposed scintillation detectors to determine a two-dimensional location of a positron event occurring between each pair of detectors, each event producing gamma rays traveling along a line of response extending between the cooperating pair of detectors, and additional electronics for measuring the difference in arrival times of coincident gamma rays from a positron event detected by the diametrically opposed scintillation detectors along the line of response, thereby enabling a determination of a third dimensional location of the positron event along the line of response to generate an event-by-event, real-time, high resolution, three-dimensional positron emission tomography image without the need for image reconstruction or CT scans for attenuation correction. The additional electronics can include a constant ratio discriminator and a digital intervalometer. In embodiments where a constant ratio discriminator and a digital intervalometer are utilized, these may be utilized as separate chip sets or, in an exemplary embodiment, may be embedded in an application specific integrated circuit. The digital intervalometer may utilize a time to amplitude converter. Each pair of diametrically opposed scintillation detectors may include a silicon photomultiplier tube operating in an avalanche photodiode mode. Each pair of diametrically opposed scintillation detectors may include a channel multiplier tube.

[0045] Various exemplary embodiments of the present invention's general concept may provide a process for capturing event-by-event, high-resolution, three-dimensional positron emission tomography images in real time, the process including providing a selected number of cooperating pairs of diametrically opposed scintillation detectors adapted to receive coincident gamma rays from a positron event, determining a two-dimensional location of the positron event generating the gamma ray photon, determining a line of response for each cooperating pair of detectors of the coincident gamma rays, and detecting and measuring in real time the time of the coincident gamma rays detected by the cooperating pair of diametrically opposed scintillation detectors along the line of response to determine a third dimension along the line of response of the positron event, thereby generating a real-time, high-resolution, three-dimensional positron emission tomography image. The process for capturing real-time, high-resolution, three-dimensional positron emission tomography images may include a constant ratio discriminator and a digital intervalometer, which may be utilized as a separate chip set or, in an exemplary embodiment, embedded in an application specific integrated circuit. The digital intervalometer may utilize a time-to-amplitude converter. Each of the cooperating pairs of diametrically opposed scintillation detectors may include a silicon photomultiplier and / or a channel multiplier. Each of the cooperating pairs of diametrically opposed scintillation detectors may include a channel multiplier.

[0046] Various exemplary embodiments of the present invention's general concept may result in an event-by-event, real-time, high-resolution, three-dimensional positron emission tomography encoder system including a plurality of cooperating pairs of diametrically opposed scintillation detectors adapted to receive gamma rays from a positron event, and electronic circuitry in electronic communication with each pair of said diametrically opposed scintillation detectors to determine a two-dimensional position of a positron event occurring between each pair of detectors, each event resulting in said gamma rays traveling along a line of response extending between the cooperating pair of detectors. Additional electronic circuitry may include a constant ratio discriminator and a digital intervalometer, utilized as a separate chip set or in an exemplary embodiment embedded in an application specific integrated circuit, that measures the difference in arrival times of coincident gamma rays from a positron event detected by the diametrically opposed scintillation detectors along the line of response, thereby enabling the determination of a third dimensional position of the positron event along the line of response to generate a real-time, high-resolution, three-dimensional positron emission tomography image without image reconstruction. The digital intervalometer may utilize a time-to-amplitude converter. Each of the pair of diametrically opposed scintillation detectors may include a silicon photomultiplier tube operating in an avalanche photodiode mode. Each of the pair of diametrically opposed scintillation detectors may include a channel multiplier tube.

[0047] Various exemplary embodiments of the present invention's general concept may provide a process for capturing event-by-event, high-resolution, three-dimensional positron emission tomography images in real time, the process including providing a selected number of cooperating pairs of diametrically opposed scintillation detectors adapted to receive coincident gamma rays from a positron event, determining a two-dimensional location of the positron event generating the gamma ray photon, determining a line of response for each cooperating pair of detectors of the coincident gamma rays, and detecting and measuring in real time the time of the coincident gamma rays detected by the cooperating pair of diametrically opposed scintillation detectors along the line of response to determine a third dimension along the line of response of the positron event using a constant ratio discriminator and a digital intervalometer, utilized as a separate chip set or, in an exemplary embodiment, embedded in an application specific integrated circuit, thereby generating a real-time, high-resolution, three-dimensional positron emission tomography image. The digital intervalometer may utilize a time-to-amplitude converter. Each of the pair of diametrically opposed scintillation detectors may include a silicon photomultiplier and / or a channel multiplier.The pair of diametrically opposed scintillation detectors may include a channel multiplier.

[0048] While the general concept of the invention has been illustrated by the description of several embodiments and exemplary embodiments have been described in detail, it is not the intention of the applicant to restrict or in any way limit the scope of the appended claims to such details. Additional modifications will be readily apparent to those skilled in the art. Thus, the general concept of the invention in its broader aspects is not limited to the specific details, representative apparatus and methods, and illustrative examples shown and described. Thus, departures may be made from such details without departing from the spirit or scope of the applicant's general patentable concept.

Claims

1. An encoder system for generating three-dimensional positron emission tomography images in real time, on an event-by-event basis, without image reconstruction and without requiring a CT scan to correct for positron emission tomography signal attenuation, a plurality of cooperating pairs of opposed gamma ray detectors adapted to detect positron events; an electronic circuit in electronic communication with each pair of opposed gamma ray detectors to determine a two-dimensional position of the positron events occurring between each pair of opposed gamma ray detectors, each event producing a gamma ray that travels along a line of response extending between a cooperating pair of detectors; additional electronic circuitry for reducing the time walk associated with a leading edge discriminator and for measuring, with picosecond resolution, the difference in arrival times of coincident gamma rays from the positron events detected by the opposing gamma ray detectors along the line of response, thereby enabling a determination of a third dimensional position of the positron events along the line of response, thereby generating an event-by-event, real-time, high resolution, three-dimensional positron emission tomography image without image reconstruction; An encoder system comprising:

2. 2. The per-event encoder system of claim 1, wherein the additional electronic circuitry for substantially reducing a time walk associated with a leading edge discriminator and for measuring differences in arrival times of coincident gamma rays includes a constant ratio discriminator and a digital intervalometer.

3. 3. The per-event encoder system of claim 2, wherein said fixed ratio discriminator and said digital intervalometer are embedded in an application specific integrated circuit.

4. 3. The per-event encoder system of claim 2, wherein the digital intervalometer utilizes a time-to-amplitude converter.

5. 2. The per-event encoder system of claim 1, wherein said pair of opposed gamma ray detectors each comprises a silicon photomultiplier tube operated in an avalanche photodiode mode.

6. 6. The per-event encoder system of claim 5, wherein said pair of opposed gamma ray detectors are arranged in opposed parallel plate arrays.

7. 2. The per-event encoder system of claim 1, wherein said pair of opposed gamma ray detectors are comprised of channel multipliers.

8. A process for generating three-dimensional positron emission tomography images in real time, on an event-by-event basis, without image reconstruction and without CT scans to correct for positron emission tomography signal attenuation, providing a selected plurality of cooperating pairs of opposed gamma ray detectors adapted to receive coincident gamma rays from a positron event; determining a two-dimensional position of the positron event that generates a gamma ray photon; determining a line of response for each cooperating pair of opposed gamma ray detectors of the coincident gamma rays; reducing a time walk associated with a leading edge discriminator; detecting and measuring, with picosecond resolution, the time difference of the coincident gamma rays being detected by cooperating pairs of gamma ray detectors opposed along the line of response, for each gamma event, to determine a third dimensional position of the positron event along the line of response, thereby producing an event-by-event, high resolution, three dimensional positron emission tomography image without image reconstruction; The process includes:

9. 9. The process of claim 8, wherein the step of detecting and measuring the time of the coincident gamma rays being detected on a gamma event by gamma event basis includes the use of a constant ratio discriminator and a digital intervalometer.

10. 10. The process of claim 9, wherein the fixed ratio discriminator and the digital intervalometer are embedded in an application specific integrated circuit.

11. The process of claim 9 , wherein the digital intervalometer utilizes a time-to-amplitude converter.

12. 9. The process of claim 8, wherein the cooperating pair of opposed gamma ray detectors is selected from the group consisting of silicon photomultipliers and channel multipliers.

13. The process of claim 8 , wherein the pair of opposed gamma ray detectors are comprised of channel multipliers.

14. 14. The process of claim 13, wherein the pair of opposed gamma ray detectors are arranged in opposed parallel plate arrays.

15. An encoder system for generating three-dimensional positron emission tomography images in real time, on an event-by-event basis, without image reconstruction and without requiring a CT scan to correct for positron emission tomography signal attenuation, a plurality of cooperating pairs of opposed gamma ray detectors adapted to detect positron events; an electronic circuit in electronic communication with each pair of opposed gamma ray detectors to determine a two-dimensional position of the positron events occurring between each pair of opposed gamma ray detectors, each event producing a gamma ray that travels along a line of response extending between a cooperating pair of detectors; additional electronic circuitry for reducing the time walk associated with a leading edge discriminator and for measuring with picosecond resolution the difference in arrival times of coincident gamma rays from the positron events detected by the opposing gamma ray detectors along the line of response, thereby enabling a determination of a third dimensional position of the positron events along the line of response, thereby generating an event-by-event, high resolution, three dimensional positron emission tomography image without image reconstruction, and additional electronic circuitry including a time-to-amplitude converter. An encoder system comprising:

16. 16. The real-time, high-resolution, three-dimensional positron emission tomography encoder system of claim 15, wherein the additional electronic circuitry is embedded in an application specific integrated circuit.

17. 16. The real-time, high-resolution, three-dimensional positron emission tomography encoder system of claim 15, wherein the pair of opposed gamma ray detectors each comprises a silicon photomultiplier tube operated in an avalanche photodiode mode.

18. 16. The real-time, high-resolution, three-dimensional positron emission tomography encoder system of claim 15, wherein the pair of opposed gamma ray detectors are comprised of channel multipliers.

19. 16. The real-time, high-resolution, three-dimensional positron emission tomography encoder system of claim 15, wherein the pair of opposed gamma ray detectors are arranged in an opposed parallel plate array.

20. A real-time, high-resolution, three-dimensional positron emission tomography encoder system for generating three-dimensional positron emission tomography images in real time, on an event-by-event basis, without image reconstruction and without requiring a CT scan to correct for positron emission tomography signal attenuation; a plurality of cooperating pairs of opposed gamma ray detectors adapted to detect positron events; an electronic circuit in electronic communication with each pair of opposed gamma ray detectors to determine a two-dimensional position of the positron events occurring between each pair of opposed gamma ray detectors, each event producing a gamma ray that travels along a line of response extending between a cooperating pair of detectors; additional electronic circuitry, including a constant ratio discriminator and a digital intervalometer, for measuring the difference in arrival times of coincident gamma rays from the positron events detected by the opposing gamma ray detectors along the line of response, thereby enabling a determination of a third dimensional position of the positron events along the line of response, thereby generating a real-time, high resolution, three dimensional positron emission tomography image without image reconstruction; 1. A real-time, high-resolution, three-dimensional positron emission tomography encoder system comprising:

21. 21. The real-time, high-resolution, three-dimensional positron emission tomography encoder system of claim 20, wherein the constant ratio discriminator and the digital intervalometer are embedded in an application specific integrated circuit.

22. 21. The real-time, high-resolution, three-dimensional positron emission tomography encoder system of claim 20, wherein the pair of opposed gamma ray detectors each comprises a silicon photomultiplier tube operated in an avalanche photodiode mode.

23. 21. The real-time, high-resolution, three-dimensional positron emission tomography encoder system of claim 20, wherein the pair of opposed gamma ray detectors are comprised of channel multipliers.

24. 21. The real-time, high-resolution, three-dimensional positron emission tomography encoder system of claim 20, wherein the pair of opposed gamma ray detectors are arranged in an opposed parallel plate array.

25. 1. A real-time, high-resolution, three-dimensional positron emission tomography encoder system, comprising: a plurality of cooperating pairs of opposed gamma ray detectors adapted to receive gamma rays from the positron events; an electronic circuit in electronic communication with each pair of opposed gamma ray detectors to determine a two-dimensional position of the positron events occurring between each pair of opposed gamma ray detectors, each event causing the gamma ray to travel along a line of response extending between a cooperating pair of detectors; additional electronic circuitry including a constant ratio discriminator and a digital intervalometer for measuring the difference in arrival times of coincident gamma rays at each of the opposing gamma ray detectors from a positron event detected by the opposing gamma ray detectors along the line of response, the digital intervalometer utilizing a time to amplitude converter whereby a position of the positron event along the line of response is determined, thereby enabling a determination of a third dimensional position of the positron event along the line of response to generate a real time, high resolution, three dimensional positron emission tomography image without image reconstruction; 1. A real-time, high-resolution, three-dimensional positron emission tomography encoder system comprising:

26. 26. The real-time, high-resolution, three-dimensional positron emission tomography encoder system of claim 25, wherein the pair of opposed gamma ray detectors each comprises a silicon photomultiplier tube operated in an avalanche photodiode mode.

27. 26. The real-time, high-resolution, three-dimensional positron emission tomography encoder system of claim 25, wherein the pair of opposed gamma ray detectors are comprised of channel multipliers.

28. 26. The real-time, high-resolution, three-dimensional positron emission tomography encoder system of claim 25, wherein the constant ratio discriminator and the digital intervalometer are embedded in an application specific integrated circuit.

29. 26. The real-time, high-resolution, three-dimensional positron emission tomography encoder system of claim 25, wherein the pair of opposed gamma ray detectors are arranged in an opposed parallel plate array.

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