System and method for time-of-flight positron emission tomography

Abstract

Systems and methods for a time of flight (TOF) positron emission tomography (PET) system are provided herein. In one example, an imaging system (10) includes one or more detector blocks (300), each detector block including an array of silicon photomultiplier (SiPM) devices (306) coupled to a scintillation crystal array (302) in a one-to-one coupling arrangement, where each SiPM device in the array of SiPM devices (306) transmits signals to an independent front-end readout circuit (800) of one or more analog-specific integrated circuits (ASICs). The front-end readout circuit (800) is configured to detect individual scintillation photons and suppress SiPM dark counts.

Description

Technical Field

[0001] Embodiments of the subject matter disclosed herein relate to positron emission tomography (PET) systems, and more particularly to time-of-flight PET bulk detectors. Background Art

[0002] Positron emission tomography (PET) scanners generate images representing the distribution of positron-emitting nuclei in the patient's body. When a positron interacts with an electron by annihilation, the entire mass of the positron-electron pair is converted into two photons (also called events). Photons are emitted in opposite directions along a line of response (LOR). The annihilation photons are detected by detectors placed on both sides of the LOR in a configuration such as a detector ring. In time-of-flight (TOF) PET, in addition to measuring the interaction position in the detector, the actual time difference between the detection of the photons is also measured with high precision to more accurately identify the position of the annihilation event along the LOR. The detector converts the incident photon into a usable electrical signal that can be used for image formation. Modern PET systems include scintillators and photodetectors, such as silicon photomultiplier (SiPM) devices. Summary of the invention

[0003] In one example, an imaging system is provided herein that includes a detector block that includes an array of silicon photomultiplier tube (SiPM) devices coupled to an array of scintillator crystals in a one-to-one coupling arrangement, wherein each scintillator crystal is coupled to a separate SiPM device in the array of SiPM devices. Signals from the SiPM devices are transmitted to independent front-end readout circuits of one or more analog application-specific integrated circuits (ASICs).

[0004] It should be understood that the above brief description is provided to introduce in a simplified form selected concepts that are further described in the detailed description. It is not meant to identify key or essential features of the claimed subject matter, the scope of which is solely defined by the claims that follow the detailed description. Furthermore, the claimed subject matter is not limited to implementations that solve any disadvantages noted above or in any part of this disclosure. BRIEF DESCRIPTION OF THE DRAWINGS

[0005] The present invention will be better understood by reading the following description of non-limiting embodiments with reference to the accompanying drawings, in which:

[0006] Figure 1 is a pictorial view of an exemplary multi-modality imaging system according to an embodiment of the present invention;

[0007] Figure 2 is a schematic block diagram of an exemplary imaging system having a detector according to an embodiment of the present invention;

[0008] Figure 3 is a schematic diagram of an exemplary detector block according to an embodiment of the present invention;

[0009] Figure 4 is a schematic diagram of a portion of an exemplary detector block having a split reflector according to an embodiment of the present invention;

[0010] Figure 5A is a schematic diagram of a portion of an exemplary detector block having a light-sharing reflector and a separate reflector according to an embodiment of the present invention;

[0011] Figure 5B According to an embodiment of the present invention Figure 5A A schematic diagram of a top view of a portion of an exemplary detector block;

[0012] Figure 6 A schematic depiction depiction of an exemplary light-sharing reflector is shown;

[0013] Figure 7 According to an embodiment of the present invention Figure 4 A block diagram of a front-end readout circuit of an exemplary detector block;

[0014] Figure 8 is a high-level block diagram of an exemplary readout circuit according to an embodiment of the present invention; and

[0015] Fig. 9 is a diagram showing an embodiment according to the present invention Figure 8 A flow chart of a method of operation of a readout circuit. DETAILED DESCRIPTION

[0016] The following description relates to various embodiments of medical imaging systems. In particular, systems and methods are provided for silicon photomultiplier (SiPM) based block detectors and readout electronics thereof for bismuth germanate oxide (BGO) time-of-flight (TOF) positron emission tomography (PET) systems. Figure 1 to Figure 2 An example of a PET imaging system according to the present technology is provided in , which includes a detector that can be used to detect energy events. Figure 3 An exemplary detector block including a scintillation crystal array and a SiPM array is shown in FIG. Figure 4 A portion of a detector block with a one-to-one coupling is shown in , illustrating the split reflector. Figure 5A A portion of a detector block with a one-to-one coupling is shown in , illustrating the light sharing reflectors. Figure 5A The detector block in Figure 5B Shown in top view. Figure 6 An example of a light-sharing reflector is illustrated in . Figure 7 The front-end readout circuit is shown in Figure 4 A block diagram of the detector block, and Figure 8 A detailed high-level schematic of its readout circuitry is shown in . Fig. 9 The flowchart in Figure 8 The operation method of the readout circuit shown in .

[0017] In PET imaging, a chemical compound (e.g., a radioactive tracer) can be injected into a subject before a PET scan is performed, wherein one or more atoms of the chemical compound have been replaced with a positron-emitting radioactive isotope. For example, the chemical compound can be glucose, or a similar compound, such as fluorodeoxyglucose (FDG). The subject can be asked to wait for a period of time while the radioactive tracer is absorbed into tissues within a region of interest in the subject's body and accumulates therein. For example, the region of interest can be the subject's brain. Due to increased metabolic activity in malignant cells, the radioactive tracer may accumulate in higher concentrations around the malignant cells, which results in increased consumption of glucose (or similar compounds). The absorption time can depend on various factors and can be as short as a few minutes or as long as 90 minutes or longer. After a predetermined waiting period of time, the subject can be placed in a PET scanner.

[0018] Radioactive tracers contain positron-emitting radionuclides that decay over time. When the radionuclide decays, a positron is emitted, which can collide with an electron after traveling a short distance (e.g., less than 1 mm). When the positron collides with the electron, the positron and the electron annihilate, producing two 511keV photons traveling in opposite directions, called gamma rays. A PET scanner may include an array of detectors positioned around the subject that can detect gamma rays. The PET scanner may also include a coincidence circuit system (e.g., a readout circuit) that measures the timing of pulses generated when the gamma rays are aligned at the detector. If two gamma rays are aligned by different detectors within a certain amount of time, it can be inferred that the two gamma rays are generated by the same annihilated positron (e.g., a coincidence event). When a coincidence event occurs, a line of response (LOR) connecting the first detector and the second detector can be determined. Based on the LOR, the location of the annihilation can be determined in a process known as electronic collimation.

[0019] TOF-PET can provide statistical noise reduction compared to conventional PET, but TOF-PET requires high timing resolution. The timing resolution can be determined in part by intrinsic scintillation properties, such as decay time, rise time, and light yield of the scintillator material. As referred to herein, the decay time can be the time it takes for the intensity of a light pulse to return to its maximum value. At later times, rise time may be the rate at which the light output of the scintillator increases once the gamma rays have been stopped in the crystal, and light yield may be the number of photons produced per unit energy deposited by the particles slowed in the scintillating medium.

[0020] Other factors affecting timing resolution may include optical transfer time spread, interaction depth (where the gamma rays of an annihilation event are detected in the scintillator at different distances from the photosensor), optical transfer efficiency, photodetection efficiency, single photon temporal resolution (SPTR), crosstalk, dark counts (e.g., the response of a photon detector in the absence of light), and noise, bandwidth, and transient response of the readout electronics.

[0021] Annihilation gamma rays travel at a speed of about 30 cm / nanosecond. The time to identify coincident pairs is called the coincidence timing window. A smaller coincidence timing window reduces the number of random coincidences collected by the scanner. The arrival time measurement depends on the timing accuracy of the detector. The timing resolution (e.g., timing accuracy) is defined by the full width at half maximum (FWHM) of the arrival time difference from a point source of annihilation gamma rays. FWHM can be the width of the curve of arrival times measured at a height that is half the value of the bell curve at its peak. Timing resolution can be applied in various aspects, including reducing accidents (overlapping detection of two unrelated gamma rays) and timing signals for positioning purposes.

[0022] Single photon time resolution (SPTR) explains the accuracy with which the arrival time of a single scintillation photon at a photosensor is determined. In certain applications, particularly for low photon count applications such as those utilizing Cherenkov detectors, SPTR can significantly impact timing resolution. As a result of gamma-ray interactions, BGO scintillator crystals emit Cherenkov light. Cherenkov light is produced when charged particles travel through matter faster than the speed of light. Photoelectrons can emit Cherenkov light of about 1ps to 20ps. When high-energy gamma rays (e.g., 511keV photons from positron annihilation) interact with BGO (the high-Z scintillator material of the detector), photoelectrons with 420keV can be generated in about 40% of the first interactions. These fast electrons can emit about 20 Cherenkov photons in a short period of time (e.g., 20ps). In comparison, each interaction produces a relatively small number of Cherenkov photons, although Cherenkov photons can significantly increase the average number of photons detected in the first few hundred ps of an event. Due to the small number of Cherenkov photons, 3 to 5 photons are detected on average, so the readout electronics needs to be triggered at a level below the single photoelectron (SPE) response and also estimate the rise time in the first 100ps to 500ps of the pulse (e.g., 1ns to 5ns).

[0023] Lutetium-based scintillators are commonly used in TOF PET systems due to the required high timing resolution and scintillation properties, while BGO with its longer decay time was previously not feasible for TOF under current timing resolution parameters. However, as mentioned above, BGO scintillator crystals do emit Cherenkov light, thus allowing for improved timing resolution in TOF PET applications where Cherenkov light can be detected. A PET detector block for a BGO-based PET system with TOF capability is presented herein.

[0024] In one example, a PET detector block can include an array of SiPM devices and an array of scintillation crystals with a one-to-one coupling between the scintillation crystals and the SiPM devices, wherein each SiPM device has an independent readout circuit by which the signal from the SiPM device is processed. All crystals in the PET detector block can be optically isolated from each other and optically coupled to individual SiPM devices without any light sharing. In this way, light emitted by a crystal is sensed / detected only by the corresponding SiPM device. In one instance, the scintillation crystal array can be segmented into multiple pairs of adjacent crystals that share light via light-sharing reflectors to provide interaction depth estimates.

[0025] Each SiPM device has an independent front-end readout circuit in an analog application-specific integrated circuit (ASIC). The front-end readout circuit of each SiPM device provides separate measurements of the timestamp and signal amplitude or charge. The signals used for timing and energy measurements are shaped by filters. A timing pick-off method implements multiple comparators for each signal to provide additional information about the photons detected on the initial part of the scintillation pulse and the corresponding rise time of each pulse, where multiple comparators are configured to have a minimum timing pick-off threshold of 1 / 3 SPE, allowing Cherenkov photons to be detected. In this way, the timing resolution in event identification can be improved via the detection of Cherenkov photons. Improving the timing resolution for BGO-based PET scanners can in turn lead to improved image quality and sensitivity.

[0026] As used herein, the phrase "reconstructing an image" is not intended to exclude embodiments of the invention in which data representing an image is generated rather than a visible image. Thus, as used herein, the term "image" broadly refers to both a visible image and data representing a visible image. However, many embodiments generate (or are configured to generate) at least one visible image.

[0027] Various embodiments of the present invention provide Figure 1 and Figure 2 The multimodal imaging system 10 is shown. The multimodal imaging system 10 can be any type of imaging system, for example, different types of medical imaging systems, such as positron emission tomography (PET), single photon emission computed tomography (SPECT), computed tomography (CT), ultrasound systems, magnetic resonance imaging (MRI), or any other system capable of generating tomographic images. The various embodiments are not limited to multimodal medical imaging systems, but can be used for single-modal medical imaging systems, such as, for example, a stand-alone PET imaging system. In addition, the various embodiments are not limited to medical imaging systems for imaging human subjects, but can include veterinary systems or non-medical systems for imaging non-human objects.

[0028] refer to Figure 1, the multimodal imaging system 10 includes a first modality unit 11 and a second modality unit 12. The two modality units enable the multimodal imaging system 10 to scan an object or patient in a second modality using the second modality unit 12. The multimodal imaging system 10 allows multiple scans in different modalities in order to improve the diagnostic capabilities of a single modality system. In one embodiment, the multimodal imaging system 10 is a computed tomography / positron emission tomography (CT / PET) imaging system 10, for example, the first modality 11 is a CT imaging system 11, and the second modality 12 is a PET imaging system 12. In some examples, the PET imaging system 12 can be a TOF PET system. The CT / PET system 10 is shown as including a gantry 13 representing a CT imaging system and a gantry 14 associated with the PET imaging system. As described above, modalities other than CT and PET can be employed with the multimodal imaging system 10.

[0029] The gantry 13 includes an X-ray source 15 that projects a beam of X-rays toward a detector array 18 on the opposite side of the gantry 13. The detector array 18 is formed of a plurality of detector blocks (not shown) including a plurality of detector elements that together sense the projected X-rays passing through a medical patient 22. Each detector element generates an electrical signal representative of the intensity of the incident X-ray beam and thereby allows for an estimation of the attenuation of the beam as it passes through the patient 22. During scanning to acquire X-ray projection data, the gantry 13 and components mounted thereon rotate about a center of rotation.

[0030] Figure 2 According to an embodiment of the present invention Figure 1 Schematic block diagram of a PET imaging system 12 is shown. The PET imaging system 12 includes a detector ring assembly 40 that includes a plurality of detector blocks. The PET imaging system 12 also includes a controller or processor 44 to control normalization, image reconstruction processes and perform calibration. The controller 44 is coupled to an operator workstation 46. The controller 44 includes a data acquisition processor 48 and an image reconstruction processor 50 interconnected via a communication link 52. The PET imaging system 12 acquires scan data and transmits the data to the data acquisition processor 48. The scanning operation is controlled from the operator workstation 46. The data acquired by the data acquisition processor 48 is reconstructed using the image reconstruction processor 50.

[0031] The detector ring assembly 40 includes a central opening through which an object or patient such as the patient 22 may be moved using, for example, a motorized table 24 (eg, Figure 1The motorized table 24 is aligned with the central axis of the detector ring assembly 40. The motorized table 24 moves the patient 22 into the central opening of the detector ring assembly 40 in response to one or more commands received from the operator workstation 46. A PET scanner controller 54 (also referred to as a PET gantry controller) is disposed (e.g., mounted) within the PET system 12. The PET scanner controller 54 responds to commands received from the operator workstation 46 via the communication link 52. Thus, the scanning operation is controlled from the operator workstation 46 via the PET scanner controller 54.

[0032] During operation, when a photon (e.g., a gamma ray) collides with a scintillation crystal 62 of a detector block 63 on the detector ring 40, a scintillation event is generated on the crystal. When a scintillation event occurs, each photomultiplier tube or photosensor (e.g., a SiPM device) generates an analog signal that is transmitted on a communication line 64. A set of acquisition circuits 66 (e.g., readout circuits) are provided to receive these analog signals. The acquisition circuits 66 generate digital signals indicating the three-dimensional (3D) location, the total energy of the event, and the time or instant when the scintillation event occurred. These digital signals are transmitted via a communication link (e.g., a cable) to an event locator circuit 68 in the data acquisition processor 48.

[0033] The data acquisition processor 48 includes an event locator circuit 68, an acquisition CPU 70, and a coincidence detector 72. The data acquisition processor 48 periodically samples the signals generated by the acquisition circuit 66. The acquisition CPU 70 controls the communication on the backplane bus 74 and the communication link 52. The event locator circuit 68 processes information about each valid event and provides a set of digital numbers or values ​​indicating the detected event. For example, the information indicates when the event occurred and the location of the scintillation crystal 62 or detector block 63 that detected the event. Event data packets are transmitted to the coincidence detector 72 via the backplane bus 74. The coincidence detector 72 receives the event data packets from the event locator circuit 68 and determines whether any two of the detected events coincide. Coincidence is determined by multiple factors. First, the time stamps in each event data packet must be within a predetermined time period (e.g., 5 nanoseconds) of each other. Second, the LOR 35 formed by the straight line connecting the two detectors that detect the coincident event should pass through the field of view in the PET imaging system 12. Events that cannot be paired are discarded. Coincident event pairs are located and recorded as coincident data packets, which are transmitted via the physical communication link 78 to the classifier / histogram generator 80 in the image reconstruction processor 50.

[0034] The image reconstruction processor 50 includes a classifier / histogram generator 80. During operation, the classifier / histogram generator 80 generates a data structure known as a histogram. The histogram includes a large number of cells, where each cell corresponds to a unique pair of detector crystals in a PET scanner. Because PET scanners typically include thousands of detector crystals, the histogram typically includes millions of cells. Each cell of the histogram also stores a count value representing the number of coincidence events detected by the pair of detector crystals of that cell during the scan. At the end of the scan, the data in the histogram is used to reconstruct an image of the patient. A complete histogram containing all the data from the scan is typically referred to as a "result histogram." The term "histogram" typically refers to the components of the scanner, such as a processor and memory, that perform the function of creating the histogram.

[0035] The image reconstruction processor 50 also includes a memory module 82, an image CPU 84, an array processor 86 and a communication bus 88. During operation, the classifier / histogram generator 80 counts all events that occur along each projection ray and organizes the events into 3D data. The 3D data or sinogram is organized as a data array 90 in an exemplary embodiment. The data array 90 is stored in the memory module 82. The communication bus 88 is linked to the communication link 52 through the image CPU 84. The image CPU 84 controls communication through the communication bus 88. The array processor 86 is also connected to the communication bus 88. The array processor 86 receives the data array 90 as input and reconstructs the image in the form of an image array 92. The resulting image array 92 is then stored in the memory module 82.

[0036] The images stored in the image array 92 are transmitted by the image CPU 84 to the operator workstation 46. The operator workstation 46 includes a CPU 94, a display 96, and an input device 98. The CPU 94 is connected to the communication link 52 and receives input, such as user commands, from the input device 98. The input device 98 can be, for example, a keyboard, a mouse, a touch screen panel, and / or a voice recognition system, etc. Through the input device 98 and the associated control panel switches, the operator can control the operation of the PET imaging system 12 and the positioning of the patient 22 for scanning. Similarly, the operator can use the program executed by the workstation CPU 94 to control the display of the resulting images on the display 96 and can perform image enhancement functions.

[0037] The detector ring assembly 40 includes a plurality of detector units (e.g., detector blocks). The detector unit may include a plurality of detectors (e.g., photoelectric sensors such as SiPM devices), light guides, scintillation crystals, and ASICs, as will be described in further detail below. For example, the detector unit may include twelve SiPM devices, four light guides, and two or more analog ASICs. For another example, the detector unit may include between 16 and 64 SiPM devices and one or more ASICs, excluding light guides.

[0038] For example, the detector block 63 may include an array of scintillation crystals and an array of SiPM devices. The detector block 63 may be arranged to have a one-to-one coupling between the scintillation crystals and the SiPM devices, wherein each scintillation crystal is optically isolated from an adjacent scintillation crystal via a separate reflector or an optically shared reflector, as will be further described below. Each SiPM device may be read out by an independent front-end readout circuit via a timing pick-up channel and an energy channel. The signals from the timing pick-up channel and the energy channel may be ultimately sent to a time-to-digital converter (TDC) or an analog-to-digital converter (ADC) after passing through one or more comparators.

[0039] Reference now Figure 3 , a schematic depiction of an exemplary PET detector block 300 is shown. The PET detector block 300 may be a block comprising Figure 2 One of the detector units of the detector ring assembly 40, for example, the PET detector block 300 may be the detector block 63. In some examples, the PET detector block 300 may be configured to have a one-to-one coupling between the scintillator crystal and the SiPM device.

[0040] PET detector block 300 includes a scintillation crystal array 302 and an SiPM device array 306. In some examples, the scintillation crystal is a BGO scintillation crystal. Scintillation crystal array 302 includes a plurality of crystals 304, and SiPM device array 306 includes a plurality of SiPM devices 308. In one example, each detector block can have between 16 and 64 SiPM devices and scintillation crystals. Each of the plurality of crystals 304 is optically coupled to one of the plurality of SiPM devices 308 via coupling wires 310 as shown. In some examples, each of the plurality of crystals 304 can be optically isolated from each other, such as Figure 4 In other examples, multiple pairs of crystals share light while being optically isolated from other adjacent crystals, such as Figure 5A As depicted in .

[0041] For example, the first scintillator crystal 314 can be optically coupled to the first SiPM device 312 so that when one or more gamma rays collide with the first scintillator crystal 314 (e.g., are stopped in the first scintillator crystal), the first SiPM device 312 detects light generated from the first scintillator crystal 314 and sends a signal to be read by a corresponding readout circuit.

[0042] Although in Figure 3 3. Although not shown in the drawings, the PET detector block 300 may be surrounded by a reflective film (eg, polytetrafluoroethylene or other mirrored or diffuse material). The reflective film may be a light-proof cover that does not allow light emitted by the scintillation crystals to pass therethrough.

[0043] like Figure 4 As shown in , in a first embodiment, each crystal of the scintillation crystal array of a detector block as disclosed herein can be optically isolated from each other. Figure 4 A schematic side view of a portion of a detector block 400 is specifically depicted, illustrating a row 402 of scintillation crystals within an array. Each scintillation crystal within the row 402 of scintillation crystals can be optically coupled to one SiPM device in a row 408 of SiPM devices. Figure 3 Like the crystal and SiPM device, the scintillation crystal described in this article can be a BGO scintillation crystal.

[0044] Each SiPM device within the detector block 400 can have a size equal to the size of the corresponding scintillation crystal. For example, the first scintillation crystal 404 can have a first width 450, and the first SiPM device 412 can have a second width 452. The first width 450 and the second width 452 can be equal to each other. In a similar manner, the area of ​​the end of the first scintillation crystal 404 can be equal to the area of ​​the end of the first SiPM device 412. In this way, each of the SiPM devices can be configured to be optically coupled to one of the crystals without overlap between them.

[0045] The detector block 400 may also include a plurality of split reflectors 406. Each of the split reflectors 406 may be positioned directly adjacent to one of the scintillation crystals 402. The plurality of split reflectors 406 may be configured to optically isolate adjacent scintillation crystals from one another so that each scintillation crystal emits light only toward one SiPM device in response to photons blocked therein. For example, the first scintillation crystal 404 may be adjacent to the second scintillation crystal 414 and the third scintillation crystal 416. The first scintillation crystal 404 may be positioned between the second scintillation crystal 414 and the third scintillation crystal 416. The first split reflector 418 may be positioned between the first scintillation crystal 404 and the second scintillation crystal 414, and the second split reflector 420 may be positioned between the first scintillation crystal 404 and the third scintillation crystal 416.

[0046] When the gamma ray 410 (eg, a photon) interacts with the first scintillator crystal 404, light is emitted, such as Figure 4 404. The first and second separate reflectors 418 and 420 can reflect the emitted light so that it is not shared with the second and third scintillator crystals 414 and 416. Therefore, the light can be sensed by the first SiPM device 412 without being sensed by any adjacent SiPM device. In this way, the signal generated by the first SiPM device 412 can be known for the gamma rays interacting with the first scintillator crystal 404.

[0047] like Figure 5A As shown, in a second embodiment, pairs of adjacent scintillation crystals can share light while being optically isolated from other scintillation crystals in the crystal array. Figure 5A A schematic side view of a portion of a detector block 500 is depicted in detail, illustrating a row 502 of scintillation crystals within the array. The row 502 of scintillation crystals may be segmented into pairs of scintillation crystals, for example a first pair 550 of crystals may include a first crystal 510 and a second crystal 512.

[0048] The detector block 500 may also include a plurality of split reflectors 504 and a plurality of light-sharing reflectors 506, wherein each pair of crystals is optically isolated from a neighboring pair of crystals by one of the plurality of split reflectors 504, and each of the plurality of light-sharing reflectors 506 is positioned between crystals of a pair of crystals. Each of the plurality of light-sharing reflectors 506 may be configured to allow emitted light to pass through and provide interaction depth information / estimation to a readout electronics device in communication with the SiPM device.

[0049] For example, the first crystal 510 and the second crystal 512 of the first pair of crystals 550 may be separated by a light-sharing reflector 514, so that light emitted by the first crystal 510 or the second crystal 512 may be shared with the other of the first crystal 510 and the second crystal 512. The first crystal 510 may be adjacent to a third crystal 516, which is one of the crystals in the second pair of crystals 552. The first crystal 510 may be optically separated from the third crystal 516 by a first separation reflector 520, so that light emitted by the first crystal 510 is not shared with the third crystal 516. The second crystal 512 may be adjacent to a fourth crystal 518, which is one of the crystals in the third pair of crystals 554. The second crystal 512 may be optically separated from the fourth crystal 518 by a second separation reflector 522.

[0050] When the gamma ray 560 interacts with the first crystal 510, light is emitted. Because the first pair of crystals 550 can share light via the light sharing reflector 514, the emitted light can be shared with the second crystal 512, and thus the light can be sensed by both the first SiPM device 524 optically coupled to the first crystal 510 and the second SiPM device 526 optically coupled to the second crystal 512. In contrast, due to the first split reflector 520 and the second split reflector 522, the light may not be shared with the third crystal 516 or the fourth crystal 518, and thus neither the third SiPM device 530 optically coupled to the third crystal 516 nor the fourth SiPM device 528 optically coupled to the fourth crystal 518 can sense the light emitted by the first crystal 510.

[0051] Having more than one SiPM device sense light emitted by a crystal can allow for increased depth of information relayed through the various readout electronics. In some examples, the amount of light shared can depend on the location of the interaction between the gamma ray (e.g., gamma ray 560) and the scintillation crystal (e.g., first crystal 510), such as height within the crystal and / or lateral position within the crystal. By measuring signals from the SiPM devices (e.g., first SiPM device 524 and second SiPM device 526) that detected the emitted light, interaction depth information can be determined and used for timing measurement corrections. As will be described with respect to Figure 7 As further described, in either the first or second embodiments, each SiPM device may have a separate front-end readout circuit, and thus with both SiPM devices providing information to the front-end readout circuit, the depth of interaction may be determined with respect to detection time correction.

[0052] Figure 5B The top view schematically depicts Figure 5AIn some examples, the row of scintillation crystals 502 may be one of a plurality of rows of scintillation crystals within a crystal array. For example, four rows of six scintillation crystals may form an array. Each row of scintillation crystals may be optically isolated from each other. Figure 5B As shown in , multiple pairs of crystals are optically isolated from each other via multiple separation reflectors 504, and the crystals of each pair are separated by a light-sharing reflector 506.

[0053] The additional split reflector 570 can be positioned perpendicular to the plurality of split reflectors 504. In some examples, the additional split reflector 570 can optically isolate the rows of scintillator crystals from each other so that the row of scintillator crystals 502 does not share light with adjacent rows. In other examples, the additional split reflector 570 optically isolates the row of scintillator crystals from other crystal arrays when the row of scintillator crystals is at the boundary of the array. In this way, light emitted from the crystals can be directed toward the corresponding target SiPM device without being directed to non-target SiPM devices or other arrays, thereby maintaining information about the location of gamma rays and annihilation events.

[0054] Now turn to Figure 6 , showing that it can be Figure 5A and Figure 5B A schematic depiction of examples of light-sharing reflectors used in detector block 500 of FIG. 5 is depicted, including a first exemplary light-sharing reflector 602 , a second exemplary light-sharing reflector 604 , and a third exemplary light-sharing reflector 606 .

[0055] The first exemplary light-sharing reflector 602 can share light based on a varying level of reflectivity and transparency. For example, the first end 610 of the first exemplary light-sharing reflector 602 can reflect all light, while the second end 612 of the first exemplary light-sharing reflector 602 transmits all light. Between the first end 610 and the second end 612, the amount of light reflection and the amount of light transmission (e.g., percentage) can gradually change. For example, a first area 614 closer to the first end 610 can reflect more light and transmit less light than a second area 616 closer to the second end 612. In some examples, the level of variation can be based on or otherwise depend on height, where areas farther from the detector block reflect or transmit more light than areas closer to the detector block.

[0056] Compared to the first exemplary light-sharing reflector 602, the second exemplary light-sharing reflector 604 and the third exemplary light-sharing reflector 606 may have a uniform amount of light reflection / light transmission throughout, but may be shaped to taper toward one end (e.g., the first end 610) so that portions of one or more corresponding scintillator crystals are not blocked (e.g., reflected) by the light-sharing reflector. In this way, light may be shared between adjacent crystals using any of the above light-sharing reflectors. The configuration of the light-sharing reflector may be known, such as to provide additional interaction depth estimates for interactions between gamma rays and scintillator crystals. For example, measurements made by the corresponding SiPM device may provide information about the location of the interaction and thus allow for time measurement corrections.

[0057] Now go to Figure 7 , the block diagram shows the individual front-end readout circuits of the SiPM devices of the detector block 400. Each SiPM device in the SiPM device row 408 and the SiPM device array composed of multiple SiPM device rows can have an independent front-end readout circuit. Figure 7 The readout electronics of the front-end buffers (e.g., front-end amplifiers) depicted in the figure can be part of one or more ASICs. Each of the front-end buffers (such as front-end buffer 702) can provide a signal for a separate measurement of a timestamp (e.g., a detection time / arrival time of an annihilation event) via a timing pick-off channel (such as timing channel 704). Each of the front-end buffers can also provide a separate signal for an energy (such as signal amplitude and / or charge) measurement via an energy channel (such as energy channel 706). Each of the timing measurement and the energy measurement can be shaped by a filter. In some examples, an optimal timing filter can shape the pulse signal to maintain the rise time but quickly restore the baseline so that the time sensing circuit system of the timing channel can accurately detect the timing of the pulse. The optimal filter of the energy channel can filter out high frequencies of the signal to improve the measurement of the total signal charge corresponding to the energy deposited in the crystal.

[0058] For example, when the gamma ray 410 interacts with the first scintillator crystal 404 and emits light therefrom, the first SiPM device 412 may detect the light and send analog signals for the timing channel and the energy channel to the front-end buffer 702. These signals may be sent to one or more comparators, TDCs, and ADCs through the timing channel and the energy channel, respectively. Figure 8 as shown in .

[0059] Figure 8A high-level block diagram of a readout circuit 800 for a SiPM device is shown. As previously described, the readout circuit 800 can receive a signal from the SiPM device when the SiPM device detects light from a scintillation crystal. In some examples, the scintillation crystal can be a BGO scintillation crystal, and the light can include Cherenkov light with detectable Cherenkov photons. A PET detector block having multiple SiPM devices can include multiple readout circuits, one for each of the SiPM devices, and each SiPM device can be read out independently of the other SiPM devices. Fig. 9 A method for signal processing via a readout circuit 800 is presented in FIG.

[0060] The readout circuit 800 of the SiPM may include a front-end buffer 802 that is communicatively coupled to both the timing pickup circuit and the energy circuit. The timing pickup circuit may include a timing channel 804 that communicates (e.g., is connected in series) with the front-end buffer 802. The timing channel 804 may also communicate with a timing filter 806 such as a timing filter, through which the pulse can be shaped. The timing channel 804 and the timing filter 806 may also communicate with a plurality of comparators 808. The plurality of comparators 808 may include a first comparator 810 (e.g., a primary comparator 810) and a second comparator 812 (e.g., a secondary comparator 812). The plurality of comparators 808 may allow additional information about the pulse rise time, wherein the signal passing through each comparator in the plurality of comparators 808 is compared with a threshold signal. In some examples, the threshold for timing pickup may be in the range of 1 / 3 SPE to 10 SPE. The comparator may also provide dark count suppression (e.g., a threshold of 1.5 SPE). The first comparator 810 can be in communication with a delay circuit 814 and, when the secondary comparator detects a signal above a threshold during the delay, produces a verified primary signal 816. Both the first comparator 810 and the second comparator 812 can be in communication with one or more TDCs 818. In this manner, the readout circuit 800, as well as other readout circuits of the PET systems described herein, can be triggered at a first level of 1 / 3 single photoelectrons (SPEs) and verified at a second level of 1.5 SPEs.

[0061] The energy circuit of the readout circuit 800 may include an energy channel 820 that communicates with an energy filter 822 (such as an energy filter by which pulses can be shaped) and one or more ADCs 828. The energy filter 822 can communicate with an energy comparator 824 in the plurality of comparators 808, which can compare signal energy (e.g., signal amplitude and charge) to a threshold. The energy comparator 824 can communicate with an energy trigger 826, which communicates with both one or more TDCs 818 and one or more ADCs 828. The one or more TDCs 818 and the one or more ADCs 828 can digitize arrival time, rise time (e.g., with multiple time stamps), and energy information (e.g., amplitude, charge, etc.).

[0062] As noted, readout circuit 800 is an example of a readout circuit system for a PET system according to the present disclosure. Each SiPM device of a PET detector can have a separate readout circuit, so that multiple readout circuit systems are included in the PET system. In this way, the threshold of the comparator can be lowered because the multiplexed signals from multiple SiPM devices require a higher threshold than that required for a single readout circuit. Therefore, each SiPM device can have a TDC and ADC, thereby providing improved timing resolution. In addition, in some examples, reconstruction of signal rise time, as provided by filters and multiple timing measurements, can include classifying events into two categories: "fast" and "slow".

[0063] Fig. 9 A diagram is shown for Figure 8 Flow chart of an exemplary method 900 for signal processing by the readout circuit 800 described above. The method 900 may be used with reference to Figures 1 to 8 For example, the method 900 may be implemented by a combination of a PET detector block (e.g., detector block 300, detector block 400, and / or detector block 500) and an analog ASIC that includes readout electronics (e.g., readout circuit 800) of a SiPM device communicatively coupled to the PET detector block. However, it should be understood that the method may be performed using different systems and components without departing from the scope of the present disclosure.

[0064] At 902, method 900 includes: providing a SiPM device (eg, Figure 4At a first SiPM device 412 of the present invention, an analog signal corresponding to a scintillation event within a scintillation crystal coupled to the SiPM device is generated. As previously described, the detector block may include a plurality of SiPM devices arranged in an array with a one-to-one coupling arrangement, with the scintillation crystals arranged in a corresponding array. When a gamma ray collides with a scintillation crystal coupled to the SiPM device, a scintillation event is generated, and light is emitted as photons (including Cherenkov photons for BGO scintillation crystals as included herein). The emitted light of the scintillation event is detected by the SiPM device, and the SiPM device generates an analog signal of the scintillation event. The analog signal may include information about arrival time, rise time, amplitude, charge, and / or other information.

[0065] At 904, method 900 includes transmitting the signal from the SiPM to a front-end buffer of the ASIC. Each SiPM of the detector block may communicate with a separate front-end buffer, and thus multiple front-end buffers may be included in the ASIC.

[0066] At 906, method 900 includes transmitting the signal to a timing pick-up circuit and an energy circuit. As described, the timing pick-up circuit may include a timing channel (e.g., timing channel 804) in communication with a timing filter (e.g., timing filter 806), and the energy circuit may include an energy channel (e.g., energy channel 820) in communication with an energy filter (e.g., energy filter 822).

[0067] At 908, method 900 includes shaping the signal via a timing filter and an energy filter. The signal may be transmitted to the energy filter via an energy channel and to the timing filter via a timing channel. The energy filter and the timing filter may be optimal filters that may change the pulse shape to allow the timing pulse and the energy pulse from the front-end buffer to be compatible with the digitizing electronics (e.g., comparators, TDCs, and ADCs) of each channel.

[0068] At 910, method 900 includes transmitting the shaped signal through a plurality of comparators. The plurality of comparators may include a first comparator, a second comparator, and an energy comparator. The first comparator and the second comparator may compare the timing signal with a threshold value, such as a range of 1 / 3 SPE to 10 SPE for time sensing. In addition, by measuring the delay between the primary comparator output and the secondary comparator output, the rise time may be estimated in the first nanosecond. The energy comparator may compare the pulse amplitude and / or charge input with a threshold value or a predefined input charge to generate an output signal.

[0069] The configuration of multiple comparators can allow for improved timing resolution by reducing the threshold and increasing the ability to detect the first photon of a scintillation event. Due to the nature of emitting Cherenkov photons, BGO scintillation crystals can produce detectable light that falls within the threshold range, and therefore using the timing picking method using multiple comparators as described herein, Cherenkov photons can be detected and timing resolution can be improved.

[0070] At 912, method 900 includes transmitting signals verified and / or generated by the plurality of comparators to a TDC and / or ADC. Figure 8 As described, the readout circuitry may include one or more TDCs (eg, one or more TDCs 818 ) and one or more ADCs (eg, one or more ADCs 828 ).

[0071] At 914, method 900 includes: generating a digital signal corresponding to the scintillation event. The one or more TDCs can measure timing signals (e.g., arrival time and / or rise time signals) and convert the signal time to a digital signal for further processing. The one or more ADCs can measure the energy signal output from the energy channel and sent to the ADC after passing through the energy filter and convert it to a digital signal. One or more digital signals corresponding to the scintillation event can be generated by the TDC and the ADC, and the one or more digital signals can provide information such as arrival time, rise time estimation, energy amplitude, energy charge, etc.

[0072] At 916, method 900 includes transmitting the digital signal to a data acquisition processor (eg, Figure 2 The data acquisition process 48) can be used for localization. The data acquisition processor can use this digital signal and other digital signals transmitted and converted from other SiPM devices of the detector ring to localize annihilation events in the patient imaged by the PET system. The improved timing resolution as described herein can allow the localization margin of the event to be reduced and therefore allow higher image quality.

[0073] The technical effect of the systems and methods described herein is that Cherenkov photons can be detected with high timing resolution via a readout circuit including multiple comparators and one or more filters, and therefore BGO scintillation crystals can be implemented into a TOF PET system. The threshold of the comparator within the circuit system can be set to 1 / 3 SPE, with verification at 1.5 to 2.5 SPE, thereby suppressing dark counts while detecting emitted Cherenkov light to increase timing resolution.

[0074] The present disclosure also provides support for an imaging system, the imaging system comprising: a detector block, the detector block comprising a silicon photomultiplier (SiPM) device array coupled to a crystal in a scintillation crystal array in a one-to-one coupling arrangement, wherein each SiPM device in the SiPM device array transmits a signal to an independent front-end readout circuit of one or more analog application-specific integrated circuits (ASICs). In a first example of the system, the detector block forms part of a detector ring assembly of the imaging system. In a second example of the system, optionally including the first example, each of the scintillation crystals in the scintillation crystal array is optically isolated from each other via a separate reflector. In a third example of the system, optionally including one or both of the first and second examples, the scintillation crystal array comprises a plurality of pairs of adjacent scintillation crystals that share light via a light-sharing reflector, and the plurality of pairs of adjacent scintillation crystals are optically isolated from each other via a separate reflector. In a fourth example of the system, optionally including one or more or each of the first to third examples, information from the plurality of pairs of adjacent scintillation crystals provides an interaction depth estimate. In a fifth example of the system, optionally including one or more or each of the first to fourth examples, the interaction depth estimate is applied to timing measurement correction. In a sixth example of the system, optionally including one or more examples or each example of the first to fifth examples, each independent front-end readout circuit includes a plurality of comparators and one or more filters. In a seventh example of the system, optionally including one or more examples or each example of the first to sixth examples, the plurality of comparators include a first comparator, a second comparator, and an energy comparator. In an eighth example of the system, optionally including one or more examples or each example of the first to seventh examples, the plurality of comparators are configured to have a minimum timing pick-up threshold of 1 / 3 single photoelectrons (SPE). In a ninth example of the system, optionally including one or more examples or each example of the first to eighth examples, the scintillation crystal array is a bismuth germanate oxide (BGO) scintillation crystal array, each BGO scintillation crystal is configured to emit Cherenkov light; and each SiPM device in the SiPM device is configured to sense the Cherenkov light emitted by the corresponding scintillation crystal. In a tenth example of the system, optionally including one or more or each of the first to ninth examples, each independent front-end readout circuit includes one or more time-to-digital converters (TDCs) and one or more analog-to-digital converters (ADCs) configured to digitize arrival time, rise time, and energy signals from the corresponding SiPM device. In an eleventh example of the system, optionally including one or more or each of the first to tenth examples, each independent front-end readout circuit is configured for dark count suppression at 1.5 SPE.

[0075] The present disclosure also provides support for a method for an imaging system, the method comprising: sensing light emitted by a scintillation crystal via a silicon photomultiplier (SiPM) device of a detector block in response to an annihilation event; transmitting an analog signal from the SiPM device to a front-end buffer of a readout circuit; shaping the analog signal with one or more filters of the readout circuit; comparing the analog signal with one or more thresholds via a plurality of comparators of the readout circuit; and digitizing the analog signal with one or more time-to-digital converters (TDCs) and one or more analog-to-digital converters (ADCs). In a first example of the method, the method further comprises: validating the analog signal to suppress dark counts. In a second example of the method, optionally including the first example, the SiPM device is one of a plurality of SiPM devices in an array, wherein each SiPM device in the array is read out by a separate readout circuit. In a third example of the method, optionally including one or both of the first and second examples, the analog signal comprises a timing signal and an energy signal, the timing signal is shaped by a timing filter in the one or more filters, and the energy signal is shaped by an energy filter in the one or more filters.

[0076] The present disclosure also provides support for a time-of-flight (TOF) positron emission tomography (PET) system, the system comprising: a detector ring, the detector ring comprising a plurality of detector blocks, each of the plurality of detector blocks comprising an array of silicon photomultiplier (SiPM) devices, each SiPM device in the array coupled to a scintillation crystal in an array of scintillation crystals, wherein each crystal in the array of scintillation crystals is configured to emit Cerenkov light in response to interaction with gamma rays; and a readout electronic device, the readout electronic device comprising a front-end readout circuit for each SiPM device in the array of SiPM devices, wherein the readout electronic device is configured to read out a signal of the detected Cerenkov light. In a first example of the system, each scintillation crystal in the array of scintillation crystals is a bismuth germanate oxide (BGO) scintillation crystal. In a second example of the system, optionally including the first example, each front-end readout circuit comprises a front-end buffer, a timing pickup circuit, and an energy circuit, wherein the timing pickup circuit comprises a plurality of comparators configured to estimate a rise time within 1 nanosecond to 5 nanoseconds of a pulse. In a third example of the system, optionally including one or both of the first and second examples, the readout electronics is configured to be triggered at a first level of 1 / 3 single photoelectrons (SPEs) and verified at a second level of 1.5 SPEs.

[0077] As used herein, the elements or steps listed in the singular and beginning with the word "one" or "a kind of" should be understood as not excluding a plurality of said elements or steps, unless such exclusion is explicitly stated. In addition, the reference to "one embodiment" of the present invention is not intended to be interpreted as excluding the existence of additional embodiments that also include the cited features. In addition, unless explicitly stated to the contrary, "comprising", "including" or "having" an embodiment of an element or multiple elements with a specific characteristic may include additional such elements that do not have the characteristic. The terms "including" and "in..." are used as the concise language equivalents of the corresponding terms "including" and "wherein". In addition, the terms "first", "second" and "third" etc. are only used as marks, and are not intended to impose numerical requirements or specific positional order on their objects.

[0078] This written description uses examples to disclose the invention, including the best mode, and also to enable a person of ordinary skill in the relevant art to practice the invention, including making and using any devices or systems and performing any included methods. The patentable scope of the invention is defined by the claims, and may include other examples that occur to a person of ordinary skill in the art. Such other examples are intended to fall within the scope of the claims if they have structural elements that do not differ from the literal language of the claims, or if they include equivalent structural elements with insignificant differences from the literal language of the claims.

Claims

1. An imaging system, comprising: A detector block (300) includes an array of silicon photomultiplier (SiPM) devices (306) coupled to a scintillation crystal array (302) in a one-to-one coupling arrangement, wherein each SiPM device in the array of SiPM devices (306) transmits a signal to an independent front-end readout circuit (800) of one or more analog application specific integrated circuits (ASICs).

2. The imaging system of claim 1, wherein the detector block (300) forms part of a detector ring assembly (40) of the imaging system (10).

3. The imaging system of claim 1, wherein each of the scintillation crystals (304) are optically isolated from each other in the scintillation crystal array (302) via a separation reflector (406).

4. The imaging system of claim 1, wherein the scintillation crystal array (302) comprises a plurality of pairs of adjacent scintillation crystals (550) that share light via a light sharing reflector (506), and the plurality of pairs of adjacent scintillation crystals are optically isolated from each other via a separation reflector (504).

5. The imaging system of claim 4, wherein information from the plurality of pairs of adjacent scintillator crystals (550) provides an interaction depth estimate.

6. The imaging system of claim 5, wherein the interaction depth estimate is applied for timing measurement correction.

7. The imaging system of claim 1 , wherein each independent front-end readout circuit (800) comprises a plurality of comparators (808) and one or more filters (806, 822), the plurality of comparators comprising a first comparator (810), a second comparator (812), and an energy comparator (824).

8. The imaging system of claim 7, wherein the plurality of comparators (808) are configured to have a minimum timing pick-off threshold of 1 / 3 single photoelectrons (SPEs).

9. The imaging system of claim 1, wherein: The scintillation crystal array (302) is a bismuth germanate oxide (BGO) scintillation crystal array, each BGO scintillation crystal being configured to emit Cherenkov light; and Each of the SiPM devices (308) is configured to sense the Cerenkov light emitted by a corresponding scintillation crystal.

10. The imaging system of claim 1, wherein each independent front-end readout circuit (800) comprises one or more time-to-digital converters (TDCs) (818) and one or more analog-to-digital converters (ADCs) (828), the time-to-digital converters and the analog-to-digital converters configured to digitize arrival time, rise time, and energy signals from the corresponding SiPM device.

11. The imaging system of claim 9, wherein each independent front-end readout circuit (800) is configured for dark count suppression at 1.5 SPE.

12. A method for an imaging system, the method comprising: In response to the annihilation event, sensing light emitted by the scintillation crystal via a silicon photomultiplier (SiPM) device of a detector block (902); transmitting an analog signal from the SiPM device to a front-end buffer of a readout circuit (904); shaping the analog signal using one or more filters of the readout circuit (908); comparing the analog signal to one or more thresholds via a plurality of comparators of the readout circuit (910); and The analog signal is digitized (914) using one or more time-to-digital converters (TDCs) and one or more analog-to-digital converters (ADCs).

13. The method of claim 12, further comprising validating the analog signal to suppress dark counts (908).

14. The method of claim 12, wherein the SiPM device is one SiPM device among a plurality of SiPM devices (308) in an array (306), wherein each SiPM device in the array is read out by a separate readout circuit (800).

15. The method of claim 12, wherein the analog signal comprises a timing signal and an energy signal, the timing signal being shaped by a timing filter (806) among the one or more filters, and the energy signal being shaped by an energy filter (822) among the one or more filters.

Images