Lateral crystal photodiode readout and switching diode network for processing core events

Abstract

A positron emission tomography (PET) scanner may include a plurality of gamma radiation detector modules arranged to form a detector ring. Each detector module may include an array of elongated scintillation crystals. Relative to the detector ring, each elongate scintillation crystal includes a proximal face, two axially oriented sides, two lateral axis oriented sides, and a distal face oriented into the detector ring to receive gamma photons. An array of photosensors is positioned along a first axially oriented side of each elongate scintillation crystal to detect scintillation photons. A reflective material is positioned on the proximal end face, the distal end face, the lateral axis oriented side and the second axially oriented side of each elongated scintillation crystal to internally reflect scintillation photons. In various embodiments, a dual channel processing circuit provides different timing and energy signals from a photosensor.

Description

CROSS-REFERENCE TO RELATED APPLICATIONS

[0001] This application claims priority to U.S. Patent Application No. 18 / 365,755, filed on August 4, 2023, entitled “Lateral Crystal Photodiode Readouts and Switched Diode Networks for Processing Nuclear Events,” which application claims priority and the benefit of U.S. Provisional Patent Application No. 63 / 378,739, filed on October 7, 2022, entitled “Lateral Crystal Photodiode Readout and Switched Diode Network for Processing Nuclear Events,” and U.S. Provisional Patent Application No. 63 / 502,115, filed on May 14, 2023, entitled “Lateral Crystal Photodiode Readout and Switched Diode Network for Processing Nuclear Events,” under 35 U.S.C. §119, each of which is incorporated herein by reference in its entirety. Background Art

[0002] The present disclosure relates to gamma radiation measurement, scintillation detector circuits, and coincidence circuit arrangements. The use of the systems, methods, practical applications, and embodiments described in the present disclosure may be understood in the context of the following disclosures, each of which is hereby incorporated by reference in its entirety: E. Berg and S. Cherry, "Innovations in instrumentation for positron emission tomography," Seminars in nuclear medicine, Vol. 48, No. 4, pp. 311-331, 2018; S. R. Cherry and M. Dahlbom, PET: Physics, Instrumentation, and Scanners, 2006; J. S. Reddin, J. S. Scheuermann, D. Bharkhada, A. M. Smith, M. Casey, M. Conti, and J. S. Karp, "Performance evaluation of the SiPM-based Siemens Biograph Vision PET / CT system," in IEEE Nuclear Science Symposium and Medical Imaging Conference Proceedings (NSS / MIC), Sydney, 2018; G. F. Knoll, Radiation Detection and Measurement, Hoboken: John Wiley & Sons, 1996. Wiley & Sons, Inc., 2000; S. Gundacker, E. Auffray, ND Vara, B. Frisch, H. Hillemanns, P. Jarron, B. Lang, T. Meyer, S. Mosquera-Vazquez, E. Vauthey, and P. Lecoq, “SiPM time resolution: From single photon to saturation,” Nuclear Instruments and Methods in Physics Research Section A: Accelerators, Spectrometers, Detectors and Associated Equipment, Volume 718, Pages 569-572, 2013; M.Conti, “Focus on time-of-flight PET: the benefits of improved time,” European Journal of Nuclear Medicine Molecular Imaging, vol. 38, pp. 1147–1157, 2011; S. Strother, M. Casey, and E. Hoffman, “Measuring PET scanner sensitivity: relating count rates to image signal-to-noise ratios using noise equivalent counts,” IEEE Trans Nuclear Science, vol. 37, pp. 783–788, 1990; and Saint Gobain, “LYSO Scintillation Crystals,” June 2018. [Online]. Available at: https: / / www.crystals.saint-gobain.com / radiation-detection-scintillators / crystal-scintillators / lyso-scintillation-crystals. . BRIEF DESCRIPTION OF THE DRAWINGS

[0003] Figure 1A A perspective view of a block diagram of a portion of a positron emission tomography (PET) detector system is shown according to one embodiment.

[0004] Figure 1B Another view of a block diagram of a portion of a PET detector system is shown according to one embodiment.

[0005] Figure 2 A positron source with multiple lines-of-response is shown according to one embodiment.

[0006] Figure 3 A tube-based block detector having four photomultiplier tubes is shown according to one embodiment.

[0007] Figure 4 is a block diagram of a ring of detector modules having multiple lines of response according to one embodiment.

[0008] Figure 5A A block diagram of a ring of detector modules with a single off-center line of response (LoR) is shown according to one embodiment.

[0009] Figure 5B A block diagram of a ring of detector modules having multiple eccentric lines of response is shown according to one embodiment.

[0010] Figure 6 Shown are response lines segmented according to their temporal and spatial relationships according to one embodiment.

[0011] Figure 7 is a simplified schematic diagram of a high-speed diode network according to one embodiment.

[0012] Fig. 8A is a two-port network lumped element model of a conductive transmission line according to one embodiment.

[0013] Figure 8B A distributed model based on lumped elements of unit length is shown according to one embodiment.

[0014] Fig. 9 Shown is a time domain diagram of two nuclear pulse signals according to one embodiment.

[0015] Fig.10 According to one embodiment, Fig. 9 Energy spectral density diagram of two nuclear pulse signals.

[0016] Fig.11 is a hierarchical block diagram of a high-speed diode network according to one embodiment.

[0017] Fig.12 A detector module having blocks of elongated scintillator crystals arranged in a two-dimensional array and having a photosensor array for lateral photosensor readout is shown according to one embodiment.

[0018] Fig.13A is a diagram of a single elongated scintillator crystal pixel having multiple photosensor arrays for lateral photosensor readout according to one embodiment.

[0019] Fig. 13B An exploded view of a single elongated scintillator crystal is shown according to one embodiment.

[0020] Fig. 13C is a diagram of a single elongated scintillator crystal pixel with a photosensor array, interposer, and connector array for lateral photosensor readout according to one embodiment.

[0021] Fig.14AA sub-divided elongated scintillator crystal pixel having a photosensor array for lateral photosensor readout is shown according to one embodiment.

[0022] Fig. 14B An exploded view of a sub-blocked scintillation pixel is shown according to one embodiment.

[0023] Fig. 14C is a diagram of a sub-blocked scintillation pixel with a photosensor array, interposer, and connector array for lateral photosensor readout according to one embodiment.

[0024] Fig.15A is a diagram of an elongated scintillator crystal block with lateral readout photosensors and a plate for thermal management according to one embodiment.

[0025] Fig. 15B is a diagram of an elongated scintillator crystal block with a photosensor array, an interposer, and a connector array according to one embodiment.

[0026] Fig. 15C is a diagram of an elongated scintillator crystal block with a photosensor array, an interposer, a connector array, and a plate for thermal management according to one embodiment.

[0027] Fig.15D is a diagram of an elongated block of scintillator crystals connected to a dual channel processing circuit according to one embodiment.

[0028] Fig.16 is a diagram of an elongated scintillator crystal block having a double-sided interposer connected to a photosensor array of adjacent elongated scintillator crystals according to one embodiment.

[0029] Fig.17 A ring of detector modules is shown with axially oriented photosensor arrays on the sides of an elongated scintillator crystal, according to one embodiment.

[0030] Fig.18 A ring of detector modules with a transverse-axis oriented photosensor array on the side of an elongated scintillator crystal is shown according to one embodiment.

[0031] Fig.19 is a graph of sensitivity gain and fill factor loss for a lateral photosensor readout according to one embodiment.

[0032] Fig. 20 is a graph of total sensitivity gain for rectangular prismatic scintillator crystals of varying lengths having square end faces according to one embodiment.

[0033] Fig.21is a graph of the photon dynamic range of a silicon photomultiplier tube photosensor versus various sizes of an elongated rectangular prismatic scintillator crystal according to one embodiment.

[0034] Fig. 22 is a graph of light output versus temperature for a scintillator according to one embodiment. DETAILED DESCRIPTION

[0035] The embodiments of the systems and methods provided in this disclosure are not intended to limit the scope of this disclosure, but are merely representative of possible embodiments. In addition, the steps of the method do not necessarily need to be performed in any particular order, or even necessarily in order, and the steps do not necessarily need to be performed only once. The following description, in conjunction with the accompanying drawings, sets forth many specific details to provide a comprehensive understanding of the possible variations of the systems and methods. However, the disclosed concepts, systems, methods, devices, etc. can be implemented without some or all of the specific details. For the sake of clarity, technical materials known in the technical field related to this disclosure are not described in detail.

[0036] Positron emission tomography (PET) is a medical imaging scanning system that incorporates multiple gamma radiation detector modules that can be arranged in a multi-ring formation for the detection and subsequent image reconstruction of annihilation radiation emitting materials located within the scanner's field of view. Medical imaging of the human body using PET can be facilitated by intravenous injection of positron-based radiotracers with sufficient radioactivity duration to enable PET scanning. Examples of such tracers include fluorine-18 labeled glucose compounds, which are intended to identify hypermetabolic tumor activity. When a positron encounters an electron, an annihilation event occurs, emitting two oppositely directed gamma photons approximately 180° apart, each with an energy of approximately 511 kiloelectronvolts (keV).

[0037] The range over which a positron migrates before encountering a free electron is called the positron range. Depending on the initial kinetic energy of the positron and the density of electrons within the medium, the in vivo range is typically less than a millimeter for fluorine-18 based radiotracers. The kinetic energy of the positrons also leads to undesirable noncollinearity effects, whereby the annihilated gamma photons traverse slightly off 180° more or less, resulting in imaging errors of about 1.8 millimeters for common scanner apertures. Errors due to positron range and noncollinearity limit the resolution of some PET imaging techniques to about 2 mm for whole body human imaging.

[0038] Figure 1AA perspective view of a PET detector system 100 and positron sources 150 and 151 within a patient 110 (illustrated as a human torso) according to one embodiment is shown. In the illustrated embodiment, the PET detector system 100 includes three rings of gamma radiation detector modules. A first ring 170 of gamma radiation detector modules is closest to an observer, a second ring 180 of gamma radiation detector modules is located in the middle, and a third ring 190 of gamma radiation detector modules is farthest from the observer. In the illustrated example, each ring 170, 180, and 190 includes forty-eight gamma radiation detector modules. The number of detector modules and the number of detector module rings can vary depending on the specific application, cost considerations, target size, target resolution, etc.

[0039] Positron sources 150 and 151 represent positron active masses that result in annihilation radiation in which two opposing gamma photons are emitted in opposing directions. The opposing gamma photons emitted by positron source 150 are detected by detector modules 183 and 187. A processor or other circuitry can calculate a line of response (LoR) 130 and determine the location of positron source 150 using the known detector locations and the measured arrival times of the gamma photons detected by detector modules 183 and 187. A processor or other circuitry can identify the detected gamma photons as significant annihilation events based on the total detected energy being equal to or within the expected threshold range of 511 keV.

[0040] The reversed gamma photons emitted by positron source 151 are detected by detector modules 181 and 185. A processor or other circuitry can calculate line of response 131 and determine the position of positron source 151 using the known detector positions and the measured arrival times of the gamma photons detected by detector modules 181 and 185. As discussed in more detail below, if positron sources 150 and 151 emit their respective pairs of reversed gamma photons simultaneously (or nearly simultaneously), there is a risk that the processor or other circuitry may incorrectly calculate line of response 135.

[0041] Throughout the present disclosure, detector modules are described and illustrated as being used in a detector ring (e.g., a detector module ring). The scintillation crystal in each detector module has a distal face facing the detector (e.g., facing the ring, where the distal face is located on the inner diameter of the ring) to receive gamma photons. However, it is understood that the detector ring may include annular detector modules without forming a complete or full ring (e.g., a case with only top and bottom detector modules). In addition, the detector may be embodied as two (or more) detector module planar panels that work together to detect annihilation events. For example, the detector may include a first planar panel of the detector module and a second relatively planar panel of the detector module. The subject may be positioned in the detector by placing the subject between the two relatively planar panels. Thus, the detector module may include a plurality of scintillation crystals (e.g., in a one-dimensional or two-dimensional array) with their distal ends facing the detector (e.g., toward a cavity or space between the relatively panels so that the distal face is located on the inner diameter of the ring).

[0042] Figure 1B A side view of another example of a PET detector system 101 and a positron source 153 within a human body 111 according to one embodiment is shown. In the illustrated embodiment, the PET detector system 101 includes a ring of seventy-two gamma radiation detector modules. Backscattered gamma photons emitted by the positron source 153 are detected by detector modules 182 and 184. A processor or other circuitry can use the detector module positions and the arrival times of the gamma photons detected by the detector modules 182 and 184 to calculate the response line 137 and determine the location of the positron source 153.

[0043] Figure 2 A block diagram of a PET detector system 200 is shown according to one embodiment. The simplified block diagram includes a single ring of sixteen gamma radiation detector modules 210-225. According to one embodiment, multiple response lines 231, 232, and 233 are shown that intersect a positron activity cluster 250. Annihilation radiation from the positron activity cluster 250 produces a large number of equal and opposite simultaneous emissions of gamma photons. The system 200 can determine the origin point range of the positron along the response line 233 extending between two detection points thereof (e.g., within detector modules 223 and 215).

[0044] Additional response lines 231 and 232 are calculated during the measurement of subsequent annihilation events occurring within the mass 250. The time period during which the annihilation event is detected from the mass 250 is called a scan. The system 200 is used to capture a scan that provides multiple response lines to locate the origin point or origin region of the positron-emitting material. A PET scanner (e.g., including or embodied as a PET detector system described herein) can use a ring of detector modules 210-225 to create images of positron-emitting materials of different concentrations and locations without using a shielded collimator (such as a shielded collimator used in single photon emission computed tomography (SPECT)) that filters the tilted starting angle. In other embodiments, the PET scanner can use axial collimators (called septa) between multiple detector rings to reduce scattered radiation from the field of view (e.g., in the human body). However, due to faster scintillator speeds and advances in electronic processing, other PET scanners operate in a "3D" mode without septa, can handle greater event rates, and thus increase system sensitivity.

[0045] The double gamma events of annihilation radiation are distinguished from (e.g., identified relative to) a large number of other unrelated gamma events (commonly referred to as "single events") by using a time-based coincidence window. The coincidence window (e.g., which may be less than five nanoseconds (e.g., two nanoseconds)) is used to filter single events that are unlikely to occur from corresponding single annihilations within the field of view given the aperture of the scanner (corresponding to the diameter of the ring of detector modules 210-225) and the propagation speed of gamma radiation. For example, for a PET scanner with detector modules located in a ring with a diameter of 70 centimeters (e.g., an aperture of 70 centimeters), a positron annihilation at the center of the aperture will require approximately 1.167 nanoseconds for each gamma photon traveling at the speed of light (29.98 centimeters per nanosecond) to reach the scintillator of one detector module in the ring. Correspondingly, a positron annihilation located along the circumference of the aperture will require twice that time for one gamma photon to scintillate into a detector module on the other side of the ring. Therefore, the coincidence window and the selected pairing of detector modules that contain only reasonable geometric response lines reduce the number of irrelevant gamma events to be recorded as coincidence events.

[0046] When two gamma events meet the criteria for being recorded as a single coincidence event (e.g., within a time window and detected by an appropriate detector module), such an event is called a "prompt" event. A prompt event may still consist of two unrelated and independent gamma events that originate from different annihilation events or other radiation sources, which are called "random" coincidence events. However, when the two gamma detections do originate from a single positron annihilation, such an event is called a "real" coincidence event. Therefore, the number of "real" coincidence events that occur from real positron annihilations is equal to the number of "prompt" coincidence events (e.g., coincidence events from unknown sources) minus the number of "random" coincidence events (known or statistically determined to be random single coincidence events measured within the coincidence window).

[0047] In various embodiments, the PET scanner may further filter "scattered" events, which are generated when a gamma photon deposits part of its energy into nearby material and then propagates its remaining energy into the detector module. Since scattered events may be deflected, such scattered events are typically filtered because their point of origin is uncertain. Scattered events typically end up in the detector with energies far below their original 511keV. Therefore, filtering may involve measuring the energy deposited by the gamma event and rejecting such events that are below a low level discrimination (LLD) threshold. In addition, high level discrimination (ULD) may also be utilized to reject the possibility of any multiple events or single higher energy nuclear particle deposition that accumulate into the detector at nearly the same time. According to various embodiments, the PET scanner may utilize an LLD threshold between 325keV and 511keV and an ULD threshold between 511keV and 675keV.

[0048] Figure 3 A tube-based detector module 300 is shown having four photomultiplier tubes 320 (PMTs) in a quadrant-based detection subsystem according to one embodiment. The four photomultiplier tubes 320 are connected to a two-dimensional array 310 of elongated scintillator crystals 310. An exemplary two-dimensional array of elongated scintillator crystals 310 includes 56 elongated scintillator crystals in a 7x8 array. Some embodiments of gamma detector modules for PET scanners include inorganic scintillators (e.g., bismuth germanate, Bi4Ge3O 12 , called BGO, or lutetium oxide silicate, Lu2(SiO4)O, called LSO, Lu 2(1-x-y) Y 2x SiO5, called LYSO etc.), which is coupled to a light guide and connected to the quadrant of the photomultiplier tube.

[0049] The configuration shown may be referred to as a block detector module. The detector module may include one or more block detectors. In this case, the block detectors share an electronic channel for processing nuclear gamma single events from multiple elongated scintillator crystals 310. As shown, the elongated scintillator crystals 310 may be pixelated into multiple elongated scintillator crystals 310, which may be etched, polished and / or wrapped in a reflective material. Pixelating the elongated scintillator crystals 310 into an array of elongated rectangular prisms helps determine the location of the impinging gamma photons within a rectangular area on the end face of each elongated scintillator crystal 310.

[0050] In various embodiments, selected areas or faces of each elongated scintillator crystal 310 may be wrapped or coated with a reflective material. The light output from any given crystal may be collected using four photomultiplier tubes 320 to perform ratiometric measurements to determine which crystal experienced a gamma interaction, thereby localizing such events in the XY plane of the detector surface.

[0051] The motivations for sharing processing electronics for a single event on a block basis are cost constraints, reduced power consumption and dissipation, and small package volume. In a block processing configuration, the detector module 300 processes a single event exclusively through the processing electronics of the block, and any other incident radiation within the processing time of the same block is typically rejected. This processing time is called detector dead-time. For PET, defining the detection area of ​​the block is a consideration because if the block is too large, sharing processing electronics may impair count rate capability and sensitivity, while too small a block may be cost-prohibitive and complicate thermal considerations for practical volume.

[0052] In various embodiments, a solid-state photosensor (e.g., a silicon-based photosensor) can be used instead of a photomultiplier tube. Conventional semiconductor-based PIN photodiodes and / or avalanche photodiodes (APDs) generate analog signals proportional to the incident photon flux. However, the APD gains from the impact ionization that occurs at its high reverse bias, where the generated electron-hole pairs further collide with nearby atoms, thereby releasing additional charge carriers beyond the photocurrent initially generated. Although APDs generally provide better sensitivity, they generally have higher noise levels than PIN photodiodes. Recent developments have given rise to single-photon avalanche photodiodes (SPADs), which operate at very high reverse biases, often referred to as Geiger-mode. SPADs are optimized for single-photon detection, which triggers a self-sustaining avalanche breakdown that must be extinguished during operation. Due to this self-sustaining breakdown, SPADs cannot provide light intensity information that is different from APDs or PIN photodiodes. However, SPADs can be configured in parallel with a large number of other SPADs, collectively referred to as silicon photomultiplier (SiPM) photosensors. The SiPM produces a unique photosensor that generates a signal proportional to the light intensity produced by the corresponding number of triggered SPAD cells. In this way, the SiPM acts as an analog photon counter, whereby each individually triggered SPAD contributes to the total photocurrent output of the sensor and produces a much higher gain than a single APD or PIN photodiode.

[0053] SiPM photosensors, which can include tens, hundreds, or even thousands of single-photon avalanche photodiodes per square millimeter on a single silicon substrate. Silicon photomultipliers can be used when timing resolution and photon dynamic range that are useful in a particular application or use case are needed, desired, and / or otherwise determined. Advantages of such devices over photomultipliers include their reduced package size, lower bias voltage, reduced sensitivity to magnetic fields, and improved single-photon timing resolution. Disadvantages of silicon photomultipliers over traditional photomultipliers include: higher dark current or counts for a given temperature and area, and complex multi-exponential pulse shapes. Solid-state photosensors can replace photomultipliers within a block on a one-to-one basis, or can be coupled one-to-one to each elongated scintillator crystal within the block.

[0054] As used herein, reference to a "photosensor," "photodiode," and / or "photodetector" is not meant to limit or exclude any particular embodiment, unless expressly stated otherwise. As used herein, the term "readout" refers to the measurement, investigation, evaluation, or processing of the referenced subject matter. Similarly, the use of "lateral" readout, "side" readout, and "side" sensing all generally refer to light sensing by a scintillator, wherein the sensing face of the photosensor is coupled to a scintillator face parallel to the transverse axis of the scanner aperture and, therefore, is not acquired from either end face parallel to the axis of the scanner aperture. In addition, as used herein, "scintillator," "crystal," "pixel," and "block" may each refer to scintillator material configured in different forms.

[0055] Figure 4 is a block diagram of a PET scanner 400 having a ring of detector modules 401 and a plurality of lines of response 471, 472, and 473, according to one embodiment. Positron annihilation 461 from the center of the ring causes the lines of response 471 to strike orthogonally the end faces of the elongated scintillator crystals of a pair of detector modules. Positron annihilations occurring off-center, such as positron annihilations 462 and 463, cause gamma photons to strike the scintillator at oblique angles, thereby producing a parallax effect. As shown in an exploded view of the elongated scintillator crystal 410 of the detector module 401 (in the upper right corner of the figure), two gamma events 481 and 482 scintillate within the same elongated scintillator crystal 410, but originate from different locations in the patient's body (at positron annihilations 462 and 463, respectively). The end face photosensor 450 on the proximal end face of the elongated scintillator crystal 410 detects scintillation photons generated by the elongated scintillator crystal 410 in response to received gamma photons.

[0056] The true response lines 472 and 473 for annihilation events 462 and 463 are shown. However, the two gamma scintillation events 481 and 482 within the same crystal 410 may be processed such that two different positron annihilation events 462 and 463 are detected as originating from the same location, resulting in imaging errors. This parallax effect results in geometric uncertainty, including the further possibility of random events, because some embodiments of the detector module 401 cannot distinguish between events perpendicular to the surface of the elongated scintillator crystal 451 and events caused by severe tilt angles. The location of the end-face photosensor 450 on the proximal end face of the elongated scintillator crystal 451 allows the end-face photosensor 450 to detect scintillation photons generated within the elongated scintillator crystal 451, but the interaction depth cannot be determined. For example, the end-face photosensor 450 does not distinguish between scintillation events 481 and 482.

[0057] Figure 5AA block diagram of a detector module ring 500 is shown, according to one embodiment, which includes detector modules 510 and 513, having a single off-center line of response 530 from a single annihilation event 520. Scintillation events within detector modules 510 and 513 are used to calculate a line of response 540 ("computed line of response") that corresponds to the true line of response 530 but is offset by parallax error, as described herein. Nevertheless, the detection of scintillation events within detector modules 510 and 513 represents a "true" coincidence event, with a pair of detected gamma photons originating from a single annihilation 520.

[0058] Figure 5B A detector module ring 500 is shown with two eccentric response lines 531 and 532 based on multiple annihilation events 521 and 522, according to one embodiment. In the example shown, annihilation event 522 produces a reverse gamma photon, one of which is absorbed and detected by detector module 511, while the reverse corresponding gamma photon escapes the detector ring by virtue of its trajectory outside the ring. Similarly, annihilation event 521 produces a reverse gamma photon; however, only one gamma photon from annihilation event 521 is detected (e.g., by detector module 514). If detector modules 511 and 514 detect a pair of scintillation photons separated from the gamma photons produced by annihilation events 521 and 522 simultaneously or within a few nanoseconds of each other, the PET scanner may erroneously calculate an artefactual response line 541, which is referred to as a "random" event. That is, the PET scanner may incorrectly estimate that the gamma photons detected by detector module 511 and the gamma photons detected by detector module 514 originate from the same annihilation event. The final response lines calculated by the PET scanner are "false" because the calculated response lines do not correspond to true or reliable response lines, where the synchronization cues are actually random events rather than true synchronization events.

[0059] Figure 6 A ring 600 of detector modules is shown with response lines 630 segmented according to their temporal and spatial relationships, according to one embodiment. The response lines 630 correspond to positron annihilations 620 detected by opposing detector modules 610 and 612. According to various embodiments described herein, the improved optoelectronic sensor technology, configuration, placement, and / or high-speed electronics described herein can help improve the measurement of the arrival time difference between two detected gamma photons. This approach, known in the art as time of flight (ToF), enhances the ability to calculate the origin point of the positron annihilations 620 along the response lines 630.

[0060] For example, a timing resolution of 500 picoseconds enables the origin of positron annihilation 620 along a line of response 630 to be localized to within approximately + / - 7.5 centimeters, thereby dividing the scanner's 70 centimeter field of view into, for example, 4.68 sections per line of response (or more than ±9 sections). The improvement in annihilation event localization reduces uncertainty or noise in image reconstruction, thereby improving the scanner's signal-to-noise ratio. This increase in signal-to-noise ratio is essentially achieved as an increase in sensitivity compared to non-ToF scanners, where the sensitivity gain achieved is proportional to the square of the improvement in signal-to-noise ratio. According to various embodiments, a PET scanner can have a timing resolution of approximately 210 picoseconds.

[0061] Various embodiments of the presently described systems and methods facilitate more accurate and precise acquisition of PET data. Examples of systems and methods are described herein that reduce and thereby improve timing resolution; improve positional delineation of response lines for a given positron annihilation event to reduce parallax; increase sensitivity to provide better image clarity, reduce scan duration, and / or reduce radiation imaging; and / or combinations thereof.

[0062] According to various embodiments, the PET scanner system can reduce such errors by longitudinally segmenting the detector crystal (e.g., into two or more crystal slices) and / or incorporating two distinct crystal time constants. For example, using two different time constants allows the PET scanner system to use the shape of the photodetector signal to discern which crystal absorbed the energy of an incident gamma photon. In this way, the PET scanner system can determine the depth of interaction, referred to as the depth of interaction (DoI), of an incident gamma photon within the length segment of each elongated scintillator crystal. This depth of interaction improves positional resolution and reduces noise in the reconstructed image.

[0063] In some embodiments, the PET system may include a plurality of detector modules, each of which includes a two-dimensional array of elongated scintillating crystals that are longitudinally segmented into two or more sub-crystals having two distinct crystal time constants. A photodetector array associated with the elongated scintillating crystal detects scintillation photons originating from scintillation events within one of the segmented volumes. The shape of the signal generated by the photosensor array will differ depending on the crystal time constant of the segmented volume from which the scintillation photons originate. Therefore, the system can determine interaction depth information (i.e., which segmented volume the scintillation light originates from) based on the shape of the signal generated by the photodetector array. As previously described, interaction depth information can be used to improve position resolution and reduce noise in reconstructed images. In embodiments where each longitudinal segmented volume in a particular scintillating crystal has a different (i.e., unique, unique and identifiable) crystal time constant, reflective material may not be positioned between adjacent segmented volumes.

[0064] According to various embodiments, the PET scanner includes one or more detector module rings. Each detector module ring includes a plurality of detector modules. Each detector module includes an array of elongated scintillation crystals (e.g., a one-dimensional or two-dimensional array of elongated scintillation crystals) positioned on the detector ring. The distal face of each elongated scintillation crystal is radially oriented into the detector ring to receive gamma photons from annihilation events. Each elongated scintillation crystal may also include two axially oriented sides, two transversely oriented sides, and a proximal face. Each elongated scintillation crystal may include an array of photosensors (e.g., a single photodetector or an array of photodetectors) positioned along one of the axially oriented sides. The photosensor array detects scintillation photons from scintillation events within a corresponding scintillation event.

[0065] The reflective material is positioned (e.g., as a layer or coating) on ​​the proximal face, the distal face, two opposing transverse-axis oriented sides, and another axially oriented side (i.e., the axially oriented side without the photosensor array) of each elongated scintillator crystal. The reflective material functions to internally reflect scintillation photons such that all or substantially all scintillation photons from scintillation events within a given elongated scintillator crystal are detected by the photosensor array associated with the given elongated scintillator crystal. Thus, the reflective material functions to prevent light sharing between adjacent scintillator crystals. In some embodiments, the reflective material (e.g., coating, layer, film, etc.) can reflect in two directions such that a single reflective material layer or coating can be shared by adjacent scintillator crystals.

[0066] In various embodiments, each elongated scintillator crystal may be shaped as an elongated N-sided polygonal prism, such as a hexagonal prism, an elongated rectangular prism, a square prism, an octagonal prism, a triangular prism, etc. The length of the elongated scintillator crystal may be, for example, between 10 mm and 30 mm, and the width in the axial and transverse directions may be, for example, between 2 mm and 10 mm. In various examples, the thickness of the photosensor array may be less than 500 microns, and the thickness of the reflective material (e.g., coating) between adjacent elongated scintillator crystals may be less than 100 microns.

[0067] As described herein, each elongated scintillator crystal in a detector module can be configured for lateral readout by a photosensor array positioned on one of the axially-oriented sides. In some embodiments, the photosensor array comprises a two-dimensional array of single photodiodes, such as a two-dimensional array of single photon avalanche diodes in silicon photomultipliers. The two-dimensional array of single photodiodes on the axially-oriented sides can be used to provide different photosensor measurements corresponding to different depth ranges between the proximal face and the distal face of each elongated scintillator crystal. In some embodiments, one of the axially-oriented sides of each elongated scintillator crystal in the detector module can include a plurality of silicon photomultipliers positioned along the length of the axially-oriented side.

[0068] In various examples, the PET scanning system also includes a cooling system to cool the detector module. The PET scanning system may also include or communicate with an imaging system connected to the detector module to generate images based on the electronic output from the detector module. The diameter of the detector ring or hole of the PET scanning system may be greater than about 25 centimeters. In many embodiments, the axial hole depth of the plurality of axially aligned rings is less than the diameter of the detector ring.

[0069] Many existing computing devices and infrastructures can be used in combination with the currently described systems and methods. Some infrastructures that can be used for the embodiments disclosed herein are already available, such as processors, microprocessors, microcontrollers, computer programming tools and techniques, digital storage media, image processor devices, imaging processing techniques and communication links. Many systems, subsystems, modules, components, etc. described herein can improve timing resolution, improve the position depiction of a given positron response line to reduce parallax errors, improve sensitivity to provide better image clarity, reduce scanning duration and / or reduce patient radiation exposure. As understood by those skilled in the art and the context of their related descriptions, systems, subsystems, modules and components can be implemented as hardware, firmware and / or software. Because there are so many possible implementations, various systems, subsystems, modules and components are described according to the functions they perform. For example, it is understood that many existing programming languages, hardware devices, frequency bands, circuits, software platforms, network infrastructures and / or data storage devices can be used alone or in combination to implement specific functions.

[0070] It should also be understood that two or more systems, subsystems, components, modules, etc. described herein can be combined into a single system, subsystem, module, or component. In addition, many systems, subsystems, components, and modules can be repeated or further divided into discrete systems, subsystems, components, or modules to perform the subtasks described herein. Any embodiment described herein can be combined with any combination of other embodiments described herein.

[0071] The figures herein describe and illustrate the components of some disclosed embodiments. Many of the parts may be arranged and designed into various configurations. In addition, the features, structures and operations associated with one embodiment may be applied to or combined with the features, structures or operations described in conjunction with another embodiment. In many cases, well-known structures, materials or operations are not shown or described in detail to avoid blurring the aspects of the present disclosure. Many illustrations are provided in block diagram format to illustrate the general configuration and may not be drawn to scale. The right to add any of the embodiments or features to any one of the figures and / or as a new figure is expressly reserved.

[0072] In some examples of the presently described systems and methods, each photosensor array includes one or more (e.g., a two-dimensional array of hundreds or thousands) high-speed diodes for digitally encoding a large number of nuclear scintillation events. A network of circuit components converts the output of the photosensors into two common digital streams. The first digital stream can encode high-frequency content for time measurement. The second digital stream can encode a lower bandwidth signal version for numerical integration, measuring nuclear energy (e.g., calculating total energy). The circuit can be used for nuclear pulse processing in PET or other applications that utilize digital signal processing of large numbers of photon signals, requiring high timing resolution for time-of-flight measurements, such as light detection and ranging (LIDAR).

[0073] In some examples, the photosensor array can provide a capacitively decoupled high frequency signal output (referred to as a "fast signal" output) and a "slow signal" output that includes a fuller but lower frequency spectrum. In other examples, the photosensor array can employ only two-terminal photosensors, where fast timing and / or energy measurements can be obtained at the same or various nodes within the bias network of resistors, capacitors, inductors, and / or active components. In various embodiments, the low noise analog circuit detects a low voltage threshold crossing of the fast signal via a comparator so that the signal can propagate to a high speed analog-to-digital converter (ADC) for digital encoding. In some embodiments, the low noise analog circuit detects a low current or low voltage threshold crossing of the fast signal via a comparator so that the signal can propagate to a high speed analog-to-digital converter for digital encoding.

[0074] Additionally, comparators may be used to select multiplexer inputs to enable separate digital encoding of slow signals (or lower bandwidth). Many of the described examples provide improvements in component propagation delays and their mutual impact on bandwidth to successfully share digital encoding of a large number of front-end photodiodes that are part of a photosensor array of one or more scintillator crystals in a detector module of a PET scanner system.

[0075] Optimized transmission lines are used to achieve target propagation delays. Many embodiments of the presently described systems and methods achieve better determination of the timing origin of nuclear scintillation events by digitally eliminating traditional analog time walks and through spectral distortion correction and machine learning algorithms. In addition, the systems and methods described herein eliminate or avoid the use of application-specific analog electronics for timing picking and provide low-noise sharing of high-speed electronics, which saves power consumption, heat dissipation, packaging volume, and cost compared to traditional approaches. Additional embodiments provide for effective thermal control or cooling of scintillator materials and photosensors, which can improve the signal-to-noise ratio. Such improvements in the timing of nuclear scintillation events have far-reaching clinical implications, for example, by reducing image noise, thereby increasing sensitivity and improving diagnostic clarity.

[0076] Figure 7 700 is a simplified schematic diagram of a high speed diode circuit network 700 according to one embodiment. The diode circuit network 700 for light sensing operates to provide high bandwidth digital encoding of nuclear scintillation events. One or more photodiodes 708 (illustrated as PD1-PDn) can be reverse biased to operate in Geiger mode. The highest bandwidth output of each photodiode 708 is capacitively decoupled and connected to a comparator 709 and a fast transmission line 710. The lower bandwidth connections of the photodiodes 708 are connected to a slow transmission line 714 for selective encoding by an analog-to-digital converter 707 through a multiplexer 706.

[0077] Comparator 709 within switch circuit block 703 compares the signal from the higher bandwidth signal to a low voltage threshold 715 Vth (which may alternatively be a low current threshold). When the signal from the higher bandwidth channel exceeds Vth, comparator 709 activates flip-flop 711. The output of flip-flop 711 biases output diode 713 "on", thereby propagating the high bandwidth photodiode signal from fast transmission line 710 through a capacitor (labeled Cdcb) for high sampling speed digital encoding by analog-to-digital converter 705 and to ground through an inductor (labeled Lr).

[0078] Circuit blocks 701 and 703 are replicated in blocks 702 and 704, respectively. Output diode 713 of circuit blocks 701 and 703 is replicated as output diode 723 of blocks 702 and 704. Circuit blocks 701 and 703 may be replicated any number of times to allow sharing of analog-to-digital converter 705. Output diode 793 represents the nth output of n replicated circuit blocks 701 and 703 for any number of photodiode groups for any number of scintillation crystals in a detector module of a PET scanner system.

[0079] Digital control logic 716 ensures that only one high-speed signal propagates through any given output diode D1-Dn (713, 723-793) at a time. Control inputs 731 (Q1-Qn) identify to control logic 716 the particular group of photodiodes 708, 718 that are triggered, as well as their respective comparators 709, 719 and flip-flops 711, 721. Control outputs 732 (EN1-ENn) are normally "on" and signal the flip-flops 711, 721 to be disabled when any particular single diode is triggered and conducting forward. After the envelope of the photodiode's core pulse signal is complete (which can be predetermined by its characteristic decay constant), control logic 716 resets any previously triggered trigger switch circuit blocks 703, 704 via the CLR signal. The digital control logic 716 selectively pulses the clear outputs 733 (CLR1-CLRn) to clear any triggered flip-flops 711, 721 to their untriggered states, thereby turning off their corresponding output diodes D1-Dn (713, 723-793). Thereafter or thereafter, all control outputs 732 are restored to "on" to provide a trigger ready state for any subsequent qualified nuclear pulse signal.

[0080] In addition, the output bus 717 of the control logic 716 is used to select the lower bandwidth output of a particular photodiode based on the trigger switch circuit input 731 (Q1-Qn) for propagation through the multiplexer 706. As shown, the slow signal propagates through the slow transmission line 714 and is received as the "energy signal" input on the multiplexer 706. The energy signal selected by the output bus 717 is sent from the multiplexer 706 to the analog-to-digital converter 707 for digital encoding. The analog-to-digital converter 707 can be a relatively low sampling speed analog-to-digital converter.

[0081] Diode network 700 digitally encodes the output of multiple groups of photodiodes (e.g., a photodiode network), which are generally represented by circuit blocks 701 and 702. Analog-to-digital converters 705 and 707 can be used to digitally encode any number of photodiode groups, where each photodiode array includes at least one photodiode. As used herein, a photosensor can include any number of photodiodes.

[0082] In some embodiments, the photodiode network may be comprised of a plurality of photodiodes that share a common high bandwidth output and a relatively low bandwidth output that are propagated via fast transmission line 710 and slow transmission line 714, respectively. The corresponding trigger switch circuit blocks 703 and 704 propagate the high speed nuclear pulses in a mutually exclusive manner for digital encoding via analog-to-digital converter 705 to provide a highly encoded signal that can be used to determine the time of origin of the nuclear pulse, while the multiplexer 706 and analog-to-digital converter 707 provide a lower bandwidth encoding for energy discrimination of the same nuclear pulse. The higher bandwidth signal may be generally referred to herein as a "fast" signal or "timing" signal, while the lower bandwidth signal may be referred to herein as a "slow" signal or "energy" signal.

[0083] The fast transmission line 710 is labeled Tx3 (or Tx4 in the corresponding block 704), which is configured to enable the comparator 709 and the trigger switch components (e.g., trigger 711, inductor 712, and output diode 713) to provide sufficient delay to conduct charge when the upcoming nuclear pulse signal (defined by threshold 715 and comparator 709) arrives at the input of the output diode 713. Similarly, the slow transmission line 714 is labeled Tx1 (or Tx2 in the corresponding block 702), which is configured to provide sufficient delay to enable the trigger switch circuit block 703 (or 704) to activate the control logic 716 and the multiplexer 706 to conduct charge, thereby fully encoding the nuclear pulse signal through the analog-to-digital converter 707. The fast transmission line 710 and the slow transmission line 714 are configured to produce sufficient signal delay to achieve the encoding of the entire nuclear scintillation pulse envelope. The fast transmission line 710 and the slow transmission line 714 are also selected to have sufficient spectral bandwidth for their respective signal content.

[0084] Fig. 8A 800 is a two-port network lumped element model of a conductive transmission line according to one embodiment. The lumped element 805 of the transmission line is modeled by discrete elements as described below. The resistive element 801 represents the series resistance of the transmission line. The inductive element 802 represents the series inductance or reactance of the transmission line. The resistive element 803 represents the shunt conductance of the transmission line, and the capacitive element 804 represents the shunt capacitance or admittance of the transmission line.

[0085] Figure 8B A distributed model 810 of a transmission line based on lumped elements 805A-805N per unit length is shown according to one embodiment. Each lumped element 805A-805N is composed of Fig. 8A805. A relatively short transmission line may be embodied as a single two-port lumped element 805. However, longer transmission lines are more accurately modeled as a distributed model based on lumped elements 805A-805N per unit length, where the values ​​of the discrete elements in each lumped element 805A-805N may be the same or different from each other, as described herein and understood in the art.

[0086] According to Oliver Heaviside's solution of the telegraph equation from James Maxwell's equations, the transfer function H(ω,x) (often called the propagation function) of a transmission line is given by:

[0087]

[0088] in: x = length of transmission line L = series inductance R = Series Resistance C = Parallel Capacitance G = parallel conductance And where R, L, G and C are on a per unit length basis.

[0089] The propagation coefficient γ(ω) is obtained by taking the negative natural logarithm of the propagation function, yielding:

[0090]

[0091] where the real components of γ represent the attenuation per unit length (in nepers) and the imaginary components represent the phase delay per unit length (in radians). For frequencies where the impedance or admittance of a reactive component greatly exceeds its respective passive resistance or conductance, the propagation function is approximated by:

[0092]

[0093] which is a linear phase delay where the propagation velocity per unit is given by its inverse:

[0094]

[0095] It can be equated to the velocity per unit due to the material:

[0096]

[0097] in: ∈ r = relative dielectric constant μ r =Relative magnetic permeability And the propagation delay is given by the reciprocal of the signal velocity of the line.

[0098] The impedance of a transmission line is given by the square root of the ratio of the series line resistance and impedance to the parallel conductance and admittance:

[0099]

[0100] And, for frequencies where the per-unit passive impedance and admittance significantly exceed the per-unit passive resistance and conductance, the characteristic impedance simply becomes the square root of the ratio of inductance to capacitance for all frequencies beyond this, and thus appears theoretically purely resistive.

[0101] For a conductive transmission line, if the total propagation delay of the line is much shorter than the rise or fall time of its input signal, this mode of operation is generally referred to as the lumped element region and can be modeled with a lumped element model 800 of a two-port network. The lumped element region can be considered when the following condition is met:

[0102] |x·γ(ω)| < 1 / 4

[0103] where x is the length of the transmission line and γ (gamma) is the propagation coefficient.

[0104] The transmission line operates at lengths beyond the lumped element boundary but at relatively low frequencies (e.g., < 1 MHz;), entering the RC region. In this operating region, the inductive reactance can be neglected, but the series resistance is still significant, especially when considered in combination with the total capacitance of the line. In this mode, the bandwidth of the line is basically determined by a single-pole response, limiting the signal bandwidth to:

[0105]

[0106] where f 3db represents the corner of the low-pass frequency response.

[0107] Beyond this frequency region, and for lengths greater than the lumped element region, the transmission line operates in the LC region (e.g., 1 MHz < f < 10 MHz), where the series inductive reactance dominates the resistance of the line and the capacitive admittance dominates the parallel conductance of the line. From a more practical perspective, non-linear effects occur at frequencies outside the LC region, including the alternating current (AC) skin effect, which attenuates the signal of the line and limits the bandwidth. Additionally, for common materials and structures, and at frequencies of approximately 1 GHz and above, dielectric losses become very significant.

[0108] The propagation delay per unit of a conductive transmission line is determined by the square root of the product of the inductance and capacitance per unit. The longer the line, the greater the delay. Ideally, at least in theory, a properly terminated transmission line appears to the signal transmitter to be merely resistive at medium frequencies and above. In practice, losses occur at higher frequencies due to skin effect and dielectric losses. These losses effectively limit the bandwidth and cause group delay. Therefore, it can be understood that the conductive transmission line delay and bandwidth are inversely proportional. This inverse relationship poses a challenge to the design and implementation of fast transmission line 710, where a delay of more than two nanoseconds is required to withstand the subsequent propagation delay caused by the implementation of elements 709, 711, 712, and 713.

[0109] In the case of a qualified propagation signal (for example, when the photoelectric sensor signal exceeds the threshold V th , 715) from the fast transmission line 710 (or another fast delay element) to the input of the output diode 713, the output diode 713 is activated. At the same time, the bandwidth of the fast transmission line 710 (or another fast delay element) is high enough to preserve the fast rising, high bandwidth signal from the photosensor 708, which is necessary for timing resolution. Similarly, the slow transmission line 714 has sufficient bandwidth while also providing sufficient propagation delay to allow the trigger switch circuit block 703, control logic 716, and multiplexer 706 to activate and enable conduction in time so that the corresponding energy signal is fully digitally encoded by the analog-to-digital converter 707.

[0110] Fig. 9 A time domain diagram 900 of two nuclear pulse signals according to one embodiment is shown. The illustrated silicon photomultiplier tube photosensor signals 901 and 902 are both from the deposition of nuclear particles into a coupled scintillator. Signal 901 is a typical capacitively decoupled (i.e., high pass filtered) silicon photomultiplier tube avalanche diode signal, referred to herein as the fast signal or timing signal of the detector module. Signal 902 represents a low pass filtered version of the same avalanche diode signal, referred to herein as the slow signal or energy signal of the detector module. Assume that the decay constant of the slow signal 902 is much slower than the decay constant of the elongated scintillator crystal (τ low-pass >5·τ scint ), then the maximum peak value of signal 902 will be approximately the total charge Q collected from the photosensor divided by the capacitance C of the circuit. In this case, the total charge collected will be proportional to the energy deposited into the scintillator. In alternative embodiments, a circuit or processing device can be used to integrate the unfiltered photosensor signal to obtain a proportional energy measurement.

[0111] The detector module can use fast timing signal 901 to determine the timing of the nuclear scintillation origin event (also referred to herein as the original or initial). Fast signal 901 and slow signal 902 are reproduced in more detail by signals 903 and 904, respectively, showing white or broadband thermal noise thereon, which has an approximately Gaussian amplitude distribution. This and other noise sources introduce timing uncertainty when using voltage discrimination, as shown by voltage level threshold 907. Using voltage level threshold 907 to identify the timing of the signal is called "timing pick".

[0112] Using voltage discrimination on signal 904 (corresponding to slow signal 902) results in timing uncertainty. The Gaussian noise of slow signal 902 results in a Gaussian timing jitter histogram, where the timing uncertainty is plotted as 906. This effect is obvious because the timing pick-up of a slow rising signal (i.e., a shallow slope) is more susceptible to the vertical amplitude of the noise, referred to herein as the slope-to-noise relationship. However, fast signal 903 (corresponding to fast signal 901) is less susceptible because it is a fast rising signal (i.e., a steep slope). Comparing fast signal 903 to voltage level threshold 907 (e.g., threshold) results in a narrower timing distribution, plotted as 905. Therefore, the effect of noise on the timing resolution of fast signals 901 and 903 is reduced or better than the effect of noise on the timing resolution of slow signals 902 and 904.

[0113] Therefore, various embodiments of the systems and methods currently described utilize timing circuits to analyze the fast capacitance decoupling output signals of one or more photodiodes to determine the timing information (e.g., timing pick-up values) of scintillation events. Similarly, various embodiments of the systems and methods currently described utilize separate energy circuits to determine the energy information (e.g., total energy values) of scintillation events. It will be appreciated that some circuit components may be shared between timing circuits and energy circuits. In some embodiments, each elongated scintillation crystal in a detector module may be associated with a dedicated energy circuit and a dedicated timing circuit, including a dedicated analog-to-digital converter. In other embodiments, multiple scintillation crystals may share energy circuits and timing circuits and / or share analog-to-digital converters. Each energy circuit and timing circuit may include multiple discrete electronic component channels, wherein each channel is dedicated to one or more scintillation crystals, and wherein these channels share the same analog-to-digital converter.

[0114] Fig.10 According to one embodiment, Fig. 9 Energy spectrum density diagram 1000 of two nuclear pulse signals in the frequency domain. Fig. 9 The energy spectral density of signal 901 in is shown by graph 1010 . Fig. 9Signal 902 in FIG. 1 is shown by plot 1020. It can be seen that the spectrum of slow signal 902 generally cuts off at around 400 MHz, with trend line 1040 being the approximate level of the noise floor beyond that point. However, for signal 901, the corresponding high frequency spectrum of plot 1010 is more pronounced, with trend line 1030 extending to approximately 1.2 GHz before approaching the noise floor. For a single pole low pass filter circuit, the 10-90% rise time of the signal acting on the circuit can be found by the following relationship and shown to be proportional to the time constant of the circuit:

[0116] in: t r = 10 to 90% rise time of the step signal, f 3db = 3db bandwidth (BW) of the circuit.

[0117] In various embodiments, high frequency content is preserved to produce a steeper timing signal than would otherwise be achievable. Some embodiments of the system optimally preserve wide bandwidth timing channels to improve timing pick-off and reduce timing jitter (e.g., uncertainty) taking into account slope-to-noise ratio and slope limitations of timing signals due to transmission line bandwidth.

[0118] Therefore, return Figure 7 , the fast transmission line 710 may be selected to have the minimum or shortest practical length to reduce high frequency attenuation while still creating enough delay to enable the remaining circuit elements (e.g., comparator 709, flip-flop 711, inductor 712, and diode 713) in the box of block 703 to activate in time to encode the capacitively decoupled "fast signal" from the photodiode 708 through the analog-to-digital converter 705. Therefore, the propagation delays of the comparator 709, flip-flop 711, bias-tee inductor 712, and output diode 713 are selected to operate as fast as possible to facilitate shorter propagation delays in the fast transmission line 710 (e.g., the fast transmission line 710 is physically shorter). For example, the bias inductor 712 may be designed for high speed switching, utilize a low loss, low relative permeability magnetic core (e.g., air core), and exhibit, for example, a low inductance (e.g., less than 100 nanohenries) to facilitate fast biasing of the diode 713. In contrast, the slowly rising energy signal from the photodiode 708 may take the longer slow transmission line 714, resulting in a greater propagation delay because the high frequency content is much lower than the high frequency content of the faster timing signal. For example, the delay of the slow transmission line 714 may be more than two times longer than the delay of the fast transmission line 710. In some embodiments, the delay of the slow transmission line 714 may be five times longer than the delay of the fast transmission line 710, or even longer.

[0119] Given the importance of the high frequency spectrum of the timing signal, special high frequency, fast switching diodes (e.g., PIN diodes using an intrinsic semiconductor region with a high level of implantation) can be used to reduce the delay required for the fast transmission line 710. For small signals at high frequencies, the forward biased PIN diode advantageously behaves as a current-controlled variable resistor, providing a resistance of less than 100 ohms at 1 GHz, and a resistance of even less than ten ohms for small currents (e.g., about ten milliamps) or in some cases. A relatively large cross-sectional area intrinsic area reduces the forward conduction (i.e., "on") resistance of the PIN diode, thereby improving (i.e., reducing) insertion loss, but undesirably increases the junction capacitance in the reverse bias state (i.e., "off"). Increasing the thickness of the intrinsic region between adjacent P and N regions reduces the capacitance while also increasing the reverse bias breakdown, both of which are beneficial to improving isolation in the "off" state. However, a thicker intrinsic region undesirably increases the switching time, requiring a correspondingly longer delay for the fast transmission line 710. In various embodiments, low junction capacitance improves operation of circuit blocks 703 and 704 in an untriggered state (e.g., to achieve optimal operation) so as to isolate photosensor noise from circuit blocks 701 and 702 from a common node present at the input of analog-to-digital converter 705.

[0120] According to various embodiments, these tradeoffs are considered and carefully accounted for in the design and selection of the fast transmission line 710 (or other delay element) and the output diode 713 to achieve acceptable timing resolution, or even optimized timing resolution. The use of high energy bandgap semiconductors with energies greater than 1.3 eV (at 3000 K), such as those provided by gallium arsenide, can increase the reverse bias breakdown voltage of silicon, thereby allowing thinner intrinsic regions to provide faster switching times. In addition, gallium arsenide provides faster carrier lifetimes, typically less than 10 nanoseconds at small currents (e.g., 1–20 mA), and in some cases, typically faster than 5 nanoseconds, thereby enabling faster removal of charge from the intrinsic region.

[0121] In some cases, the "on" (i.e., forward conduction mode) time of a PIN diode may be faster than the "off" (i.e., reverse bias configuration) time due to differences in the injection and removal of charge carriers. In some embodiments, this characteristic of a PIN diode is acceptable given the unique nature of PET signals. This slow turn-off time tolerance is provided by the relatively infrequent nature of PET single events on a typical block basis, whereby each event typically requires a long duration to decay (e.g., approximately 200 nanoseconds for lutetium oxysilicate-based scintillators), so that in this unique situation (i.e., fast turn-on and slow turn-off of PET single events), several orders of magnitude of turn-off duration relative to turn-on can be tolerated. Desirable characteristics of a PIN diode used to implement output diode 713 include low insertion loss (e.g., <1 dB) and high turn-off isolation (e.g., >15 dB), which characteristics are typically present in the frequencies of gamma scintillation photodiode signals, such as Fig.10 The frequencies shown can be up to several gigahertz (e.g., 4 GHz) or higher. High off isolation is achieved using relatively low junction capacitance (e.g., less than one hundred fifty (150) femtofarads) at these frequencies. The cathode of the diode can optionally be negatively offset to lower the threshold voltage that causes the diode to be in the on state (e.g., forward activated / conducting). Fast conducting, low resistance, and ultra-low capacitance gallium arsenide (GaAs) or aluminum gallium arsenide (AlGaAs) diodes can be used to achieve the target fast switching time for the target bandwidth of transmission line 710. The target rise time of the fast signal of photodiode 708 can be, for example, less than one nanosecond at the expected PET scanner noise level to achieve precise timing. The target bandwidth of transmission line 710 can be, for example, greater than 350 MHz over its entire length.

[0122] Analog-to-digital converters 705 and 707 may sample at or above the Nyquist rate (i.e., half the highest bandwidth period of the corresponding input signal) to encode timing and energy signals, respectively. The sampling rate may be selected based on the Nyquist-Shannon sampling theorem. Alternatively, the conversion rate of the analog-to-digital converter may be selected based on the propagation of uncertainty to achieve optimal timing pick-off, where one or more data points of the signal attenuation curve may be utilized to minimize the propagation of uncertainty, as given by:

[0123] In this case, each data point has an uncertainty σ, which is also related to the other data points by a correlation coefficient ρ. The system can use this equation to reduce or minimize the uncertainty of the timing pick. The correlation coefficient can also be given by the instrument response function.

[0124] The specific components and circuit designs provided herein are merely examples of possible implementations. Those skilled in the art will appreciate that the general concept can be implemented by other circuits, components, and techniques to achieve the same or similar results. For example, the slow energy analog-to-digital converter 707 can be eliminated, and the signal from the high-speed analog-to-digital converter 705 can also be digitally integrated to determine the total energy of the signal. In addition, the trigger circuit 703 can be implemented using alternative components, such as using a two-stage comparator with an RC delay circuit in between, because the time constant of the scintillator and the signal envelope are uniquely known. In addition, the switch element 712 (shown as an inductor element) and the output diode 713 can be implemented with a transistor (e.g., with a common base or a common gate) instead and operated in an open drain configuration. Even further, the fast transmission line 710 and the slow transmission line 714 can be implemented optically using optical fibers or other light guides, respectively. The photodiode 708 can also be configured in a different manner, wherein the pull-up resistor is connected to the cathode and the fast signal is decoupled from its capacitance, while still optionally retaining the anode resistance for low-frequency energy signal detection. As a further alternative, the anode of the photodiode can be directly connected to the input of a low input impedance transimpedance amplifier.

[0125] Fig.11 1 is a hierarchical block diagram of a high speed diode network 1100 according to one embodiment. As described herein, a gamma radiation detector module may include a photosensor array associated with each elongated scintillator crystal. The photosensor array may include any number of photodiodes, such as arranged in a photodiode array (e.g., one or more single photon avalanche photodiodes or silicon photomultipliers). The illustration includes two electronic processing channels 1110 and 1120. The first electronic processing channel 1110 is associated with a first photodiode array, including a first photodiode array 1101 and a second photodiode array 1102 as shown. The second electronic processing channel 1120 is associated with a second photodiode array, including a third photodiode array 1111 and a fourth photodiode array 1112 as shown.

[0126] According to various embodiments, the processing circuit associated with the detector module includes two processing channels to ensure that signals from adjacent scintillation crystals can be processed simultaneously. However, it is understood that the processing circuit of the detector module can include any number of channels that share any number of analog-to-digital converters. Each channel in the N-channel processing circuit can be associated with any number of adjacent or non-adjacent scintillation crystals in the scintillation crystal block of the detector module.

[0127] In the illustrated example, the detector module may include a single high speed diode network 1100 having two electronic processing channels 1110 and 1120. Half of the elongated scintillator crystals in the detector module (e.g., every other one) may be associated with the first electronic processing channel 1110. More specifically, the photodiodes associated with half of the elongated scintillator crystals in the detector module are connected to the first electronic processing channel 1110 and share a single fast timing analog-to-digital converter 1105 and a single slow or energy analog-to-digital converter 1107.

[0128] Thus, the first electronic processing channel 1110 can be associated with any number of photodiode arrays, including the illustrated first photodiode array 1101 and the second photodiode array 1102. The photodiode arrays 1101 and 1102 can have their own trigger switch circuits 1103 and 1104 and / or share a single trigger switch circuit. The trigger switch circuits 1103 and 1104 (or the shared single trigger switch circuit) can be, for example, combined with Figure 7 Multiplexer 1106 is used to select between various photodiode arrays (including photodiode arrays 1101 and 1102 ) for encoding energy signals by energy analog-to-digital converter 1107 .

[0129] The photodiode arrays of the other half of the elongated scintillator crystals in the detector module can be associated with a second electronic processing channel 1120. These photodiode arrays share a single fast timing analog-to-digital converter 1115 and a single slow or energy analog-to-digital converter 1117. The second electronic processing channel 1120 can be associated with any number of photodiode arrays, including the third photodiode array 1111 and the fourth photodiode array 1112 shown. The photodiode arrays 1111 and 1112 can have their own trigger switch circuits 1113 and 1114 and / or share a single trigger switch circuit. The trigger switch circuits 1113 and 1114 (or share a single trigger switch circuit) can be, for example, combined with Figure 7 Multiplexer 1116 is used to select between various photodiode arrays (including photodiode arrays 1111 and 1112) for energy signal encoding by energy analog-to-digital converter 1117. Figure 7 As described above or variations thereof, digital control logic may be used to control diode biasing in multiplexers 1106 and 1116 and / or trigger switch circuits 1103 , 1104 , 1113 , and 1114 .

[0130] As described in more detail herein in conjunction with other figures, portions (or all) of the electronic processing channels 1110 and 1120 for a particular scintillation crystal in a detector module may be integrated within an interposer. The interposer may be positioned on the back side of the photoelectric sensor on the side of the elongated scintillation crystal. In some embodiments, each elongated scintillation crystal is associated with a different interposer. In other embodiments, multiple elongated scintillation crystals may share a single interposer. According to any of the various embodiments described herein, an interposer associated with one or more elongated scintillation crystals may, for example, include biasing and signal shaping circuits of circuit block 701.

[0131] In other embodiments, additional portions of the high-speed diode circuit network 700 may be incorporated into an inserter associated with one or more elongated scintillator crystals. According to any of the various embodiments described herein, the inserter may include and / or be connected to a timing circuit and an energy circuit. The energy circuit may, for example, include a first analog-to-digital converter, a multiplexer for selecting between the output signals of the photodiodes for conversion by the first analog-to-digital converter, a signal delay transmission line for connecting the photodiodes to the multiplexer, and digital control logic for controlling the multiplexer to select the output signal of the photodiode. The timing circuit may, for example, include a second analog-to-digital converter (for high-speed digital encoding of the output signal of each photodiode), a trigger or other triggering component (for selectively triggering the encoding of the second analog-to-digital converter), and a comparator for comparing the output signal of each photodiode with a threshold value.

[0132] Fig.12 A block of scintillation crystals 1200 arranged in a two-dimensional array according to one embodiment is shown. The illustrated view shows the distal faces of the elongated scintillation crystals. As shown in the expanded view of the upper left elongated scintillation crystal, each elongated scintillation crystal includes a reflective material as a layer or coating applied to the distal face 1230, the transversely oriented side faces 1231 and 1233, and the first axial face 1232. The reflective material is also positioned on the proximal face (not shown). In some embodiments, the reflective material is a thermally conductive reflective material, such as, but not limited to, one or more of aluminum, magnesium, silver, stainless steel, and / or combinations thereof.

[0133] The photosensor array 1201 (e.g., one or more photodiode arrays) is positioned along an opposite second axial face 1234 that is not coated with a reflective material. Thus, scintillation photons from a scintillation event within one of the elongated scintillator crystals are internally reflected within the elongated scintillator crystal until received by a laterally positioned photosensor. The reflective material prevents light sharing between adjacent crystals in the two-dimensional array of elongated scintillator crystals. The simplified illustration includes an array of 16 scintillator crystals. However, it will be appreciated that the two-dimensional array of elongated scintillator crystals may include M elongated scintillator crystals in the axial direction of the detector ring and N elongated scintillator crystals in the transverse direction of the detector ring, where M and N are both integer values.

[0134] The elongated scintillator crystals in the two-dimensional array are arranged seamlessly so that each elongated scintillator crystal, its reflective layer / coating and its photosensor are adjacent to each other. Alternating shading is used for adjacent elongated scintillator crystals to indicate the electronic processing channels associated with each elongated scintillator crystal in the two-dimensional array. In the illustrated embodiment, the elongated scintillator crystals shown with a dot fill pattern may utilize a first electronic processing channel (e.g., Fig.11 The elongated scintillator crystals shown with the cross-hatched fill pattern may utilize a second electronic processing channel (e.g., Fig.11 electronic processing channels 1120). In this configuration, inter-crystal scattered radiation of gamma photons can partially deposit energy into one crystal and then deposit its remaining energy into an adjacent crystal and can be processed without being discarded due to the dead time of a single processing channel per block.

[0135] While the reflective material may be thin relative to the size of each elongated scintillator crystal (e.g., as a thin sheet, deposited layer, applied film, coating, etc.), the thickness of the photodiode may be much thicker. This thickness of the reflective material and sensor consumes space between the crystals, thereby reducing the scintillator area (and volume) compared to the same geometric block detector without the lateral photodetector and reflective coating. The percentage of the scintillator area on the radiating face (the distal face shown) to the defined block area is referred to herein as the fill factor of the block. The lower the fill factor, the lower the sensitivity of the detector. Minimizing the thickness of the reflective material and / or the thickness of the laterally positioned (e.g., laterally mounted, laterally integrated, and / or otherwise laterally positioned) photosensors can achieve higher fill factors.

[0136] By coupling a photosensor (which may consist of a bare silicon sensor) to a scintillator crystal, using a micro lead frame package (also known as flat leadless) or wafer level package, optionally with through silicon vias, and electrically bonded to a flexible polyimide printed circuit board substrate or similar material, photosensor thicknesses of less than 400 μm can be achieved. The bonding or coupling of the photosensor to the side of the elongated scintillator crystal can be achieved by using an optical coupling agent or grease (such as silicone grease or clear epoxy resin), which helps to improve the efficiency of light transmission and can also be used as a glue to assist in manufacturing and / or increase durability. This coupling between the scintillator crystal, coupling agent material and / or photosensor can be photon optimized by selecting materials with a narrow range of refractive indices to maximize the critical angle of reflection, thereby providing efficient photon coupling with low reflection. For lutetium orthosilicate (LSO), the refractive index is about 1.82. For bismuth germanate (BGO), the refractive index is about 2.15, and for perovskites, the refractive index ranges from about 2.0 to 2.3.

[0137] The elongated scintillator crystal and / or the photosensor surface area may also be processed, such as by polishing or roughening, to enhance photon transmission. Bonding between the silicon package and the interposer and / or printed circuit board may be achieved using solder balls, conductive amalgam paste, conductive epoxy, ultrasonic welding, etc. For scintillator crystals of various sizes having an elongated rectangular prism shape, fill factors above 85% and even above 88% may be achieved using these techniques.

[0138] In some cases, semiconductor wafer support materials, such as glass, are used during the photosensor manufacturing process. The glass or other support material can be adhered to the photosensitive surface of the wafer. Such glass or other transparent or translucent material may also be favored for its optical properties used during photosensing. The glass or translucent material may be thinned to achieve an optimal (increased) fill factor, for example, where the transparent or translucent optical material may be thinner than the finished semiconductor photosensor. Alternatively, a scintillator material may be used instead of glass or other transparent or translucent material as a wafer support material.

[0139] PET has an optimum for scintillator crystals with side photosensors employing one or more side photodiodes. For example, scintillators such as BGO or LSO often exhibit a bi-exponential response, simulating the finite time required for the scintillator to initially fill the luminescent state, followed by fluorescence decay. The overall shape of the response can often be modeled as:

[0140]

[0141] The time constant T r corresponds to the generally fast luminescent state population, while T f Simulates the normally slower falloff of light.

[0142] The light output (e.g., expressed as the number of photons released, N) is proportional to the energy absorbed by the scintillator. For example, an LSO emits about 25,000 photons for every 1 megaelectronvolt (MeV) absorbed, and the decay time constant (T f ) is approximately 47 nanoseconds (ns). The statistical uncertainty or sample variance of the photon output is approximately equal to the square root of the number of measured photons. Since the bi-exponential response results in a specific peak time, increased light output results in a steeper signal leading edge (i.e., rise time) regardless of the amplitude. Correspondingly, the timing uncertainty or jitter is inversely proportional to the scintillator photosensor signal amplitude, as shown in graph 900, so:

[0144] Therefore, the timing jitter t jitter It decreases inversely proportional to the square root of the number of measured photons N.

[0145] For PET using time-of-flight, the reduction in temporal uncertainty reduces image noise, thereby increasing the image signal-to-noise ratio (SNR) by

[0146]

[0147] in: D = effective diameter of the imaged object, c = speed of light, and SNR non-ToF = Signal-to-noise ratio without time-of-flight application.

[0148] The SNR of a PET scanner is proportional to the square root of its noise effective counts (NEC). Therefore, the gain in SNR is comparable to the gain in simultaneous events. Therefore, as mentioned earlier, the sensitivity of a PET scanner is therefore improved in proportion to the square of the SNR improvement. However, a loss in fill factor reduces sensitivity. Since the PET image is made up of simultaneous response lines (i.e., from two detectors), any loss in fill factor is affected by the square of the fill factor. Thus, for example, a fill factor of 90% will result in a loss of PET sensitivity of 81%. Therefore, there is an optimal design for lateral readout relative to the volume displaced by the lateral photosensor.

[0149] In various embodiments, a plurality of detector module rings are used to form a detection cylinder. The plurality of detector module rings forming a cylinder can be described as having an axial depth d of the cylinder (e.g., a hole). In many embodiments, the axial depth is much smaller than the diameter D of the ring, especially for human body tomography scanners. The blocks of the detector modules are usually axially adjacent in sequence to define a certain number of rings in the entire cylinder depth.

[0150] Coupling side or lateral photodiodes to multiple crystals in a circumferential or transverse manner increases the diameter of the cylinder. Increasing the diameter opens the aperture, increasing sensitivity to radiation outside the field of view, thereby increasing random or scattered radiation, resulting in noise in the acquired image. In addition, increasing the aperture reduces the radiation acceptance angle within the scanner, or the angle of the lines of response produced along the axial length of the aperture, thereby reducing sensitivity. The geometric efficiency of a scanner is related to the solid angle coverage provided by its detector for a given point source in the field of view:

[0152] In contrast, the thickness of the photodiodes oriented laterally or laterally along the axial length of the aperture (i.e., along the axially oriented side of each elongated scintillator crystal) extends the depth of the aperture but does not affect the cylindrical diameter. Elongating the aperture with inactive material (i.e., non-scintillator material) reduces sensitivity due to a reduction in fill factor. However, this reduction in fill factor is the same regardless of whether the photosensors are positioned on the axially oriented sides of the elongated scintillator crystals or the transversely oriented sides of the elongated scintillator crystals. Although there is a loss in sensitivity based on the fill factor per unit length, an increase in the length of the aperture does increase sensitivity due to an increase in the acceptance angle. Therefore, positioning the photosensors on the axially oriented faces of the elongated scintillator crystals reduces the noise of random or scattered radiation and improves sensitivity through an improved acceptance angle, compared to positioning the photodiodes on the transversely oriented faces (e.g., along the circumferential sequence of the apertures).

[0153] For reference, elongated scintillator crystal 1250 is axially adjacent to elongated scintillator crystal 1251 and axially adjacent to elongated scintillator crystal 1253 (i.e., adjacent in the axial direction). Conversely, elongated scintillator crystal 1250 is transversely adjacent to elongated scintillator crystal 1252 and transversely adjacent to elongated scintillator crystal 1254 (i.e., adjacent in the transverse direction).

[0154] Fig.13Ais a diagram of a single scintillator crystal pixel 1300 with a photosensor array 1301-1306 for lateral readout according to one embodiment. An elongated scintillator crystal 1350 has a rectangular prism shape and is covered (i.e., enclosed or coated) on five sides by a reflective material (e.g., a coating or thin layer) and is covered on one side by the photosensor array 1301-1306. The elongated scintillator crystal 1350 is depicted as having four sides, two axially oriented and two transversely oriented. A distal face 1312 is depicted as a radiating face and is radially oriented into the detector ring to receive gamma photons from annihilation events. A proximal face 1311 is positioned at the other end of the length of the elongated scintillator crystal 1350. The photosensor arrays 1301-1306 are positioned on the axially oriented sides of the elongated scintillator crystal 1350. Each photosensor array 1301, 1302, 1303, 1304, 1305, and 1306 may include a plurality (e.g., a two-dimensional array) of photodiodes. For example, the photosensor array 1301 may include a two-dimensional array of single-photon avalanche photodiodes operating in Geiger mode, such that the photosensor array 1301 is a silicon photomultiplier tube.

[0155] For example, gamma radiation (e.g., gamma photons) may impact location 1307 near the distal face 1312 of the elongated scintillator crystal 1350. In prior embodiments where a single photosensor is positioned proximally, the scintillation light must propagate from location 1307 to the proximal end of the elongated scintillator crystal 1350, possibly undergoing absorption and / or multiple reflections along the length of the elongated scintillator crystal 1350 and propagating at different angles, resulting in a distribution of photon arrival times at the proximal detection end face. This propagation causes the scintillation light to attenuate, and further results in a distribution of photon arrival times, thereby slowing down the signal rise time and subsequent overall photodetection (e.g., as combined with Fig. 9 discussed), which results in timing jitter.

[0156] In contrast, the laterally positioned photosensor arrays 1301-1306 detect scintillation light along the sides of the elongated scintillator crystals 1350, which can achieve higher light collection efficiency and shorter photon propagation paths. The faster detection provided by the laterally positioned photosensor arrays 1301-1306 exhibits reduced attenuation and range of photon arrival time distribution, thereby increasing photosensor output signal and slope.

[0157] The laterally positioned (e.g., laterally mounted, laterally integrated, or otherwise laterally positioned) photosensor arrays 1301-1306 also facilitate determining the interaction depth or origin of the scintillation (e.g., locating location 1307), thereby providing the scanner system with the ability to reduce parallax errors, as previously described, and also achieve improved time-of-flight resolution by reducing the spatiotemporal uncertainty of the gamma photon absorption / scintillation point. For example, a scintillation point originating from a gamma photon at location 1307, absorbed within the distal portion of the elongated scintillator crystal 1350, will result in an illumination (or photon impact) gradient in magnitude (or photon flux) across the photosensor arrays 1306-1301. By measuring the output magnitude of each photosensor array 1301-1306, the gamma photon absorption region can be determined.

[0158] As a second example, a gamma photon 1308 is emitted at an oblique angle to the elongated scintillator crystal 1350, resulting in scintillation as shown, resulting in greater illumination of the photosensor arrays 1304 and 1305 than the photosensor array 1306, and resulting in a gradient of light extending in two directions (i.e., proximal and distal), thereby enabling the determination of a gamma absorption point within the elongated scintillator crystal 1350 in a ratiometric manner. As another example, a gamma photon 1309 is absorbed within the proximal end of the elongated scintillator crystal 1350, resulting in the greatest degree of light in the photosensor array 1302 relative to all other photosensor arrays 1301 and 1303-1306. From this absorption point, a gradient of light will extend in the proximal and distal directions of the elongated scintillator crystal 1350, resulting from the absorption of light along the lateral length of the elongated scintillator crystal 1350, again enabling the determination of a ratiometric measure of the absorption point by comparing the output amplitude and / or time of each photosensor. One method of determining the absorption region is by the following method for a set of lateral photosensors, for example, S = {1301, 1302, ... 1306}:

[0159] in, AF i = the amplitude fraction of photosensor i, where i∈S, AP i = peak amplitude of photoelectric sensor i, |S| represents the cardinality of set S, and For all sensors j in the set S, I = argmax(AF[j]) where I is the index of the sensor with the largest amplitude, representing the geometric area of ​​scintillation (or part of the scintillation depending on the inter-crystal scattering).

[0160] The amplitude peak AP can be the peak value of signal 902, for example, representing the approximate total charge collected by the photosensor, as previously described. Alternatively, other methods can optionally be used to obtain AP, such as integrating the fast timing signal 901 or taking the peak value of signal 901 through digital signal processing. In addition, only certain amplitude ranges can be optionally considered valid, thereby rejecting scattered or piled energy (for example, by using a low-level discriminator (LLD) and a high-level discriminator (ULD)). Alternatively, scattered energy deposited in multiple crystals (i.e., less than the LLD value) can be used to interpolate scintillation locations occurring between two or more crystals, where the total energy sum between the two or more crystals is between the LLD and ULD values ​​and is therefore generally considered to be caused by a single gamma photon with inter-crystal scattering. Further, the scattered energy measured between the various crystals can be used to reconstruct the geometric response line of a single gamma photon, where the total energy is summed between the LLD and ULD (for example, summed to approximately 511 keV). A gamma photon that causes inter-crystal scattering can be determined to terminate at the point of its last partial scintillation location, the initial point of its partial scintillation, or any point in between.

[0161] In some embodiments, the interaction depth is determined by measuring the time at which photons arrive at each photosensor array 1301-1306. For example, for a gamma photon 1309 absorbed within the proximal face 1311 of the elongated scintillator crystal 1350, the earliest detected scintillation light will occur at the nearest photosensor array 1302, followed by 1301 and 1303, followed by 1304, 1305, and finally 1306. The location of the scintillator event can be determined by measuring the first photon detection time of each photosensor (or photosensor array) in the photosensor arrays 1301-1306 based on the known speed within the corresponding elongated scintillator crystal 1350 (e.g., the speed of light traveling through the medium).

[0162] Fig. 13B An exploded view of a single scintillator crystal pixel 1300 is shown with the photosensor arrays 1301 - 1306 and reflective material panels 1321 - 1325 separated from the elongated scintillator crystal 1350 according to one embodiment.

[0163] Fig. 13Cis a diagram of an elongated scintillator crystal 1350 having a photosensor array 1301-1306 for lateral photosensor readout, as well as an interposer 1314 and a connector array 1315, according to one embodiment. The interposer 1314 is connected to, or alternatively integrated into, the photosensor array 1301-1306 along a first axially-oriented side of the elongated scintillator crystal 1350. The interposer 1314 may be used to transmit and / or modify electrical signals from the photosensor array 1301-1306 of the elongated scintillator crystal 1350 to a proximal face 1311 (e.g., to the connector array 1315).

[0164] The electrical signals from the photodiode semiconductor chips forming the photosensor array can be directly connected to or integrated within a very thin electrical interposer 1314. The electrical interposer 1314 can be made of, for example, silicon, glass, fiberglass, or an organic substrate to form a homogenous or heterogeneous integrated component. In various embodiments, the interposer 1314 can include integrated connections to the individual photodiodes within the photosensor array. In various embodiments, the interposer 1314 can include, for example, Figure 7 The circuit components shown in the circuit block 701 of FIG. The interposer 1314 can achieve a very high fill factor crystal array. The interposer 1314 can be interconnected using any combination of metal, polysilicon, through silicon vias, conductive pads, wire bonding, or similar technologies. The interposer 1314 can optionally use solder balls, micro bumps, conductive amalgam paste, conductive epoxy, ultrasonic welding, etc. The interposer 1314 can connect, actively and / or passively modify and / or network and / or transmit the signals of multiple photodiodes to the proximal end of the elongated scintillation crystal 1350. The interposer can be embodied as an integrated circuit. The interposer 1314 at the proximal face 1311 can be connected by a connector array 1315, which can include, for example, solder balls, such as a ball grid array (BGA), or a rigid metal conductor.

[0165] In some embodiments, the BGA connections on the interposer 1314 may optionally form connections to a flexible circuit board, thereby enabling the interposer signals to be routed out of the elongated scintillator crystal array volume or otherwise routed therefrom. Thus, the interposer 1314 is distinct from an (optional) flexible circuit substrate, such as a flexible circuit substrate made of polyimide or a thick printed circuit board substrate, such as a substrate made of fiberglass (e.g., typically 1.6 mm thick or thicker).

[0166] In some embodiments, the interposer 1314 receives power or signals from external circuits through the connector array 1315. The interposer 1314 can be passive or active, with the ability to multiplex, amplify, or otherwise process signals (e.g., including conversion to optical signals). The active components for such processing can be fabricated directly within the interposer 1314, such as in the case of a silicon interposer, or they can be bonded and directly connected to the interposer 1314.

[0167] In some embodiments, interposer 1314 provides very thin mechanical support for the photosensor array. Interposer 1314 is also capable of routing power and signals to and from their respective photodiodes. In this way, power and signal routing form a series chain through interposer 1314 as electrical elements, the series chain including the photodiodes in photosensor arrays 1301-1306, interposer 1314, and finally electrical connector array 1315.

[0168] Fig.14A 1 is a diagram of another embodiment of a sub-divided scintillation crystal pixel 1400 having a photosensor array 1401-1406 for lateral photosensor readout according to one embodiment. As shown, the elongated scintillation crystal 1450 has a rectangular prism shape. The elongated scintillation crystal 1450 is divided into a plurality of segmented volumes along its length, wherein a reflective layer is positioned between adjacent faces of the segmented volumes. In the illustrated embodiment, the elongated scintillation crystal 1450 is segmented into six cubic volumes 1451-1456, each separated by a reflective layer 1420 (see FIG. 1 ). Fig. 14B ). Thus, each segmented cubic volume 1451-1456 contains the scintillation light within its unique volume, sharing light only with its corresponding photosensor array 1401-1406. Little to no scintillation light is shared between cubic volumes 1451-1456.

[0169] For example, scintillation light generated by impinging gamma photons within cubic volume 1454 is internally reflected within cubic volume 1454 and received (e.g., detected) by photosensor array 1403. Similarly, scintillation light generated within cubic volume 1451 is detected only or almost only by photosensor array 1406. Figure 14A-Figure 14C One advantage of the embodiment shown in is that the maximum amount of scintillation light can be identified as originating from a specific area (i.e., cubic volume). In addition, complete collection of light occurs within the specific cubic volume, thereby reducing the photon propagation length and therefore reducing the timing variation of the photons arriving at the photosensor.

[0170] Fig. 14BAn exploded view of a sub-blocked scintillation crystal pixel 1400 is shown according to one embodiment, wherein the photosensor arrays 1401-1406 and the reflective material panel 1420 separate each segmented cubic volume 1451-1456. The elongated scintillation crystal may be segmented into more or less than six segmented volumes. Furthermore, each segmented volume need not be a cube. For example, the elongated scintillation crystal may be segmented into only two volumes, each still in the shape of an elongated rectangular prism. In other embodiments, the elongated scintillation crystal may not have square end faces, so the segmented volumes are not necessarily cubic volumes.

[0171] Fig. 14C is a diagram of a sub-blocked scintillation pixel having a photosensor array 1401-1406 for lateral photosensor readout, as well as an interposer 1414 and a connector array 1415 according to one embodiment. The interposer 1414 may be configured to combine Fig. 13C The inserter 1314 of the embodiment described above can be operated and / or configured in any of the variants and embodiments. Similarly, the connector array 1415 can be configured in accordance with the combination of Fig. 13C The connector array 1415 of the embodiment of the present invention may be operated and / or configured in accordance with any of the variations and embodiments described herein. As in all figures, the sizes, proportions, shapes and dimensions of the various components and elements are not drawn to scale and are not necessarily drawn in proportion relative to each other. For example, the connector array 1415 may be much smaller and / or differently shaped than shown. In addition, the inserter 1414 and / or the connector array 1415 may be closer to the proximal face 1411 of the elongated scintillator crystal 1450 or extend further from the proximal face 1411 of the elongated scintillator crystal 1450.

[0172] Another advantage of lateral light sensing is realized when using highly attenuated scintillator crystals (which might otherwise be incompatible with end readout approaches, which is undesirable). By arranging the photosensors on the sides of a rectangular prism, light localized to the scintillation can be efficiently captured, which would otherwise be significantly attenuated in an end readout approach. Furthermore, for scintillators of low density and stopping power, where longer crystals may be required to achieve similar conversion efficiency (or stopping power) as shorter, denser crystals, the advantage of lateral light sensing is greater because the photon arrival time range can be greatly reduced by side light sensing, whereas end readout would produce a longer photon path and reduce timing resolution. Thus, lateral light sensing (or readout) can enable the use of more economical or higher performance scintillators that would otherwise be incompatible or undesirable for end readout configurations.

[0173] Fig.15Ais a schematic diagram of a scintillation crystal block 1500 of a detector module according to one embodiment, the scintillation crystal block 1500 having lateral readout photosensors 1501-1504 and a heat plate 1575 for thermal management. The use of lateral light sensing by the photosensor array 1501-1504 allows the heat plate 1575 to be placed on the proximal face 1511 of the scintillation crystal block 1500 for thermal conduction. The distal face 1512 of the scintillation crystal block 1500 is oriented to face radially toward the aperture of the scanner to receive gamma radiation therefrom. The heat plate 1575 conceptually illustrates a thermal conductor in contact with the proximal face 1511 of each respective elongated scintillation crystal in the scintillation crystal block 1500. In some embodiments, the reflective material on the five faces of each elongated scintillation crystal can be thermally conductive.

[0174] Thus, the heat plate 1575 can conduct heat through the reflective material on the proximal face 1511 of each elongated scintillator crystal to cool the elongated scintillator crystal. The reflective material (e.g., a coating or thin material layer) on the proximal face 1511 of each elongated scintillator crystal can be thermally coupled to the reflective material on the other faces of each elongated scintillator crystal. In such an embodiment, the elongated scintillator crystal (and the laterally positioned photosensor array) can be cooled from all sides as heat is conducted through the reflective material into the heat plate 1575.

[0175] In some embodiments, the heat plate 1575 also acts as a reflector. For example, the heat plate 1575 can be used as a reflector at the same time by polishing, sandblasting or other finishing processes to provide specular reflection or diffuse reflection. In some embodiments, the heat plate 1575 can be connected to the proximal end face 1511 of the elongated scintillator crystal block by thermally conductive paste, or connected to the reflector of each elongated scintillator crystal by welding, or both. Examples of thermally conductive and reflective materials that can be used to form the reflector (or reflective coating) and / or the heat plate 1575 include, but are not limited to, aluminum, magnesium, silver or compounds thereof (e.g., Al2O3, MgO) and stainless steel. In some embodiments, the heat plate 1575 can be slotted or otherwise have holes or electrical conduits to support the routing of signals into and out of the crystal block volume that supports side readout. In some embodiments, the heat plate 1575 is formed as part of a reflector material that includes a continuous metal member formed to wrap around the sides and ends of each crystal. In some embodiments, the heat plate 1575 is formed as a continuous metal member to wrap around the distal end 1512 and the sides of the scintillation crystal block 1500 .

[0176] In the illustrated embodiment, the two-dimensional array of elongated scintillator crystals comprises a 4x4 array of a total of 16 elongated scintillator crystals. It will be appreciated that any number of elongated scintillator crystals in a one-dimensional or two-dimensional array may be used to form a scintillator crystal block in a detector module. In the illustrated embodiment, each elongated scintillator crystal is associated with six discrete photosensor arrays, each of which may include a plurality of photodiode sensors. In an alternative embodiment, a single elongated photosensor array may be positioned on an axially oriented side of each elongated scintillator crystal.

[0177] The use of thermally conductive materials in contact with elongated scintillating crystals and in direct or indirect contact with adjacent photosensors can stabilize and / or increase scintillator light emission and photosensor gain by cooling. Thermal fluctuations in elongated scintillating crystals and / or photosensors result in a modulated signal output for a given energy absorption, thereby reducing the resolution of the measurement. In various embodiments, the PET scanner is intended to be installed under controlled indoor conditions. In some cases, the PET scanner can use water at ambient temperature or lower for use by the scanner. In other cases, an air-to-air heat exchanger can be used to remove convective electron heat from the inside of the scanner to a cooler surrounding environment.

[0178] In other cases, heat may be conducted from a thermally connected plate (e.g., heat plate 1575) to the external ambient air (i.e., outside the scanner housing) by liquid conduction (e.g., through water) using an air-to-liquid heat exchanger. In other embodiments, a PET scanner may conduct heat using a liquid-to-liquid heat exchanger. In many cooling configurations, heat conduction or convection is only generally or extensively applied within the housing of the scanner without any direct or significant heat conduction from or to the array of elongated scintillation crystals within a single detector module (e.g., between laterally adjacent crystal faces within a single detector module).

[0179] Furthermore, in many embodiments, the PET scanner may not have a cooling system configured to cool the system below room temperature. One reason a PET scanner may be configured to operate at or near ambient temperature is due to the risk of condensation on very cold actively cooled surfaces that could damage components therein. Therefore, the scanner thermal control system may monitor dew point conditions and their temperature thresholds to keep the cooling surfaces above the condensation temperature.

[0180] However, some embodiments of the presently described systems and methods utilize a cooling system to directly cool the elongated scintillator crystals and / or the photosensors to improve thermal regulation and output resolution (e.g., scintillator light and photosensor signals). In addition, some embodiments of the presently described systems and methods cool the elongated scintillator crystals to below ambient temperature to increase scintillator light output and photosensor gain. By isolating cooling surfaces at temperatures below the dew point so that the temperature of their outer surfaces exposed to ambient conditions is above the dew point threshold, the elongated scintillator crystals and / or photosensors in the detector modules of a PET scanner can operate at temperatures much lower than ambient room conditions. In some examples, each gamma radiation detector module is sealed with an insulating layer to encapsulate the elongated scintillator crystals, photosensor array, and reflective material. In some embodiments, routing wires can pass through the sealed housing.

[0181] Fig. 15B 1 is a schematic diagram of a scintillation crystal block 1590 according to one embodiment, the scintillation crystal block having photosensors 1501, 1502, 1503, and 1504, interposers 1514, 1515, 1516, and 1517, and connector arrays 1524, 1525, 1526, and 1527. In various embodiments, the detector module receives gamma photons on the distal face 1512. The gamma photons scintillate within one (or more) elongated scintillation crystals to produce lower frequency scintillation photons. The scintillation photons are received by lateral photosensors associated with the elongated scintillation crystals. The electrical signals generated by the lateral photosensors are transmitted by the electrical interposers to the connector array (e.g., a ball grid array). In the illustrated embodiment, each of the sixteen elongated scintillation crystals is associated with six photosensor arrays and one laterally positioned interposer. Thus, scintillator crystal block 1590 includes sixteen elongated scintillator crystals, sixteen interposers, and ninety-six photosensor arrays (eg, ninety-six silicon photomultipliers, each silicon photomultiplier including a plurality of photodiodes), not all of which are labeled with reference numbers.

[0182] In some embodiments, a single laterally positioned interposer may be associated with a plurality of elongated scintillator crystals. For example, two adjacent scintillator crystals may share one interposer. As another example, four scintillator crystals per column may share a single interposer (e.g., spanning the sides of all four scintillator crystals). In such an example, the scintillator crystal block 1590 would include sixteen elongated scintillator crystals, four interposers, and ninety-six photosensor arrays (each of which may include a plurality of individual photodiodes). Even further, a single laterally positioned interposer may integrate photosensors on the faces of two opposing sides, thereby optically matching two or more crystals on each side of the interposer.

[0183] In various embodiments, a PET scanning system includes one or more detector module rings that are arranged with photosensors and electrical inserters that are oriented axially relative to the rings. In such embodiments, the axial width of each detector module ring includes the sum of the widths of the elongated scintillator crystals in the axial direction, the sum of the widths of the axially arranged photosensor arrays on the sides of each elongated scintillator crystal, the sum of the widths of the axially arranged reflectors or reflective coatings, and the sum of the widths of the axially arranged inserters connected to the photosensors. The illustrated embodiment is simplified to include only four elongated scintillator crystals in the axial direction, each of which has square ends (e.g., the height and width of each elongated scintillator crystal are the same).

[0184] Thus, the width of the scintillation crystal block 1590 in the axial direction is equal to four times the sum of the widths of the interposer, the photosensors, and the elongated scintillation crystals, plus a relatively negligible width of the reflective coating on the axially aligned surfaces of the elongated scintillation crystals. The width of the detector module in the transverse direction is equal to four times the height of the elongated scintillation crystals, plus a relatively negligible width of the reflective coating on the axially aligned surfaces of the elongated scintillation crystals.

[0185] Fig. 15C yes Fig. 15B Schematic diagram of a scintillation crystal detector block 1590. In which heat plates 1541, 1542, 1543 and 1545 are added to the proximal end face 1511 of each elongated scintillation crystal column of the scintillation crystal block 1590. The heat plates 1541, 1542, 1543 and 1545 can be combined according to Fig.15A Any of the various embodiments and variations described for the thermal plate 1575 can be configured and / or operated. The use of multiple thermal plates 1541, 1542, 1543, and 1545 facilitates routing of signals through the interposer to the connector array and / or facilitates connection with the connector array.

[0186] Fig.15D is a diagram of a detector block of scintillation crystals 1590 connected to a dual channel processing circuit 1580 according to one embodiment. In the illustrated embodiment, the dual channel processing circuit 1580 includes connection features to connect to the connector array 1520 of the interposer 1514 associated with the corresponding elongated scintillation crystal 1550. The interposer 1514 operates to transmit electrical signals from the plurality of photosensor arrays 1501 of each elongated scintillation crystal 1550 to the corresponding connector array 1520. A thermal plate 1540 is mounted to facilitate thermal cooling of the electronic and conductive components within the elongated scintillation crystal 1550, photosensors 1501, and / or interposer 1514.

[0187] In the illustrated embodiment, each scintillation crystal 1550 is associated with a different inserter 1514. In some embodiments, a single inserter may be associated with multiple scintillation crystals. For example, a single inserter may have a width that spans the sides of multiple scintillation crystals. A single inserter may, for example, be associated with four or five scintillation crystals.

[0188] As combined Figure 7 and Fig.12 As described, the first channel of the processing circuit 1580 includes an energy circuit and a timing circuit to process the signal with dot-shaped shadows from the elongated scintillation crystal 1550. The second channel of the processing circuit 1580 includes an energy circuit and a timing circuit to process the signal with cross-hatching from the elongated scintillation crystal 1550. The PET scanner may include a plurality of detector modules, each of which includes one or more scintillation crystal blocks 1590 and a processing circuit 1580. As is understood in the art, the image processing circuit of the PET scanner may receive the timing signal and the energy signal from the dual-channel processing circuit 1580 of each detector module to reconstruct an image.

[0189] Fig.16 is a schematic diagram of an elongated scintillation crystal block 1600 having double-sided interposers 1614 and 1615 connected to the photosensor arrays of adjacent elongated scintillation crystals according to one embodiment. As described in connection with other embodiments, thermal plates 1641, 1643, and 1645 are spaced apart to contact the proximal face 1611 of the elongated scintillation crystal block 1600. The specific dimensions of thermal plates 1641, 1643, and 1645 may be selected to facilitate routing of signals via interposers 1614 and 1615 through connector arrays 1620 and 1622.

[0190] In the illustrated example, the elongated scintillator crystal block 1600 includes sixteen elongated scintillator crystals, including numbered elongated scintillator crystals 1650, 1651, 1652, 1653, and 1654, and eight inserters, including numbered inserters 1614 and 1615. It will be appreciated that the scintillator crystal block may include any number of elongated scintillator crystals arranged in a one-dimensional or two-dimensional array, with an even or odd number of elongated scintillator crystals. Each elongated scintillator crystal has a reflective material or coating on five faces, so that any optical radiation generated by a scintillation event within any given elongated scintillator crystal is reflected for detection by an associated photosensor array. In the illustrated example, each inserter has a photosensor array on opposite faces in the direction of the axial hole.

[0191] As shown, the interposer 1614 has a first photosensor array 1601 on a first side to detect optical radiation within (e.g., scintillation events therefrom) the elongated scintillator crystal 1650. A second photosensor array 1602 on a second side of the interposer 1614 detects optical radiation within the elongated scintillator crystal 1651. Electrical signals from the first photosensor array 1601 and the second photosensor array 1602 are transmitted, processed, and / or partially processed through the interposer 1614. The interposer 1614 transmits the processed or partially processed electrical signals from the photosensors 1601 and 1602 to a connector array (not visible) at the proximal end 1611 of the block 1600.

[0192] Reiterating the functionality of the double-sided interposer, the interposer 1615 has a first photosensor array 1603 on a first side to detect optical radiation within the elongated scintillator crystal 1652. A second photosensor array 1604 on a second side of the interposer 1615 detects optical radiation within the elongated scintillator crystal 1653. Electrical signals from the first photosensor array 1603 and the second photosensor array 1604 are transmitted, processed, and / or partially processed through the interposer 1615. The interposer 1615 transmits the processed or partially processed electrical signals from the photosensors 1603 and 1604 to the connector array 1620 at the proximal end 1611 of the block 1600. In the illustrated example, each of the first photosensor array 1603 and the second photosensor array 1604 includes six discrete photosensor arrays. It will be appreciated that each photosensor array may include more or fewer discrete photosensors.

[0193] In the illustrated embodiment, the block 1600 includes an even number of scintillation crystals in the direction of the axial hole, so that an even number of double-sided inserters are used. In some embodiments, the elongated scintillation crystal block includes an odd number of scintillation crystals in the direction of the axial hole. In such embodiments, an odd number of inserters can be used, where the inserters at one end of the block are single-sided inserters, each inserter servicing a single elongated scintillation crystal.

[0194] Fig.17A ring 1700 of detector modules 1790 (e.g., blocks) according to one embodiment is shown, having axially oriented photosensors on the sides of elongated scintillating crystals. Each detector module 1790 is simplified but is intended to represent various embodiments of detector modules described herein. For example, although not shown in each example, each detector module 1790 may include an inserter positioned on an axially aligned side of a photosensor, which in turn is positioned on an axially aligned side of each corresponding elongated scintillating crystal. The distal face 1712 of each detector module 1790 is positioned toward the ring 1700 to receive gamma photons from annihilation events within the patient 1750. The ring 1700 has a diameter 1710 and an axial width 1720. The ring 1700 defines an axial direction (through the hole) and a transverse direction around the circumference of the ring 1700.

[0195] In the illustrated embodiment, the slender scintillator crystals of each detector module 1790 or "block" are seamlessly positioned in the axial direction of the detector ring with no substantial gap therebetween, such that two adjacent slender scintillator crystals are separated from each other by: (i) a photosensor array on a first axially-oriented side of one of the adjacent slender scintillator crystals, (ii) a reflective material on a second axially-oriented side of another adjacent slender scintillator crystal, and (iii) optionally, an inserter connected to the photosensor array on the first axially-oriented side of each respective slender scintillator crystal.

[0196] Fig.18 A ring 1800 of detector modules 1890 is shown with transversely-oriented photosensors on the sides of elongated scintillator crystals according to one embodiment. Again, each detector module 1890 in the ring 1800 is simplified but is intended to represent any of the various embodiments of detector modules described herein. For example, each detector module 1890 in the ring 1800 can include an interposer positioned on the transversely-aligned side of the photosensors, which in turn are positioned on the transversely-aligned side of each respective elongated scintillator crystal.

[0197] The transverse orientation of the inserter and photosensor results in the axial width 1820 of the ring 1800 being less than Fig.17 The axial width 1720 of the middle ring 1700 is greater than that of the middle ring 1700, even though the same number of detector modules are used to form the ring. However, the transverse orientation of the interposer and the photosensors results in the inner diameter 1810 of the ring 1800 being greater than that of the middle ring 1700. Fig.17 The inner diameter 1710 of the middle ring 1700 is the same even when the same number of detector modules are used. As described herein, Fig.17 The axial orientation results in a smaller inner diameter 1710 for a given number of detector modules. For the reasons described herein, in some applications, a smaller diameter is advantageous.

[0198] Fig.19 1900 is a graph of sensitivity gain and fill factor loss for lateral photosensor readout using end-readout as a baseline versus various crystal geometries according to one embodiment. As shown, lateral sensitivity (compared to end-readout) is plotted by the dashed line (labeled lateral gain) as the length of the elongated scintillator crystal relative to the width of the elongated scintillator crystal end face (plotted as a 2 mm dimension). 2 , 3mm 2 and 4mm 2 square end faces), increases almost linearly. In contrast, as the aspect ratio of length to width increases accordingly (i.e., for thinner crystals), the photosensor thickness (plotted as 370 μm) becomes significant and the fill factor loss increases, as shown by the dashed line (labeled Fill Factor Loss), thereby reducing the lateral readout gain. Nevertheless, the sensitivity gain produced by the lateral photosensor readout outweighs the fill factor loss, as shown by the solid line (labeled Combined Gain) plotting the ratio.

[0199] Fig. 20 2000 is a graph of overall lateral sensitivity gain (compared to end readout) for rectangular prismatic scintillator crystals of varying lengths (e.g., 10-30 mm long) with square end faces according to various embodiments. As shown, longer / higher aspect ratios exhibit higher side readout gain. Fill factor loss becomes prevalent in thinner crystals (e.g., less than about 2 mm), while short, wide crystals have minimal gain due to the reduced lateral photosensor surface area compared to end readout.

[0200] exist Fig.19 and Fig. 20 In the context of the information conveyed in, various embodiments of the presently described systems and methods include a detector module having a two-dimensional or three-dimensional array of scintillation crystals whose target size is for relatively higher gain relative to the gain of the end readout. In some embodiments, the size of the elongated scintillation crystals is selected to increase the overall gain while reducing or minimizing the fill factor loss. For example, the two-dimensional array of elongated scintillation crystals in the detector module can utilize scintillation crystals having a size and aspect ratio that provides a fill factor greater than 85%. In various embodiments, the length between the distal and proximal end faces of each elongated scintillation crystal is between 10 mm and 30 mm, the ratio of the length to the axial width is 3-15, and the ratio of the length to the transverse width is 3-15. In various embodiments, the end faces of the elongated scintillation crystals are square so that the transverse width and the axial width are equal.

[0201] Fig.212100 is a diagram of the photon dynamic range of a silicon photomultiplier photosensor having 1,600 single photon avalanche photodiode microcells per square millimeter with different sizes according to one embodiment. The dynamic range is proportional to the crystal lateral surface area. In various embodiments, each elongated scintillator crystal is associated with a photosensor array. The photosensor array can include one or more silicon photomultiplier photosensors positioned on an axially oriented side of the elongated scintillator crystal.

[0202] Fig. 22 is a graph 2200 of the light output of the scintillator LYSO:Ce versus temperature according to one embodiment. According to various embodiments, the elongated scintillator crystals may be maintained at a temperature below ambient temperature to obtain a higher light yield output. For example, in some embodiments, an approximately 4% increase in light output may be achieved by operating the elongated scintillator crystals within each detector module at -40°C rather than at room temperature of approximately 21°C. It is noteworthy that the voltage change per hertz bandwidth of the thermally induced electronic noise of the photosensor array (referred to in the art as Johnson-Nyquist noise) is given by the following formula:

[0203] in, k B = Boltzmann's constant in joules per kelvin, T = the absolute temperature of the resistor in Kelvin, and R = quenching resistance of the photoelectric sensor or other equivalent resistance.

[0204] Therefore, for a decrease in temperature, all other factors remaining constant, the thermal noise decreases as given by:

[0205] in, T1 = upper operating temperature in degrees Kelvin, and T2 = lower operating temperature, in degrees Kelvin.

[0206] Thus, for example, a decrease from 21°C to -40°C results in a total reduction in root mean square (RMS) noise of approximately 11%. For example, combined with an approximately 4% gain in light output, the SNR can be improved by nearly 17%. In addition to the reduction in thermal noise of the silicon photomultiplier tube photosensor at lower temperatures, the photoelectric gain is also improved. According to various embodiments, a detector module with lateral light sensing and a cooled scintillator crystal below ambient temperature can improve light collection efficiency and include a thermal management component (e.g., a heat plate) to effectively reduce the temperature, thereby improving scintillator light output, increasing photosensor conduction, and reducing noise.

[0207] In some embodiments, condensation on cooling surfaces in a PET scanning environment can be eliminated or reduced by controlling the surrounding environment and / or by isolating the cooling surfaces from warmer, humid air. At a surface temperature of -40°C, under standard conditions, the absolute humidity is approximately 119.3 mg / m 3 (or 79ppm) may cause condensation. In some embodiments, the cooled detector module is enclosed in a substantially airtight chamber. In some embodiments, the airtight chamber is positively pressurized with an inert gas (e.g., nitrogen). The water vapor concentration in the sealed and positively pressurized chamber can be reduced to less than 1ppm. In some embodiments, the air in the sealed chamber is evacuated using a vacuum pump to reduce the pressure to about ten millibars or less. Dry nitrogen can be used to purge the chamber. Thereafter, dry nitrogen can be used to pressurize the insulating chamber to above atmospheric pressure.

[0208] The present disclosure has been carried out with reference to various exemplary embodiments (including best mode). However, those skilled in the art will recognize that the exemplary embodiments can be changed and modified without departing from the scope of the present disclosure. Although the principles of the present disclosure have been shown in various embodiments, many modifications of structure, arrangement, ratio, elements, materials and parts can be adjusted for specific environments and / or operational requirements without departing from the principle and scope of the present disclosure. These and other changes or modifications are intended to be included within the scope of the present disclosure.

[0209] Item 1. A gamma radiation detector module comprising: an array of scintillation crystals positioned on the detector, wherein each scintillation crystal comprises a proximal face, a distal face oriented into the detector to receive gamma photons, and four sides comprising a first side, a second side, a third side, and a fourth side; an array of photosensors positioned along the first side of each scintillation crystal to detect scintillation photons; and a reflective material positioned on the proximal face, the distal face, and the second side, the third side, and the fourth side of each scintillation crystal to internally reflect scintillation photons.

[0210] Item 2. A gamma radiation detector module according to the item, wherein each scintillation crystal comprises an elongated scintillation crystal, wherein the elongated scintillation crystal array is configured to be positioned on a detector ring, wherein the four sides of each scintillation crystal comprise two axially oriented sides and two transversely oriented sides, wherein the distal end face is radially oriented into the detector ring to receive the gamma photons.

[0211] Clause 3. The gamma radiation detector module of Clause 2, wherein the first side is axially oriented such that the photosensor is positioned along the axially oriented side.

[0212] Clause 4. The gamma radiation detector module of Clause 1, wherein the array of elongated scintillator crystals is configured to be positioned on a detector panel that operates in conjunction with an opposing detector panel.

[0213] Clause 5. The gamma radiation detector module of Clause 1, wherein each scintillator crystal is in the shape of an elongated N-sided polygonal prism, where N is an integer value.

[0214] Clause 6. The gamma radiation detector module of Clause 1, wherein each scintillator crystal comprises a cubic shaped scintillator crystal.

[0215] Clause 7. The gamma radiation detector module of Clause 1, wherein each scintillator crystal comprises an elongated scintillator crystal in the shape of an elongated rectangular prism.

[0216] Clause 8. The gamma radiation detector module of Clause 7, wherein the distal face of each elongated scintillator crystal is square.

[0217] Item 9. A gamma radiation detector module according to Item 8, wherein the length between the distal face and the proximal face of each elongated scintillator crystal is between 10 millimeters and 30 millimeters, and wherein the ratio of the length of each elongated scintillator crystal to the width of each elongated scintillator crystal is between three and ten (3-10).

[0218] Clause 10. The gamma radiation detector module of Clause 3, wherein the array of elongated scintillator crystals comprises a one-dimensional array of elongated scintillator crystals arranged along an axial direction of the detector ring.

[0219] Item 11. A gamma radiation detector module according to Item 10, wherein the slender scintillator crystals are seamlessly positioned in the axial direction of the detector ring without gaps therebetween, so that a given slender scintillator crystal is separated from an axially adjacent slender scintillator crystal by: (i) a photosensor on a first axially oriented side of the given slender scintillator crystal and (ii) a reflective material on a second axially oriented side of an axially adjacent slender scintillator crystal.

[0220] Clause 12. A gamma radiation detector module according to Clause 3, wherein the slender scintillation crystal array comprises a two-dimensional array of slender scintillation crystals, wherein the two-dimensional array of slender scintillation crystals comprises M slender scintillation crystals in the axial direction of the detector ring and N slender scintillation crystals in the transverse direction of the detector ring, wherein M and N are both integer values.

[0221] Item 13. A gamma radiation detector module according to Item 12, wherein the slender scintillation crystals are seamlessly positioned adjacent to each other in the axial and transverse directions of the detector ring without any gaps therebetween, so that a given slender scintillation crystal is separated from an axially adjacent slender scintillation crystal by: (i) a photosensor on a first axially oriented side of the given slender scintillation crystal and (ii) a reflective material on a second axially oriented side of the axially adjacent slender scintillation crystal, and so that the given slender scintillation crystal and the transversely adjacent slender scintillation crystal are separated from each other in the transverse direction by the reflective material.

[0222] Clause 14. A gamma radiation detector module according to Clause 13, wherein the thickness of the photosensors and reflective materials between adjacent elongated scintillator crystals in the axial direction of the detector ring is less than 500 microns, and wherein the thickness of the reflective material between adjacent elongated scintillator crystals in the transverse direction of the detector ring is less than 100 microns.

[0223] Item 15. A gamma radiation detector module according to Item 14, wherein the length between the distal face and the proximal face of each slender scintillator crystal is between 10 mm and 30 mm, wherein the ratio of the length of each slender scintillator crystal to the axial width is between three and ten (3-10), and wherein the ratio of the length of each slender scintillator crystal to the transverse width is between three and ten (3-10).

[0224] Clause 16. The gamma radiation detector module of Clause 13, wherein the fill factor of the area of ​​the distal face of the elongated scintillator crystal to the total area is greater than 85%.

[0225] Clause 17. The gamma radiation detector module of Clause 1, wherein the reflective material comprises a thermally conductive reflective material.

[0226] Clause 18. The gamma radiation detector module of Clause 17, wherein the thermally conductive reflective material comprises at least one of aluminum, magnesium, silver, stainless steel, and combinations thereof.

[0227] Clause 19. The gamma radiation detector module of Clause 17, further comprising a thermally conductive plate coupled to the proximal faces of at least some of the scintillator crystals.

[0228] Clause 20. A gamma radiation detector module according to clause 1, further comprising: a sealed housing for encapsulating the scintillation crystals, the photosensors and the reflective material; and wiring for routing the electrical signal from the photosensor of each scintillation crystal from its first side toward its proximal face and through the sealed housing.

[0229] Clause 21. The gamma radiation detector module of Clause 1, further comprising wiring for routing electrical signals from the photosensor of each scintillator crystal from its first side toward its proximal face.

[0230] Clause 22. A gamma radiation detector module according to Clause 3, further comprising: an inserter connected to the photosensor along a first axially-oriented side of each corresponding elongated scintillator crystal, wherein the inserter is operable to transmit an electrical signal from the photosensor of each corresponding elongated scintillator crystal toward a proximal face of each corresponding elongated scintillator crystal.

[0231] Clause 23. A gamma radiation detector module according to Clause 22, wherein the slender scintillation crystal array comprises a two-dimensional array of slender scintillation crystals, wherein the two-dimensional array of slender scintillation crystals comprises M slender scintillation crystals in the axial direction of the detector ring and N slender scintillation crystals in the transverse direction of the detector ring, wherein M and N are both integer values.

[0232] Clause 24. A gamma radiation detector module according to clause 23, wherein the slender scintillation crystals are seamlessly positioned adjacent to each other in the axial and transverse directions of the detector ring without any gaps therebetween, so that a given slender scintillation crystal is separated from an axially adjacent slender scintillation crystal by: (i) a photosensor on a first axially oriented side of the given slender scintillation crystal and (ii) a reflective material on a second axially oriented side of the axially adjacent slender scintillation crystal, and so that the given slender scintillation crystal and the transversely adjacent slender scintillation crystal are separated from each other in the transverse direction by the reflective material.

[0233] Clause 25. The gamma radiation detector module of Clause 1, wherein the photosensor array of each scintillator crystal comprises at least one two-dimensional array of individual photodiodes.

[0234] Clause 26. The gamma radiation detector module of Clause 25, wherein each photodiode comprises a single photon avalanche diode (SPAD) operated in Geiger mode.

[0235] Clause 27. The gamma radiation detector module of Clause 1, wherein the photosensor array of each scintillator crystal comprises a plurality of discrete photosensors to provide different photosensor measurements corresponding to different depth ranges between the proximal face and the distal face of each scintillator crystal.

[0236] Clause 28. The gamma radiation detector module of Clause 27, wherein the photosensor array of each scintillator crystal comprises a plurality of two-dimensional arrays of individual photodiodes.

[0237] Clause 29. The gamma radiation detector module of Clause 28, wherein each photodiode comprises a single photon avalanche diode (SPAD) operated in Geiger mode.

[0238] Clause 30. The gamma radiation detector module of Clause 29, wherein each photosensor comprises a silicon photomultiplier (SiPM), such that each scintillator crystal is associated with a plurality of SiPMs extending along a first side thereof.

[0239] Item 31. A gamma radiation detector module according to Item 30, which also includes a processing circuit, which includes: a plurality of energy circuits, each of which is configured to generate an energy signal, which is the sum of the energy detected by one or more silicon photomultiplier tubes; and a plurality of timing signal circuits, each of which is configured to generate different timing signals for scintillation photons detected by one or more silicon photomultiplier tubes.

[0240] Clause 32. The gamma radiation detector module of clause 31, wherein the processing circuit is a dual channel high speed circuit, wherein each channel includes circuitry for digitally encoding a timing signal and an energy signal for one or more silicon photomultipliers.

[0241] Clause 33. The gamma radiation detector module of Clause 32, wherein each scintillation crystal in the array of scintillation crystals uses different high speed circuitry than an adjacent scintillation crystal.

[0242] Clause 34. The gamma radiation detector module of Clause 31, wherein the processing circuit includes a different timing signal circuit for each group of three silicon photomultipliers.

[0243] Clause 35. The gamma radiation detector module of Clause 31, wherein the processing circuit comprises a different energy circuit for each silicon photomultiplier.

[0244] Clause 36. The gamma radiation detector module of Clause 31, wherein the processing circuit comprises a different energy circuit for each group of six silicon photomultipliers.

[0245] Clause 37. The gamma radiation detector module of Clause 27, wherein each scintillator crystal is divided into a plurality of segmented volumes along its length.

[0246] Clause 38. The gamma radiation detector module of Clause 37, wherein each segmented volume of each scintillator crystal has a different crystal time constant.

[0247] Clause 39. The gamma radiation detector module of Clause 38, wherein processing circuitry associated with the detector module determines which segmented volume the scintillation photons originated from based on a time constant of the detected scintillation photons.

[0248] Item 40. A gamma radiation detector module according to Item 27, wherein each scintillation crystal is divided into a plurality of segmented volumes along its length, and a reflective layer is positioned between adjacent faces of the segmented volumes, wherein each segmented volume is associated with a different photosensor array so that scintillation photons generated within the segmented volume are reflected within the segmented volume to be detected by the corresponding photosensor array.

[0249] Item 41. A positron emission tomography (PET) scanning system comprising: a plurality of gamma radiation detector modules positioned to form a detector ring, wherein each detector module comprises: an array of elongated scintillator crystals, each scintillator crystal comprising a proximal face, two axially oriented sides, two transversely oriented sides, and a distal face positioned radially into the detector ring to receive gamma photons, an array of photoelectric sensors positioned along a first axially oriented side of each elongated scintillator crystal to detect scintillation photons, and reflective material positioned on the proximal face, distal face, radial side, and second axially oriented side of each elongated scintillator crystal to internally reflect scintillation photons; a cooling system for cooling the detector modules; and an imaging system connected to the detector modules to generate images based on electronic output from the detector modules.

[0250] Clause 42. The PET scanning system of clause 41, wherein the detector ring is arranged into a plurality of axially aligned detector module rings, wherein each axially aligned ring comprises at least four detector modules.

[0251] Clause 43. A PET scanning system according to Clause 42, wherein the diameter of the detector ring is greater than 25 centimeters, and the axial bore depth of the plurality of axially aligned rings forming the detector ring is less than the diameter of the detector ring.

[0252] Clause 44. The PET scanning system of Clause 41, wherein each elongated scintillator crystal is shaped as an elongated N-sided polygonal prism, where N is an integer value.

[0253] Clause 45. The PET scanning system of Clause 41, wherein each elongated scintillator crystal is shaped as an elongated rectangular prism.

[0254] Clause 46. The PET scanning system of Clause 45, wherein the distal face of each elongated scintillator crystal is square.

[0255] Clause 47. A PET scanning system according to Clause 46, wherein the length between the distal face and the proximal face of each elongated scintillator crystal is between 10 millimeters and 30 millimeters, and wherein the ratio of the length of each elongated scintillator crystal to the width of each elongated scintillator crystal is between three and ten (3-10).

[0256] Clause 48. The PET scanning system of Clause 41, wherein the elongated scintillator crystal array of each detector module comprises a one-dimensional array of elongated scintillator crystals arranged along an axial direction of the detector ring.

[0257] Item 49. A PET scanning system according to Item 48, wherein the slender scintillator crystals of each detector module are seamlessly positioned in the axial direction of the detector ring without gaps therebetween, so that a given slender scintillator crystal is separated from an axially adjacent slender scintillator crystal by: (i) a photosensor array on a first axially oriented side of the given slender scintillator crystal and (ii) a reflective material on a second axially oriented side of an axially adjacent slender scintillator crystal.

[0258] Item 50. A PET scanning system according to Item 41, wherein the slender scintillator crystal array of each detector module includes a two-dimensional array of slender scintillator crystals, wherein the two-dimensional array of slender scintillator crystals of each detector module includes M slender scintillator crystals in the axial direction of the detector ring and N slender scintillator crystals in the transverse direction of the detector ring, wherein M and N are both integer values.

[0259] Item 51. A PET scanning system according to Item 50, wherein the slender scintillator crystals of each detector module are seamlessly positioned adjacent to each other in the axial and transverse directions of the detector ring without any gap therebetween, so that a given slender scintillator crystal is separated from an axially adjacent slender scintillator crystal by: (i) a photosensor array on a first axially oriented side of the given slender scintillator crystal and (ii) a reflective material on a second axially oriented side of the axially adjacent slender scintillator crystal, and so that the given slender scintillator crystal and the transversely adjacent slender scintillator crystal are separated from each other in the transverse direction by the reflective material.

[0260] Item 52. A PET scanning system according to Item 51, wherein the thickness of the photosensor array and the reflective material between adjacent elongated scintillator crystals in each detector module in the axial direction of the detector ring is less than 500 microns, and wherein the thickness of the reflective material between adjacent elongated scintillator crystals in each detector module in the transverse direction of the detector ring is less than 100 microns.

[0261] Item 53. A PET scanning system according to Item 52, wherein the length between the distal face and the proximal face of each slender scintillator crystal is between 10 mm and 30 mm, wherein the ratio of the length of each slender scintillator crystal to the axial width is between three and ten (3-10), and wherein the ratio of the length of each slender scintillator crystal to the transverse width is between three and ten (3-10).

[0262] Clause 54. The PET scanning system of Clause 51, wherein a fill factor of an area of ​​a distal face of the elongated scintillator crystal in each detector module to a total area of ​​each detector module is greater than 85%.

[0263] Clause 55. The PET scanning system of Clause 41, wherein the reflective material comprises a thermally conductive reflective material.

[0264] Clause 56. The PET scanning system of Clause 55, wherein the thermally conductive reflective material comprises at least one of aluminum, magnesium, silver, stainless steel, and combinations thereof.

[0265] Clause 57. The PET scanning system of Clause 56, further comprising a thermally conductive plate coupled to the proximal face of at least some of the elongated scintillator crystals in the array of elongated scintillator crystals of each detector module.

[0266] Item 58. A PET scanning system according to Item 41, wherein each detector module further comprises: a sealed housing for encapsulating the elongated scintillator crystal, the photosensor array and the reflective material; and wiring for routing the electrical signal from the photosensor array of each elongated scintillator crystal from its first side toward its proximal face and through the sealed housing to the imaging system.

[0267] Clause 59. The PET scanning system of Clause 41, further comprising wiring to route electrical signals from the photosensor of each elongated scintillator crystal of each detector module from the first side thereof toward the proximal face thereof.

[0268] Item 60. A PET scanning system according to Item 41, further comprising: an inserter connected to the photosensor array along a first axially-oriented side of each corresponding elongated scintillator crystal of each detector module, wherein the inserter is operable to transmit electrical signals from the photosensor array of each corresponding elongated scintillator crystal toward the proximal face of each corresponding elongated scintillator crystal.

[0269] Item 61. A PET scanning system according to Item 60, wherein the slender scintillator crystal array of each detector module includes a two-dimensional array of slender scintillator crystals, wherein the two-dimensional array of slender scintillator crystals of each detector module includes M slender scintillator crystals in the axial direction of the detector ring and N slender scintillator crystals in the transverse direction of the detector ring, wherein M and N are both integer values.

[0270] Item 62. A PET scanning system according to Item 61, wherein the slender scintillator crystals of each detector module are seamlessly positioned adjacent to each other in the axial and transverse directions of the detector ring without any gap therebetween, so that a given slender scintillator crystal and an axially adjacent slender scintillator crystal are separated from each other in the axial direction by: (i) a photosensor array on a first axially oriented side of a given slender scintillator crystal, (ii) an inserter of a given slender scintillator crystal, and (iii) a reflective material on a second axially oriented side of an axially adjacent slender scintillator crystal, and so that a given slender scintillator crystal and a transversely adjacent slender scintillator crystal are separated from each other in the transverse direction by the reflective material.

[0271] Clause 63. The PET scanning system of Clause 41, wherein each photosensor array of elongated scintillator crystals comprises a two-dimensional array of individual photodiodes.

[0272] Clause 64. A PET scanning system as described in Clause 41, wherein the photosensor array of each elongated scintillator crystal includes a plurality of photosensors to provide different photosensor measurements corresponding to different depth ranges between the proximal face and the distal face of each elongated scintillator crystal.

[0273] Clause 65. The PET scanning system of Clause 64, wherein each photosensor array of elongated scintillator crystals comprises a two-dimensional array of individual photodiodes.

[0274] Clause 66. The PET scanning system of Clause 65, wherein each photodiode comprises a single photon avalanche diode (SPAD).

[0275] Clause 67. The PET scanning system of clause 66, wherein each photosensor comprises a silicon photomultiplier tube (SiPM), such that each elongated scintillator crystal is associated with a plurality of SiPMs extending along a first axially-oriented side thereof.

[0276] Item 68. A PET scanning system according to Item 67, which also includes: an energy circuit for each slender scintillator crystal of each detector module, which is used to generate an energy signal, which is the sum of the energy detected by the photosensor array of each corresponding slender scintillator crystal; and a timing signal circuit for each slender scintillator crystal, which is used to generate different timing signals for the scintillation photons detected by each individual photosensor of each corresponding slender scintillator crystal.

[0277] Item 69. A PET scanning system according to Item 64, wherein each elongated scintillator crystal is divided into a plurality of segmented volumes along its length, wherein a reflective layer is positioned between adjacent faces of the segmented volumes, wherein each segmented volume is associated with a different photosensor array so that scintillation photons generated within the segmented volume are reflected within the segmented volume to be detected by the corresponding photosensor array.

[0278] Item 70. A PET scanning system according to Item 64, further comprising: a high-speed circuit connected to the photosensor array of each elongated scintillator crystal of each detector module to digitally encode scintillation events detected within each elongated scintillator crystal into: a timing signal for each of the multiple photosensors associated with each elongated scintillator crystal; and an energy signal that sums the total energy detected by the multiple photosensors associated with each elongated scintillator crystal.

[0279] Clause 71. A PET scanning system as described in Clause 70, wherein the high speed circuit is a dual channel high speed circuit, wherein each channel includes circuitry to digitally encode timing signals and energy signals for a plurality of photosensors associated with each elongated scintillator crystal.

[0280] Clause 72. The PET scanning system of Clause 71, wherein each elongated scintillator crystal in the array of elongated scintillator crystals of each detector module uses a different high speed circuit than an adjacent elongated scintillator crystal.

[0281] Item 73. A gamma radiation detector module comprising: an array of elongated scintillation crystals positioned on a detector ring, wherein each elongated scintillation crystal comprises a proximal face, two axially oriented sides, two transversely oriented sides, and a distal face oriented radially into the detector ring to receive gamma photons; an array of photosensors having a detection face for detecting scintillation photons, wherein the detection face of the photosensor array is positioned along the length of the first of the axially oriented sides of each corresponding elongated scintillation crystal; and an inserter connected to an output face of the photosensor array along the length of the first axially oriented side of each corresponding elongated scintillation crystal, wherein the inserter is operable to transmit an electrical signal from the output face of the photosensor array of each corresponding elongated scintillation crystal toward the proximal face of each corresponding elongated scintillation crystal.

[0282] Item 74. A gamma radiation detector module according to Item 73, wherein each photosensor array of elongated scintillator crystals comprises a two-dimensional array of single photon avalanche diodes (SPADs), and wherein the inserter comprises a bias circuit to operate each corresponding two-dimensional array of SPADs of elongated scintillator crystals in Geiger mode.

[0283] Clause 75. The gamma radiation detector module of clause 73, wherein the fill factor of the distal face area of ​​the elongated scintillator crystal to the total distal face area (including the distal face of the photosensor array and the distal face of the inserter of each corresponding elongated scintillator crystal) is greater than 85%.

[0284] Clause 76. A gamma radiation detector module according to clause 73, comprising: reflective material positioned on the proximal face, the distal face, two transversely oriented side faces and the second axially oriented side face of each slender scintillator crystal, wherein the reflective material acts to internally reflect scintillation photons generated within each corresponding slender scintillator crystal.

[0285] Item 77. A gamma radiation detector module according to Item 76, wherein each elongated scintillator crystal is divided into a plurality of segmented volumes along its length, a reflective layer is positioned between adjacent faces of the segmented volumes, wherein each segmented volume of each elongated scintillator crystal is associated with a unique photosensor array so that scintillation photons generated within the segmented volume are reflected within the segmented volume to be detected by the corresponding photosensor array.

[0286] Clause 78. The gamma radiation detector module of Clause 76, wherein each elongated scintillator crystal is shaped as an elongated N-sided polygonal prism, where N is an integer value.

[0287] Clause 79. The gamma radiation detector module of Clause 76, wherein each elongated scintillator crystal is shaped as an elongated rectangular prism.

[0288] Clause 80. The gamma radiation detector module of Clause 79, wherein the distal face of each elongated scintillator crystal is square.

[0289] Clause 81. A gamma radiation detector module according to clause 80, wherein the length between the distal face and the proximal face of each slender scintillator crystal is between 10 millimeters and 30 millimeters, and wherein the ratio of the length of each slender scintillator crystal to the width of each slender scintillator crystal is between three and ten (3-10).

[0290] Clause 82. The gamma radiation detector module of Clause 76, wherein the array of elongated scintillator crystals comprises a one-dimensional array of elongated scintillator crystals arranged along an axial direction of the detector ring.

[0291] Item 83. A gamma radiation detector module according to Item 82, wherein the slender scintillator crystals are seamlessly positioned in the axial direction of the detector ring without any gaps therebetween, so that a given slender scintillator crystal and an axially adjacent slender scintillator crystal are separated from each other by: (i) a photosensor array on a first axially oriented side of a given slender scintillator crystal, (ii) an inserter of a given slender scintillator crystal, and (iii) a reflective material on a second axially oriented side of an axially adjacent slender scintillator crystal.

[0292] Item 84. A gamma radiation detector module according to Item 76, wherein the slender scintillation crystal array includes a two-dimensional array of slender scintillation crystals, wherein the two-dimensional array of slender scintillation crystals includes M slender scintillation crystals in the axial direction of the detector ring and N slender scintillation crystals in the transverse direction of the detector ring, wherein M and N are both integer values.

[0293] Item 85. A gamma radiation detector module according to Item 84, wherein the slender scintillation crystals are seamlessly positioned adjacent to each other in the axial and transverse directions of the detector ring without any gap therebetween, so that a given slender scintillation crystal and an axially adjacent slender scintillation crystal are separated from each other in the axial direction by: (i) a photosensor array on a first axially oriented side of a given slender scintillation crystal, (ii) an inserter of a given slender scintillation crystal and (iii) a reflective material on a second axially oriented side of an axially adjacent slender scintillation crystal, and so that a given slender scintillation crystal and a transversely adjacent slender scintillation crystal are separated from each other in the transverse direction by the reflective material.

[0294] Item 86. A gamma radiation detector module according to Item 85, wherein the thickness of the photosensor array and the reflective material between adjacent slender scintillator crystals in the axial direction of the detector ring is less than 500 microns, and wherein the thickness of the reflective material between adjacent slender scintillator crystals in the transverse direction of the detector ring is less than 100 microns.

[0295] Item 87. A gamma radiation detector module according to Item 86, wherein the length between the distal face and the proximal face of each slender scintillator crystal is between 10 mm and 30 mm, wherein the ratio of the length of each slender scintillator crystal to the axial width is between three and ten (3-10), and wherein the ratio of the length of each slender scintillator crystal to the transverse width is between three and ten (3-10).

[0296] Clause 88. The gamma radiation detector module of Clause 85, wherein a fill factor of an area of ​​the distal face of the elongated scintillator crystal to a total area is greater than 85%.

[0297] Clause 89. The gamma radiation detector module of Clause 76, wherein the reflective material comprises a thermally conductive reflective material.

[0298] Clause 90. The gamma radiation detector module of clause 89, wherein the thermally conductive reflective material comprises at least one of aluminum, magnesium, silver, stainless steel, and combinations thereof.

[0299] Clause 91. The gamma radiation detector module of Clause 89, further comprising a thermal conductor plate coupled to the proximal end faces of at least some of the elongated scintillator crystals in the array of elongated scintillator crystals.

[0300] Clause 92. The gamma radiation detector module of Clause 73, further comprising: a sealed housing enclosing the elongated scintillator crystal, the photosensor array, and the interposer.

[0301] Clause 93. A gamma radiation detector module as described in Clause 73, wherein the photosensor array of each elongated scintillator crystal includes a plurality of photosensors to provide different photosensor measurements corresponding to different depth ranges between the proximal and distal faces of each elongated scintillator crystal.

[0302] Clause 94. The gamma radiation detector module of Clause 93, wherein the photosensor array of each elongated scintillator crystal comprises a two-dimensional array of individual photodiodes.

[0303] Clause 95. The gamma radiation detector module of Clause 94, wherein each photodiode comprises a single photon avalanche diode (SPAD).

[0304] Clause 96. The gamma radiation detector module of clause 95, wherein each photosensor comprises a silicon photomultiplier (SiPM) such that each elongated scintillator crystal is associated with a plurality of SiPMs extending along a first axially-oriented side thereof.

[0305] Item 97. A circuit for processing nuclear scintillation events, comprising: a plurality of photosensors for detecting scintillation photons generated by scintillation events within a scintillation crystal and outputting electrical signals; an energy circuit for generating a digital signal representing the total energy detected by one or more of the plurality of photosensors during a detection time period; and a timing circuit for generating a digital signal having different timing signals for the scintillation photons detected by one or more photosensors during the detection time period.

[0306] Clause 98. A circuit according to clause 97, wherein the photosensor includes a plurality of single photon avalanche diodes (SPADs) biased to operate in Geiger mode, wherein the bias circuit is integrated within the interposer, and wherein each of the plurality of photosensors is positioned on a side of the scintillator crystal.

[0307] Item 99. A circuit as described in Item 97, wherein the energy circuit includes: an analog-to-digital converter; a multiplexer for selecting between the output signals of each of a plurality of photosensors for conversion by the analog-to-digital converter; a signal delay transmission line for connecting one or more of the plurality of photosensors to the multiplexer; and digital control logic for controlling the multiplexer's selection between the output signals of the plurality of photosensors.

[0308] Item 100. A circuit according to item 97, wherein the timing circuit includes: an analog-to-digital converter for high-speed digital encoding of the capacitive decoupled output of one or more of the plurality of photosensors; a transmission line for connecting the capacitive decoupled output of one or more of the plurality of photosensors to the analog-to-digital converter through an output diode; a trigger for selectively triggering the analog-to-digital converter to encode the capacitive decoupled output of one or more of the plurality of photosensors by biasing the output diode into an on state; and a comparator for comparing the capacitive decoupled output of one or more of the plurality of photosensors with one of a low voltage threshold and a low current threshold, wherein the output of the comparator selectively triggers the trigger.

[0309] Clause 101. A circuit as described in clause 100, wherein the output diode comprises a PIN diode characterized by having at least one of the following: a high energy bandgap semiconductor of at least 1.3 eV, a reverse bias junction capacitance of less than 150 femtofarads, a nominal carrier lifetime of less than 10 nanoseconds at an operating current of less than 20 milliamperes, and a forward bias on-resistance of less than 10 ohms at an operating frequency of 1 gigahertz.

[0310] Clause 102. The circuit of clause 100, wherein the output diode comprises a PIN diode having a high energy bandgap semiconductor of at least 1.3 eV.

[0311] Clause 103. The circuit of clause 100, wherein the output diode comprises a PIN diode having a reverse biased junction capacitance of less than 150 femtofarads.

[0312] Clause 104. The circuit of clause 100, wherein the output diode comprises a PIN diode having a nominal carrier lifetime of less than 10 nanoseconds at an operating current of less than 20 milliamperes.

[0313] Clause 105. The circuit of clause 100, wherein the output diode comprises a PIN diode having a forward biased on-resistance of less than 10 ohms at an operating frequency of 1 gigahertz.

[0314] Clause 106. A circuit according to clause 97, wherein the energy circuit includes: a first analog-to-digital converter; a multiplexer for selecting between the output signals of each of a plurality of photosensors for conversion by the first analog-to-digital converter; a signal delay transmission line for connecting each photosensor to the multiplexer; and digital control logic for controlling the multiplexer's selection between the output signals of the plurality of photosensors; and wherein the timing circuit includes: a second analog-to-digital converter for high-speed digital encoding of the capacitive decoupled outputs of one or more of the plurality of photosensors; a transmission line for connecting the capacitive decoupled outputs from one or more of the plurality of photosensors to the analog-to-digital converter via an output diode; a trigger for selectively triggering the second analog-to-digital converter to encode the capacitive decoupled outputs from one or more of the plurality of photosensors by biasing the output diode into a conducting state; and a comparator for comparing the output signals of one or more of the plurality of photosensors with a threshold, wherein the output of the comparator selectively triggers the trigger to bias the output diode into a conducting state.

[0315] Clause 107. The circuit of clause 106, wherein the comparator comprises a voltage comparator for comparing a voltage level of an output signal of one or more of the plurality of photosensors to a voltage threshold.

[0316] Clause 108. The circuit of clause 106, wherein the comparator comprises a current comparator to compare a current level of an output signal of one or more of the plurality of photosensors to a current threshold value.

[0317] Clause 109. The circuit of clause 106, wherein the output diode comprises a PIN diode having a high energy bandgap semiconductor of at least 1.3 eV.

[0318] Clause 110. The circuit of clause 106, wherein the output diode comprises a PIN diode having a reverse biased junction capacitance of less than 150 femtofarads.

[0319] Clause 111. The circuit of clause 106, wherein the output diode comprises a PIN diode having a nominal carrier lifetime of less than 10 nanoseconds at an operating current of less than 20 milliamperes.

[0320] Clause 112. The circuit of clause 106, wherein the output diode comprises a PIN diode having a forward biased on-resistance of less than 10 ohms at an operating frequency of 1 gigahertz.

[0321] Item 113. A circuit as described in item 106, wherein the output diode comprises a PIN diode characterized by having: a high energy bandgap semiconductor of at least 1.3 eV, a reverse bias junction capacitance of less than 150 femtofarads, a nominal carrier lifetime of less than 10 nanoseconds at an operating current of less than 20 milliamperes, and a forward bias on-resistance of less than 10 ohms at an operating frequency of one gigahertz.

[0322] The present disclosure should be considered illustrative rather than restrictive, and all such modifications are intended to be included within its scope. Likewise, benefits, other advantages, and solutions to problems have been described above for various embodiments. However, benefits, advantages, solutions to problems, and any elements that may cause any benefit, advantage, or solution to occur or become more apparent should not be construed as critical, required, or essential features or elements. The present disclosure includes and encompasses at least the claims set forth.

Claims

1. A gamma radiation detector module, comprising: an array of scintillation crystals positioned on the detector, wherein each scintillation crystal comprises a proximal face, a distal face oriented into the detector to receive gamma photons, and four sides comprising a first side, a second side, a third side, and a fourth side; an array of photosensors positioned along said first side of each scintillation crystal to detect scintillation photons; and A reflective material is positioned on the proximal face, the distal face, and the second, third, and fourth sides of each scintillator crystal to internally reflect scintillation photons.

2. The gamma radiation detector module of claim 1 , wherein each scintillation crystal comprises an elongated scintillation crystal, wherein the array of elongated scintillation crystals is configured to be positioned on a detector ring, wherein the four sides of each scintillation crystal include two axially oriented sides and two transversely oriented sides, wherein the distal face is oriented radially into the detector ring to receive the gamma photons, and The first side is axially oriented so that the photoelectric sensor is positioned along the axially oriented side.

3. The gamma radiation detector module according to claim 2, wherein the slender scintillator crystal array comprises a two-dimensional array of slender scintillator crystals, wherein the two-dimensional array of slender scintillator crystals comprises M slender scintillator crystals in the axial direction of the detector ring and N slender scintillator crystals in the transverse direction of the detector ring, wherein M and N are both integer values.

4. The gamma radiation detector module of claim 3, wherein the elongated scintillation crystals are seamlessly positioned adjacent to each other in the axial direction and the transverse direction of the detector ring without gaps therebetween, such that a given elongated scintillator crystal is separated from an axially adjacent elongated scintillator crystal by: (i) the photosensor on the first axially oriented side of the given elongated scintillator crystal and (ii) the reflective material on the second axially oriented side of the axially adjacent elongated scintillator crystal, and The given elongated scintillator crystal and the elongated scintillator crystals adjacent to each other in the transverse axis are separated from each other by the reflective material in the transverse axis direction.

5. The gamma radiation detector module of claim 4, wherein the thickness of the photosensors and reflective materials between adjacent elongated scintillator crystals in the axial direction of the detector ring is less than 500 microns, and The thickness of the reflective material between adjacent elongated scintillator crystals in the transverse direction of the detector ring is less than 100 microns.

6. The gamma radiation detector module of any one of claims 1-5, wherein the reflective material comprises a thermally conductive reflective material.

7. The gamma radiation detector module of claim 2, further comprising: an interposer connected to said photosensor along said first axially-oriented side of each respective elongated scintillator crystal, wherein the interposer is operable to transmit an electrical signal from the photosensor of each respective elongated scintillator crystal toward the proximal face of each respective elongated scintillator crystal.

8. The gamma radiation detector module of claim 7, wherein the elongated scintillator crystal array comprises a two-dimensional array of elongated scintillator crystals, The two-dimensional array of elongated scintillator crystals includes M elongated scintillator crystals in the axial direction of the detector ring and N elongated scintillator crystals in the transverse direction of the detector ring, wherein M and N are both integer values.

9. The gamma radiation detector module of claim 8, wherein the elongated scintillation crystals are seamlessly positioned adjacent to each other in the axial direction and the transverse direction of the detector ring without gaps therebetween, such that a given elongated scintillator crystal is separated from an axially adjacent elongated scintillator crystal by: (i) the photosensor on the first axially oriented side of the given elongated scintillator crystal and (ii) the reflective material on the second axially oriented side of the axially adjacent elongated scintillator crystal, and The given elongated scintillator crystal and the elongated scintillator crystals adjacent to each other in the transverse axis are separated from each other by the reflective material in the transverse axis direction.

10. A gamma radiation detector module according to any one of claims 1-5, wherein the photosensor array of each scintillator crystal comprises at least one two-dimensional array of individual photodiodes.

11. The gamma radiation detector module of claim 10, wherein each photodiode comprises a single photon avalanche diode (SPAD) operating in Geiger mode.

12. A gamma radiation detector module according to any one of claims 1-5, wherein the photosensor array of each scintillator crystal includes a plurality of discrete photosensors to provide different photosensor measurements corresponding to different depth ranges between the proximal face and the distal face of each scintillator crystal.

13. The gamma radiation detector module of claim 12, wherein the photosensor array of each scintillator crystal comprises a plurality of two-dimensional arrays of individual photodiodes, and Each of the photodiodes comprises a single photon avalanche diode (SPAD) operating in Geiger mode.

14. The gamma radiation detector module of claim 13, wherein each photosensor comprises a silicon photomultiplier (SiPM), such that each scintillator crystal is associated with a plurality of SiPMs extending along the first side thereof.

15. The gamma radiation detector module of claim 14, further comprising a processing circuit, the processing circuit comprising: a plurality of energy circuits, wherein each energy circuit is configured to generate an energy signal that is a sum of energies detected by one or more silicon photomultiplier tubes; as well as A plurality of timing signal circuits, wherein each timing signal circuit is configured to generate a different timing signal for the scintillation photons detected by one or more silicon photomultiplier tubes.

16. The gamma radiation detector module of claim 15, wherein the processing circuit is a dual channel high speed circuit, wherein each channel includes circuitry for digitally encoding the timing signal and the energy signal of one or more of the silicon photomultiplier tubes, and Each scintillation crystal in the scintillation crystal array uses a different high speed circuit than an adjacent scintillation crystal.

17. A positron emission tomography (PET) scanning system comprising: a plurality of gamma radiation detector modules positioned to form a detector ring, wherein each detector module comprises: an array of elongated scintillation crystals, each scintillation crystal comprising a proximal end face, two axially oriented side faces, two transversely oriented side faces, and a distal end face positioned radially into the detector ring to receive gamma photons, an array of photosensors positioned along a first said axially oriented side of each elongated scintillation crystal to detect scintillation photons, and a reflective material positioned on the proximal face, the distal face, the radial side face and the second of the axially oriented sides of each elongated scintillator crystal to internally reflect scintillation photons; a cooling system for cooling the detector module; and An imaging system is coupled to the detector module to generate an image based on the electronic output from the detector module.

18. The PET scanning system according to claim 17, further comprising: an interposer coupled to the photosensor array along the first axially-oriented side of each respective elongated scintillator crystal of each detector module, wherein the interposer is operable to transmit electrical signals from the photosensor array of each respective elongated scintillator crystal toward the proximal face of each respective elongated scintillator crystal.

19. The PET scanning system according to claim 18, wherein: The slender scintillation crystal array of each detector module includes a two-dimensional array of slender scintillation crystals, wherein the two-dimensional array of slender scintillation crystals of each detector module includes M slender scintillation crystals in the axial direction of the detector ring and N slender scintillation crystals in the transverse direction of the detector ring, wherein M and N are both integer values.

20. The PET scanning system according to claim 19, wherein: The elongated scintillation crystals of each detector module are positioned adjacent to each other seamlessly in the axial direction and the transverse direction of the detector ring without gaps therebetween, such that a given elongated scintillator crystal and an axially adjacent elongated scintillator crystal are separated from each other in the axial direction by: (i) the photosensor array on the first axially oriented side of the given elongated scintillator crystal, (ii) the inserter of the given elongated scintillator crystal, and (iii) the reflective material on the second axially oriented side of the axially adjacent elongated scintillator crystal, and The given elongated scintillator crystal and the elongated scintillator crystals adjacent to each other in the transverse axis are separated from each other by the reflective material in the transverse axis direction.

21. A gamma radiation detector module comprising: an array of elongated scintillator crystals positioned on a detector ring, wherein each elongated scintillator crystal comprises a proximal end face, two axially oriented side faces, two transversely oriented side faces, and a distal end face oriented radially into the detector ring to receive gamma photons; a photosensor array having a detection surface for detecting scintillation photons, wherein said detection surface of said photosensor array is positioned along the length of a first said axially-oriented side of each respective elongated scintillation crystal; as well as an interposer connected to the output face of the photosensor array along the length of the first axially-oriented side of each respective elongated scintillator crystal, wherein the interposer is operable to transmit electrical signals from the output face of the photosensor array of each respective elongated scintillator crystal toward the proximal face of each respective elongated scintillator crystal.

22. The gamma radiation detector module of claim 21, wherein the photosensor array of each elongated scintillator crystal comprises a two-dimensional array of single photon avalanche diodes (SPADs), and wherein the interposer includes bias circuitry to operate the two-dimensional array of SPADs of each corresponding elongated scintillator crystal in Geiger mode.

23. The gamma radiation detector module according to claim 21 or 22, comprising: A reflective material is positioned on the proximal face, the distal face, the two transverse-axis oriented sides and the second of the axially oriented sides of each elongated scintillator crystal, wherein the reflective material operates to internally reflect scintillation photons generated within each respective elongated scintillator crystal.

24. The gamma radiation detector module of claim 23, wherein the elongated scintillator crystal array comprises a one-dimensional array of elongated scintillator crystals arranged along an axial direction of the detector ring, and wherein the elongated scintillator crystals are seamlessly positioned in the axial direction of the detector ring without gaps therebetween, such that a given elongated scintillator crystal and an axially adjacent elongated scintillator crystal are separated from each other by: (i) the photosensor array on the first axially oriented side of the given elongated scintillator crystal, (ii) the inserter of the given elongated scintillator crystal and (iii) the reflective material on the second axially oriented side of the axially adjacent elongated scintillator crystal.

25. A gamma radiation detector module according to claim 23, wherein the slender scintillator crystal array comprises a two-dimensional array of slender scintillator crystals, wherein the two-dimensional array of slender scintillator crystals comprises M slender scintillator crystals in the axial direction of the detector ring and N slender scintillator crystals in the transverse direction of the detector ring, wherein M and N are both integer values.

26. The gamma radiation detector module of claim 25, wherein the elongated scintillator crystals are positioned adjacent to each other seamlessly in the axial direction and the transverse direction of the detector ring without gaps therebetween, such that a given elongated scintillator crystal and an axially adjacent elongated scintillator crystal are separated from each other in the axial direction by: (i) the photosensor array on the first axially oriented side of the given elongated scintillator crystal, (ii) the inserter of the given elongated scintillator crystal, and (iii) the reflective material on the second axially oriented side of the axially adjacent elongated scintillator crystal, and The given elongated scintillator crystal and the elongated scintillator crystals adjacent to each other in the transverse axis are separated from each other by the reflective material in the transverse axis direction.

27. A circuit for processing a nuclear scintillation event, comprising: A plurality of photoelectric sensors for detecting scintillation photons generated by scintillation events in the scintillation crystal and outputting electrical signals; an energy circuit for generating a digital signal representing a total energy detected by one or more photosensors of the plurality of photosensors during a detection time period; as well as A timing circuit is used to generate a digital signal of different timing signals of the scintillation photons detected by the one or more photosensors during the detection time period.

28. The circuit of claim 27, wherein the photosensors comprise a plurality of single photon avalanche diodes (SPADs) biased to operate in Geiger mode, wherein the bias circuit is integrated within an interposer, and wherein each of the plurality of photosensors is positioned on a side of the scintillator crystal.

29. The circuit of claim 27, wherein the energy circuit comprises: Analog-to-digital converters; a multiplexer for selecting between output signals of each of the plurality of photosensors for conversion by the analog-to-digital converter; a signal delay transmission line for connecting one or more of the plurality of photosensors to the multiplexer; as well as Digital control logic for controlling the multiplexer to select between the output signals of the plurality of photosensors.

30. The circuit of claim 27, wherein the timing circuit comprises: an analog-to-digital converter for high-speed digital encoding of the capacitively decoupled output of one or more of the plurality of photosensors; a transmission line for connecting the capacitively decoupled output of one or more of the plurality of photosensors to the analog-to-digital converter via an output diode; a trigger for selectively triggering the analog-to-digital converter to encode the capacitively decoupled output of one or more of the plurality of photosensors by biasing the output diode into a conducting state; as well as A comparator is configured to compare the capacitive decoupling output of one or more of the plurality of photosensors with one of a low voltage threshold and a low current threshold, wherein the output of the comparator selectively triggers the trigger.

31. The circuit of claim 30, wherein the output diode comprises a PIN diode, characterized in that At least one of the following: High bandgap semiconductors of at least 1.3 eV, less than 150 femtofarads of reverse-biased junction capacitance, A nominal carrier lifetime of less than 10 nanoseconds at an operating current of less than 20 mA, and Less than 10 ohms forward bias on-resistance at 1 GHz operating frequency.

32. The circuit of claim 27, wherein the energy circuit comprises: a first analog-to-digital converter; a multiplexer for selecting between output signals of each of the plurality of photosensors for conversion by the first analog-to-digital converter; a signal delay transmission line for connecting each photosensor to the multiplexer; and digital control logic for controlling the selection of said multiplexer between said output signals of said plurality of photosensors; and Wherein, the timing circuit comprises: a second analog-to-digital converter for high-speed digital encoding of the capacitively decoupled output of one or more of the plurality of photosensors; a transmission line for connecting the capacitively decoupled output from one or more of the plurality of photosensors to the analog-to-digital converter via an output diode; a trigger for selectively triggering the second analog-to-digital converter to encode the capacitively decoupled output from one or more of the plurality of photosensors by biasing the output diode into a conducting state; and A comparator is configured to compare the output signal of one or more of the plurality of photosensors with a threshold value, wherein the output of the comparator selectively triggers the trigger to bias the output diode into a conducting state.

33. The circuit of claim 32, wherein the comparator comprises a voltage comparator for comparing a voltage level of the output signal of one or more of the plurality of photosensors with a voltage threshold.

34. The circuit of claim 32, wherein the comparator comprises a current comparator for comparing a current level of the output signal of one or more of the plurality of photosensors with a current threshold.

35. The circuit of claim 32, wherein the output diode comprises a PIN diode comprising a high energy bandgap semiconductor of at least 1.3 eV.

36. The circuit of claim 32, wherein the output diode comprises a PIN diode having a reverse biased junction capacitance of less than 150 femtofarads.

37. The circuit of claim 32, wherein the output diode comprises a PIN diode having a nominal carrier lifetime of less than 10 nanoseconds at an operating current of less than 20 milliamperes.

38. The circuit of claim 32, wherein the output diode comprises a PIN diode having a forward biased on-resistance of less than 10 ohms at an operating frequency of 1 GHz.

39. The circuit of claim 32, wherein the output diode comprises a PIN diode, characterized in that have: A high energy bandgap semiconductor of at least 1.3 eV, a reverse bias junction capacitance of less than 150 femtofarads, a nominal carrier lifetime of less than 10 nanoseconds at an operating current of less than 20 milliamps, and a forward bias on-resistance of less than 10 ohms at an operating frequency of 1 gigahertz.

Images