Lateral crystal photodiode readout and switch diode network for processing nuclear events.

Abstract

A positron emission tomography (PET) scanner includes multiple gamma radiation detector modules arranged to form a detector ring. Each detector module includes an array of elongated scintillation crystals. With respect to the detector ring, each elongated scintillation crystal includes a proximal end face, two axially-oriented side faces, two cross-axially-oriented side faces, and a distal end face oriented radially into the detector ring to receive gamma photons. An array of photosensors is disposed along a first of the axially-oriented side faces of each elongated scintillation crystal to detect scintillation photons. Reflective material is disposed on a second of the proximal end face, the distal end face, the cross-axially-oriented side face, and the axially-oriented side face of each elongated scintillation crystal to internally reflect scintillation photons. In various embodiments, a dual-channel processing circuit provides separate timing and energy signals from the photosensors.

Description

Detailed Description of the Invention

[0001] (CROSS-REFERENCE TO RELATED APPLICATIONS) This application claims priority to U.S. patent application Ser. No. 18 / 365,755, filed Aug. 4, 2023, entitled "Lateral Crystal Photodiode Readouts and Switched Diode Networks for Processing Nuclear Events," and claims priority and the benefit under 35 U.S.C. § 119 of U.S. Provisional Patent Application No. 63 / 378,739. This application claims priority to and the benefit under 35 U.S.C. § 119 of U.S. Provisional Patent Application No. 63 / 378,739, filed 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 May 14, 2023, also entitled "Lateral Crystal Photodiode Readout and Switched Diode Network for Processing Nuclear Events," each of which is incorporated herein by reference in its entirety.

[0002] (background) The present disclosure relates to gamma radiation measurement, circuits for scintillation detectors, and coincidence circuit arrangements. The systems, methods, practical applications, and uses of the embodiments described in this disclosure can be understood in the context of the following publications, each of which is incorporated herein 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,Inc.,2000;S. Gundacker, E. Auffray, NDVara, 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,vol.718,pp.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,p.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: https: / / www.crystals.saint-gobain.com / radiation-detection-scintillators / crystal-scintillators / lyso-scintillation-crystals. BRIEF DESCRIPTION OF THE DRAWINGS FIG. 1A is a perspective view of a block diagram of a portion of a positron emission tomography (PET) detector system according to one embodiment.

[0003] FIG. 1B is another illustration of a block diagram of a portion of a PET detector system according to one embodiment.

[0004] FIG. 2 is a diagram illustrating a positron source having multiple lines of response, according to one embodiment.

[0005] FIG. 3 shows a tube-type block detector with four photomultiplier tubes, according to one embodiment.

[0006] FIG. 4 is a block diagram of a ring of detector modules having multiple response lines, according to one embodiment.

[0007] FIG. 5A is a block diagram of a ring of detector modules with a single off-center line of response (LoR), according to one embodiment.

[0008] FIG. 5B is a block diagram of a ring of detector modules having multiple off-center response lines, according to one embodiment.

[0009] FIG. 6 is a diagram illustrating response lines segmented in time and space according to one embodiment.

[0010] FIG. 7 is a simplified circuit diagram of a high speed diode network according to one embodiment.

[0011] FIG. 8A is a two-port network lumped element model of a conductive transmission line, according to one embodiment.

[0012] FIG. 8B illustrates a distribution model of lumped elements per unit length, according to one embodiment.

[0013] FIG. 9 shows a time domain graph of two nuclear pulse signals, according to one embodiment.

[0014] FIG. 10 shows a graph of the energy spectral density of the two nuclear pulse signals of FIG. 9, according to one embodiment.

[0015] FIG. 11 is a hierarchical block diagram of a high speed diode network according to one embodiment.

[0016] FIG. 12 shows a detector module having blocks of elongated scintillation crystals arranged in a two-dimensional array with an array of photosensors for lateral photosensor readout, according to one embodiment.

[0017] FIG. 13A is a diagram of a single elongated scintillation crystal pixel with multiple arrays of photosensors for lateral photosensor readout, according to one embodiment.

[0018] FIG. 13B is an exploded view of a single elongated scintillation crystal according to one embodiment.

[0019] FIG. 13C is a diagram of a single elongated scintillation crystal pixel with an array of photosensors for side photosensor readout, a middle substrate, and a connector array, according to one embodiment.

[0020] FIG. 14A shows a subdivided elongated scintillation crystal pixel having an array of photosensors for lateral photosensor readout, according to one embodiment.

[0021] FIG. 14B shows an exploded view of a subdivided scintillation pixel, according to one embodiment.

[0022] FIG. 14C is a diagram of a sub-divided scintillation pixel having an array of photosensors for side photosensor readout, a middle substrate, and a connector array, according to one embodiment.

[0023] FIG. 15A is a diagram of a block of an elongated scintillation crystal with a side-reading photosensor and a thermal management plate, according to one embodiment.

[0024] FIG. 15B is a diagram of a block of an elongated scintillation crystal with an array of photosensors, an intermediate substrate, and a connector, according to one embodiment.

[0025] FIG. 15C is a diagram of a block of elongated scintillation crystal with an array of photosensors, an intermediate substrate, a connector array, and a thermal management plate, according to one embodiment.

[0026] FIG. 15D is a diagram of a block of elongated scintillation crystals connected to a dual-channel processing circuit, according to one embodiment.

[0027] FIG. 16 is a diagram of a block of elongated scintillation crystals having a double-sided intermediate substrate connected to the photosensor arrays of adjacent elongated scintillation crystals, according to one embodiment.

[0028] FIG. 17 illustrates a ring of detector modules with axially oriented photosensor arrays flanking an elongated scintillation crystal, according to one embodiment.

[0029] FIG. 18 shows a ring of detector modules having a transverse array of photosensors on the sides of an elongated scintillation crystal, according to one embodiment.

[0030] FIG. 19 is a graph of sensitivity gain and fill factor loss for side photosensor readout, according to one embodiment.

[0031] FIG. 20 is a graph of the overall sensitivity gain for square-ended rectangular prism scintillation crystals of various lengths, according to one embodiment.

[0032] FIG. 21 is a graph of the photonic dynamic range of a silicon photomultiplier photosensor for various dimensions of an elongated rectangular prism scintillation crystal, according to one embodiment.

[0033] FIG. 22 is a graph of scintillator light output versus temperature, according to one embodiment.

[0034] (Detailed explanation) The system and method embodiments provided within this disclosure are not intended to limit the scope of the disclosure but are merely representative of possible embodiments. Additionally, the steps of the methods do not necessarily have to be performed in a particular order, nor sequentially, nor do the steps have to be performed only once. The following description, in combination with the figures, sets forth numerous specific details to provide a thorough understanding of possible variations of the systems and methods. However, the disclosed concepts, systems, methods, apparatus, etc., can be practiced without some or all of the specific details. For clarity, technical materials known in the technical fields relevant to this disclosure will not be described in detail.

[0035] Positron emission tomography (PET) is a medical imaging system incorporating multiple gamma radiation detector modules, which can be arranged in a multi-ring configuration for the detection of radiation-emitting materials placed within the scanner's field of view and subsequent image reconstruction. Human medical imaging using PET is facilitated by the intravenous injection of a positron-based radioactive tracer with sufficient emission duration to enable PET scanning. An example of such a tracer is a fluorine-18 labeled glucose compound, which is used to identify highly metabolic tumor activity. When a positron encounters an electron, an annihilation event occurs, resulting in the emission of two opposing gamma photons, approximately 180° apart and each with an energy of approximately 511 kiloelectron volts (keV).

[0036] The range a positron travels before encountering a free electron is called the positron range. Depending on the initial kinetic energy of the positron and the electron density in the medium, the in vivo range is typically less than 1 mm for fluorine-18-based radiotracers. The positron's kinetic energy also introduces undesirable non-collinearity effects, causing annihilation gamma photons to traverse more or less from 180°, resulting in an image error of approximately 1.8 mm for a typical scanner bore size. The positron range and errors caused by non-collinearity limit the resolution of some PET whole-body imaging to approximately 2 mm.

[0037] 1A is a perspective view of a PET detector system 100 and positron sources 150 and 151 within a patient 110 (shown as a human torso), according to one embodiment. 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 the viewer, a second ring 180 of gamma radiation detector modules is in the center, and a third ring 190 of gamma radiation detector modules is furthest from the viewer. In the illustrated example, each ring 170, 180, 190 includes 48 gamma radiation detector modules. The number of detector modules and the number of rings of detector modules may vary based on the particular application, cost considerations, target size, target resolution, etc.

[0038] Positron sources 150 and 151 represent a bolus of positron activity resulting in annihilation radiation in which two opposing gamma photons are emitted in opposite directions. The opposing gamma photons emitted by positron source 150 are detected by detector modules 183 and 187. A processor or other circuitry can use the known detector positions and measured arrival times of the gamma photons detected by detector modules 183 and 187 to calculate the line of response (LoR) 130 and determine the location of positron source 150. The processor or other circuitry can identify the detected gamma photons as significant annihilation events based on the total energy detected being equal to the expected 511 keV or within a threshold range.

[0039] Opposing gamma photons emitted by positron source 151 are detected by detector modules 181 and 185. A processor or other circuitry can use the known detector positions and the measured arrival times of the gamma photons detected by detector modules 181 and 185 to calculate line of response 131 and determine the position of positron source 151. As will be described in more detail below, if positron sources 150 and 151 emit their respective pairs of opposing gamma photons simultaneously (or nearly simultaneously), the processor or other circuitry may erroneously calculate line of response 135.

[0040] Throughout this disclosure, detector modules are described and illustrated as being used within a detector ring (e.g., a ring of detector modules). The scintillation crystal of each detector module has a distal end face facing into the detector (e.g., into the ring, with the distal end face located on the inner diameter of the ring) to receive gamma photons. However, it is understood that a detector ring may include ring-shaped detector modules without forming a complete or complete ring (e.g., only the top and bottom detector modules are in place). Furthermore, a detector may be embodied as two (or more) planar panels of detector modules that cooperate to detect annihilation events. For example, a detector may include a first planar panel of a detector module and a second, opposing planar panel of a detector module. An object of interest may be placed within the detector by being positioned between the two opposing planar panels. In this manner, a detector module may include multiple scintillation crystals (e.g., a one-dimensional or two-dimensional array) with their distal ends oriented toward the detector (e.g., within the cavity or space between the opposing panels, with the distal end face located on the inner diameter of the ring).

[0041] 1B is a side view of another example of PET detector system 101 and positron source 153 within human 111, according to one embodiment. In the illustrated embodiment, PET detector system 101 includes a ring having 72 gamma radiation detector modules. Opposing gamma photons emitted by positron source 153 are detected by detector modules 182 and 184. A processor or other circuitry can use the positions of the detector modules and the arrival times of the gamma photons detected by detector modules 182 and 184 to calculate line of response 137 and determine the position of positron source 153.

[0042] 2 is a block diagram of a PET detector system 200 according to one embodiment. The simplified block diagram includes a single ring of 16 gamma radiation detector modules 210-225. According to one embodiment, multiple response lines 231, 232, and 233 are shown intersecting a bolus of positron activity 250. Annihilation radiation from the positron bolus 250 generates multiple simultaneous emissions of equal and opposite gamma photons. System 200 can determine the range of origin of a positron along response line 233, which extends between two of its detection points (e.g., within detector modules 223 and 215).

[0043] Additional lines of response 231 and 232 are calculated during measurement periods during which subsequent annihilation events occur within the bolus 250. The period during which annihilation events are detected from the bolus 250 is called a scan. System 200 is used to acquire scans that provide multiple lines of response for identifying the origin or region of positron material. A PET scanner (e.g., including or embodied as a PET detector system described herein) can use rings of detector modules 210-225 to develop images of varying concentrations and locations of positron material without the use of shielded collimators, such as those employed in single-photon emission computed tomography (SPECT), which filter oblique origin angles. In other embodiments, PET scanners may utilize axial collimators, called septa, between multiple detector rings to reduce scattered radiation emanating from the field of view (e.g., within the human body). However, other PET scanners operate in a "3D" mode without septa due to faster scintillators and advances in electronic processing. This allows for larger event rates and improves system sensitivity.

[0044] Annihilation double gamma radiation events are distinguished (e.g., relatively distinguished) from numerous other unrelated gamma radiation events (commonly referred to as "single events") by using a time-based coincidence window. The coincidence window, which may be less than 5 nanoseconds (e.g., 2 nanoseconds), is used to filter single events that could not have arisen from a corresponding single annihilation within the field of view, given the scanner's bore diameter (corresponding to the diameter of the ring of detector modules 210-225) and the propagation speed of gamma radiation. For example, in a PET scanner with detector modules in a 70-centimeter diameter ring (e.g., a 70-centimeter bore diameter), a positron annihilation at the center of the bore requires 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 of the detector modules in the ring. Correspondingly, if the positron annihilations are located along the circumference of the bore, it would take twice as long for one of the gamma photons to scintillate into a detector module on the opposite side of the ring. Thus, together with a selected pairing of detector modules consisting of only reasonable geometric lines of response, the coincidence window reduces uncorrelated gamma radiation events that would otherwise be recorded as coincidences.

[0045] If two gamma radiation events meet the criteria to be recorded as a single coincident event (e.g., detected by the appropriate detector module within a time window), such an event is called a "prompt" event. Prompt events are two uncorrelated, independent gamma radiation events that originate from different annihilation events or other radiation sources, and are called "random" coincidence events. However, if two gamma radiation detections truly result from a single positron annihilation, such an event is called a "true" coincidence event. Therefore, the number of "true" coincidence events resulting from a true positron annihilation is equal to the number of "prompt" coincidence events (e.g., coincidence events from an unknown source) minus the number of "random" coincidence events (coincidence events known or statistically determined to be random singles measured within the coincidence window).

[0046] In various embodiments, PET scanners may further filter "scatter" events. This occurs when a gamma photon deposits some of its energy in nearby material and propagates the remaining energy toward the detector module. Such scatter events are typically filtered because their origin is unclear given their possible deflection. Scatter events typically arrive at the detector with significantly less energy than the original 511 keV. Therefore, filtering involves measuring the amount of energy deposited by a gamma radiation event and rejecting such events below a low-level discrimination (LLD) threshold. Additionally, upper-level discrimination (ULD) may be utilized to reject the possibility of near-simultaneous deposition of multiple events on the detector or the deposition of a single high-energy nuclear particle. According to various embodiments, the PET scanner may utilize an LLD threshold between 325 keV and 511 keV and a ULD threshold between 511 keV and 675 keV.

[0047] 3 illustrates a tube-based detector module 300 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 of elongated scintillation crystals 310. An exemplary two-dimensional array of elongated scintillation crystals 310 includes a 7x8 array of 56 elongated scintillation crystals. Some embodiments of gamma detector modules for PET scanners use an inorganic scintillator called LYSO (e.g., bismuth germanate, called BGO, Bi4Ge3O4) coupled to a light guide and connected to the quadrants of the photomultiplier tubes. 12 , or LSO, Lu2(SiO4)O, Lu 2(1-x-y) Y 2x SiO5, etc.

[0048] The illustrated configuration is sometimes referred to as a block detector module. A detector module may be composed of one or more block detectors. In this case, the block detectors share a channel of electronics for processing single nuclear gamma radiation events from multiple elongated scintillation crystals 310. The elongated scintillation crystals 310 may be pixelated into multiple elongated scintillation crystals 310, as shown. The elongated scintillation crystals 310 may be etched, polished, and / or encased in a reflective material. Pixelating the elongated scintillation crystals 310 into an array of elongated rectangular prisms facilitates determining the position of incident gamma photons within a rectangular area at the end face of each elongated scintillation crystal 310.

[0049] In various embodiments, selected areas or faces of each elongated scintillation crystal 310 can be wrapped or coated with a reflective material. By utilizing four photomultiplier tubes 320, the collection of light output from any crystal can be measured ratiometrically to determine which crystals have undergone gamma radiation interactions, thereby allowing such events to be located on the XY plane of the detector plane.

[0050] Sharing single-event processing electronics across blocks is motivated by cost limitations, reduced power consumption and dissipation, and compact packaging volume. In a block processing configuration, the detector module 300 processes a single event exclusively through the block's processing electronics, and other incident radiation within the same block's processing time is typically rejected. This processing time is called detector dead time. A defined block detection area is a consideration for PET, since sharing processing electronics across too large a block can reduce count rate capability and sensitivity, while a block that is too small can be cost prohibitive and involve complex thermal considerations relative to a practical volume.

[0051] In various embodiments, solid-state photosensors (e.g., silicon-based photosensors) can be used instead of photomultiplier tubes. Traditional semiconductor-based PIN photodiodes and / or avalanche photodiodes (APDs) generate analog signals proportional to the incident light flux. However, APDs experience gain through impact ionization at high reverse bias, where the generated electron-hole pairs collide with nearby atoms, liberating additional charge carriers beyond the initially generated photocurrent. APDs typically have much better sensitivity but generally suffer from higher noise levels than PIN photodiodes. Recent developments have resulted in single-photon avalanche photodiodes (SPADs), which operate at very high reverse bias, commonly referred to as Geiger mode. SPADs are optimized for single-photon detection and undergo self-sustained avalanche breakdown, which must be quenched during operation. As a result of this self-sustained breakdown, SPADs do not provide information about light intensity like APDs or PIN photodiodes. However, SPADs can be configured in parallel with many other SPADs, collectively referred to as silicon photomultiplier (SiPM) photosensors. The SiPM is a unique photosensor that generates a signal proportional to the light intensity resulting from a corresponding number of triggered SPAD cells. In this way, the SiPM acts as an analog photon counter, with each individually triggered SPAD contributing to the total photocurrent output of the sensor, resulting in a much higher gain than a single APD or PIN photodiode.

[0052] SiPM photosensors can contain tens, hundreds, or even thousands of single-photon avalanche photodiodes per square millimeter on a single silicon substrate. Silicon photomultipliers are used when both timing resolution and photonic dynamic range are required, desired, and / or deemed useful for a particular application or use case. Silicon photomultipliers offer advantages over conventional photomultiplier tubes, including reduced packaging volume, reduced bias voltage, reduced sensitivity to magnetic fields, and improved single-photon timing resolution. Disadvantages of silicon photomultiplier tubes over conventional photomultiplier tubes include higher dark current (counts) for a given temperature and area and complex multi-exponential pulse shapes. Solid-state photosensors can be used one-to-one in place of photomultiplier tubes in a block, or they can be coupled one-to-one to an elongated scintillation crystal in a block.

[0053] As used herein, references to "photosensors," "photodiodes," and / or "photodetectors" are not meant to be limiting or exclusive to a particular embodiment unless expressly stated otherwise. The term "readout," as used herein, refers to the measurement, investigation, evaluation, or processing of the referenced topic. Similarly, the use of "lateral" readout, "side" readout, and "side" sensing all generally refer to light sensing of a scintillator in which the conversion surface of the photosensor is coupled to a scintillator surface parallel to the lateral axis of the scanner bore, and therefore is not imaged from either end surface parallel to the axis of the scanner bore. Additionally, as used herein, "scintillator," "crystal," "pixel," and "block" may each refer to scintillator materials configured in different forms.

[0054] 4 is a block diagram of a PET scanner 400 having a ring-shaped detector module 401 and multiple lines of response 471, 472, and 473, according to one embodiment. Positron annihilation 461 from the center of the ring results in line of response 471 impinging orthogonally on the end face of the elongated scintillation crystal of a pair of detector modules. Positron annihilations occurring off-center, such as positron annihilations 462 and 463, result in gamma photons impinging on the scintillator at oblique angles, resulting in a parallax effect. As shown in the exploded view of the elongated scintillation crystal 410 of detector module 401 (top right corner of the drawing), two gamma radiation events 481 and 482 scintillate within the same elongated scintillation crystal 410 but originate from separate locations within the patient (positron annihilations 462 and 463, respectively). An end-face photosensor 450 at the proximal end face of the elongated scintillation crystal 410 detects scintillation photons produced by the elongated scintillation crystal 410 in response to received gamma photons.

[0055] True response lines 472 and 473 for annihilation events 462 and 463 are shown. However, these two gamma radiation 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 error. Such parallax effects introduce geometric uncertainty, further including the possibility of random events, given that some embodiments of the detector module 401 cannot distinguish between events perpendicular to the face of the elongated scintillation crystal 451 and events arising from severe oblique angles. Due to the location of the end photosensor 450 on the proximal end face of the elongated scintillation crystal 451, the end photosensor 450 can detect scintillation photons generated within the elongated scintillation crystal 451, but does not have the ability to determine the depth of interaction. For example, the end photosensor 450 does not distinguish between scintillation events 481 and 482.

[0056] 5A shows a block diagram of a ring 500 of detector modules including detector modules 510 and 513 having a single, off-center response line 530 from a single annihilation event 520, according to one embodiment. The scintillation events in detector modules 510 and 513 are used to calculate a response line 540 (a "calculated response line") that corresponds to the true response line 530 but is offset by parallax error, as described herein. Nevertheless, the detection of the scintillation events in detector modules 510 and 513 represents a "true" coincidence event with a set of detected gamma photons resulting from a single annihilation 520.

[0057] 5B shows a ring 500 of detector modules having two off-center response lines 531 and 532 based on multiple annihilation events 521 and 522, according to one embodiment. In the illustrated example, annihilation event 522 generates opposing gamma photons such that one gamma photon is absorbed and detected by detector module 511, while the opposite corresponding gamma photon escapes the detector ring due to its trajectory being outside the ring. Similarly, annihilation event 521 generates opposing gamma photons, but only one of the gamma photons from annihilation event 521 is detected (e.g., by detector module 514). If pairs of scintillation photons dissociated from the gamma photons generated by annihilation events 521 and 522 are detected by detector modules 511 and 514 simultaneously or within a few nanoseconds of each other, the PET scanner may erroneously calculate an artificial response line 541, referred to as a "random" event. That is, the PET scanner may erroneously assume that a gamma photon detected by detector module 511 and a gamma photon detected by detector module 514 originate from the same annihilation event. As a result, the response line calculated by the PET scanner is "artificial" in that the calculated response line does not correspond to a true or bona fide response line, where the coincidence prompt is actually a random event rather than a true coincidence event.

[0058] 6 illustrates a ring 600 of detector modules with temporally and spatially separated lines of response 630, according to one embodiment. The lines of response 630 correspond to positron annihilations 620 detected by opposing detector modules 610 and 612. According to various embodiments described herein, improved photosensor technology, configurations, arrangements, and / or high-speed electronics described herein can facilitate improved measurement of the difference in arrival time between two detected gamma photons. Such methodology, referred to in the art as time-of-flight (ToF), improves the ability to calculate the origin of positron annihilations 620 along the lines of response 630.

[0059] For example, a timing resolution of 500 picoseconds allows the origin of a positron annihilation 620 along a line of response 630 to be localized to within approximately ±7.5 centimeters, e.g., dividing the scanner's 70-centimeter field of view into 4.68 sections (or ±9 sections or more) for each line of response. Improved localization of annihilation events reduces image reconstruction uncertainty (noise) and improves the scanner's signal-to-noise ratio. This signal-to-noise improvement is essentially realized as improved sensitivity compared to non-ToF scanners, with the realized sensitivity improvement being proportional to the square of the signal-to-noise improvement. According to various embodiments, a PET scanner can have a timing resolution of approximately 210 picoseconds.

[0060] Various embodiments of the presently described systems and methods facilitate more accurate and precise PET data acquisition. Example systems and methods described herein reduce and therefore improve temporal resolution, improve the localization of the line of response for a given positron annihilation event to reduce parallax error, improve sensitivity for better image clarity, reduce scan time, and / or reduce radiation imaging, and / or combinations thereof.

[0061] According to various embodiments, a PET scanner system can reduce such errors by segmenting the detector crystal lengthwise (e.g., into two or more crystal pieces) and / or incorporating two different crystal time constants. For example, by using two different time constants, the PET scanner system can use the shape of the photodetector signal to identify which crystal absorbed the energy of an incident gamma photon. In this way, the PET scanner system can determine the depth at which the incident gamma photon interacted (referred to as the depth of interaction (DoI)) within each elongated scintillation crystal length segment. Such depth of interaction improves positional resolution and reduces noise in the reconstructed image.

[0062] In some embodiments, a PET system can include multiple detector modules, each including a two-dimensional array of elongated scintillation crystals segmented longitudinally into two or more sub-crystals with two different crystal time constants. An array of photodetectors associated with the elongated scintillation crystals detects scintillation photons originating from scintillation events within one of the segmented volumes. The shape of the signal generated by the array of photosensors differs based on the crystal time constant of the segmented volume from which the scintillation photon originates. Thus, the system can determine depth-of-interaction information (i.e., from which segmented volume the scintillation light originated) based on the shape of the signal generated by the array of photodetectors. As described above, depth-of-interaction information can be used to improve positional resolution and reduce noise in reconstructed images. In embodiments in which each longitudinally segmented volume within a particular scintillation crystal has a different (i.e., distinct, unique, distinguishable) crystal time constant, reflective material may not be disposed between adjacent segmented volumes.

[0063] According to various embodiments, a PET scanner includes one or more rings of detector modules. Each ring of detector modules includes multiple 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) arranged on the detector ring. A distal end face of each elongated scintillation crystal faces radially of the detector ring to receive gamma photons from annihilation events. Each elongated scintillation crystal may also include two axially oriented side faces, two cross-axially oriented side faces, and a proximal end face. Each elongated scintillation crystal may include an array of photosensors (e.g., a single photodetector or an array of photodetectors) arranged along one of the axially oriented side faces. The array of photosensors detects scintillation photons from scintillation events within a corresponding scintillation event.

[0064] A reflective material is disposed (e.g., as a layer or coating) on ​​the proximal end face, the distal end face, two opposing lateral axially facing sides, and the other axially facing side (i.e., the axially facing side without the array of photosensors) of each elongated scintillation crystal. The reflective material operates to internally reflect scintillation photons such that all or substantially all of the scintillation photons from a scintillation event within a given elongated scintillation crystal are detected by the array of photosensors associated with the given elongated scintillation crystal. Thus, the reflective material operates to prevent light sharing between adjacent scintillation crystals. In some embodiments, the reflective material (e.g., coating, layer, film, etc.) may be bidirectionally reflective such that a single reflective material layer or coating may be shared by adjacent scintillation crystals.

[0065] In various embodiments, each elongated scintillation crystal may be in the shape of an elongated N-sided polygonal prism, such as a hexagonal prism, an elongated rectangular prism, a square prism, an octagonal prism, or a triangular prism. The length of the elongated scintillation crystal may be, for example, between 10 millimeters and 30 millimeters, and the axial and cross-axial widths may be, for example, between 2 millimeters and 10 millimeters. In various examples, the thickness of the array of photosensors may be less than 500 micrometers, and the thickness of the reflective material (e.g., coating) between adjacent elongated scintillation crystals may be less than 100 micrometers.

[0066] As described herein, each elongated scintillation crystal in a detector module can be configured for lateral readout via an array of photosensors disposed on one of its axially oriented sides. In some embodiments, the array of photosensors includes a two-dimensional array of individual photodiodes, such as a two-dimensional array of single-photon avalanche diodes in a silicon photomultiplier. The two-dimensional array of individual photodiodes on the axially oriented side can be used to provide separate photosensor measurements corresponding to separate depth ranges between the proximal and distal end faces of each elongated scintillation crystal. In some embodiments, one of the axially oriented sides of each elongated scintillation crystal in a detector module can include multiple silicon photomultipliers disposed along the length of the axially oriented side.

[0067] In various embodiments, the PET scanning system further includes a cooling system for cooling the detector module. The PET scanning system further includes communication with an imaging system connected to the detector module for generating an image based on the electronic output from the detector module. The diameter of the detector ring (bore) of the PET scanning system may be greater than about 25 centimeters. In many embodiments, the axial bore depth of the multiple axially aligned rings is less than the diameter of the detector ring.

[0068] Many existing computing devices and infrastructures can be used in conjunction with the presently described systems and methods. Some of the infrastructure that can be used with the embodiments disclosed herein, such as processors, microprocessors, microcontrollers, computer programming tools and techniques, digital storage media, image processing devices, image processing techniques, and communication links, is already available. Many of the systems, subsystems, modules, components, etc. described herein improve temporal resolution, improve positional delineation for a given positron line of response to reduce parallax error, improve sensitivity for better image clarity, reduce scan time, and / or reduce patient radiation exposure. The systems, subsystems, modules, and components may be implemented as hardware, firmware, and / or software, as will be understood by those skilled in the art and in the context of the relevant description. The various systems, subsystems, modules, and components are described in terms of the functions they perform, as a wide variety of such possible implementations exist. For example, it is understood that many existing programming languages, hardware devices, frequency bands, circuits, software platforms, networking infrastructures, and / or data stores can be utilized, alone or in combination, to implement particular functions.

[0069] It is also understood that two or more of the systems, subsystems, components, modules, etc. described herein may be combined into a single system, subsystem, module, or component. Furthermore, many of the systems, subsystems, components, and modules may be duplicated or further divided into discrete systems, subsystems, components, or modules to perform subtasks of those described herein. Any of the embodiments described herein may be combined with any combination of the other embodiments described herein.

[0070] Some components of the disclosed embodiments are described and illustrated in the figures herein, many of which can be arranged and designed in a wide variety of different configurations. Furthermore, features, structures, and operations related to one embodiment may be applied to or combined with features, structures, or operations described in connection with another embodiment. In many cases, well-known structures, materials, or operations have not been shown or described in detail to avoid obscuring aspects of the disclosure. Many of the illustrations are provided in block diagram form to illustrate general configurations and may not be drawn to scale. The right to add described embodiments or features to any figure and / or as new figures is expressly reserved.

[0071] In some embodiments of the presently described systems and methods, each array of photosensors includes one or more (e.g., a two-dimensional array of hundreds or thousands) of 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. One of the digital streams encodes the high-frequency content for time measurement. The second digital stream encodes a low-bandwidth version of the signal for numerical integration to measure nuclear energy (e.g., calculate total energy). This circuitry can be used for nuclear pulse processing in PET or other applications utilizing digital signal processing of a large number of optical signals where high time resolution is required for time-of-flight measurements, such as light detection and ranging (LIDAR).

[0072] In some examples, the array of photosensors can provide both a capacitively decoupled high-frequency signal output (referred to as a "high-speed signal" output) and a "low-speed signal" output that is more comprehensive but includes a lower-frequency spectrum. In other examples, the array of photosensors may employ only two-terminal photosensors, where high-speed timing and / or energy measurements can be obtained at the same node or various nodes within a bias network of resistive, capacitive, inductive, and / or active components. In various embodiments, the low-noise analog circuitry detects low-voltage threshold crossings of the high-speed signal via a comparator, enabling signal propagation to a high-speed analog-to-digital converter (ADC) for digital encoding. In some embodiments, the low-noise analog circuitry detects low-current or low-voltage threshold crossings of the high-speed signal via a comparator, enabling signal propagation to a high-speed analog-to-digital converter (ADC) for digital encoding.

[0073] Additionally, comparators can be utilized to select multiplexer inputs to enable separate digital encoding of slower speed signals (or lower bandwidths). Many of the described embodiments provide improvements in component propagation delays and their interrelated effects on bandwidth to successfully enable shared digital encoding of multiple front-end photodiodes that are part of an array of photosensors of one or more scintillation crystals of a detector module of a PET scanner system.

[0074] Transmission paths are utilized that are optimized to achieve target propagation delays. Many embodiments of the presently described systems and methods digitally eliminate traditional analog time walks and enable superior determination of the timing origin of nuclear scintillation events through spectral distortion correction and machine learning algorithms. Furthermore, the systems and methods described herein eliminate or avoid the use of application-specific analog electronics used for timing pickoff, providing low-noise sharing of high-speed electronics that saves power consumption, heat dissipation, packaging volume, and cost compared to traditional approaches. Additional embodiments may provide efficient thermal control or cooling of the scintillator material and photosensor, providing a signal-to-noise boost. Such improvements in nuclear scintillation event timing offer significant clinical implications, for example, by reducing image noise, thereby increasing sensitivity and improving diagnostic clarity.

[0075] 7 is a simplified schematic diagram of a high-speed diode circuit network 700 according to one embodiment. The photosensing diode circuit network 700 operates to provide high-bandwidth digital encoding of nuclear scintillation events. One or more photodiodes 708 (shown as PD1-PDn) may be reverse-biased to operate in Geiger mode. The highest bandwidth output of each photodiode 708 is capacitively isolated and connected to a comparator 709 and a high-speed transmission line 710. The lower bandwidth connection of the photodiode 708 is connected via a multiplexer 706 to a low-speed transmission line 714 for selective encoding by an analog-to-digital converter 707.

[0076] A comparator 709 in the switch circuit block 703 compares the signal from the high bandwidth channel with a low voltage threshold 715, Vth (which could alternatively be a low current threshold). The comparator 709 activates a flip-flop 711 when the signal from the high bandwidth channel exceeds Vth. The output of the flip-flop 711 biases the output diode 713 'on', thereby allowing the high bandwidth photodiode signal to propagate from the high speed transmission line 710 through a capacitance (labeled Cdcb) and to ground through an inductor (labeled Lr) for digital encoding at a high sampling rate by the A / D converter 705.

[0077] 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 in blocks 702 and 704. Circuit blocks 701 and 703 may be replicated any number of times to allow for sharing of AD converter 705. Output diode 793 represents the nth output of n replicated circuit blocks 701 and 703 for any number of sets of scintillation crystal photodiodes in a detector module of a PET scanner system.

[0078] 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), along with their respective comparators 709, 719 and flip-flops 711, 721, identify to the control logic 716 the specific set of photodiodes 708, 718 to be triggered. Control outputs 732 (EN1-ENn) are normally "on," signaling which flip-flops 711, 721 to disable while a particular single diode is triggered and forward conducting. Upon completion of the photodiode nuclear pulse signal envelope (which may be predetermined by its characteristic decay constant), control logic 716 resets the previously fired 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 state, thereby turning off their corresponding output diodes D1-Dn (713, 723-793). This turns off the corresponding output diodes D1-Dn (713, 723-793). This returns all control outputs 732 to "on" and ready to be triggered by a subsequent eligible nuclear pulse signal.

[0079] Additionally, output bus 717 of control logic 716 is used to select the low bandwidth output of a particular photodiode for propagation through multiplexer 706 based on trigger switch circuit inputs 731 (Q1-Qn). As shown, the low speed signal is propagated through low speed transmission path 714 and received as the "energy signal" input of multiplexer 706. The energy signal selected by output bus 717 is sent from multiplexer 706 to analog-to-digital converter 707 for digital encoding. Analog-to-digital converter 707 may be a relatively low sampling rate analog-to-digital converter.

[0080] Diode network 700 digitally encodes the output of a set of multiple photodiodes (e.g., a photodiode network). This is generally illustrated by circuit blocks 701 and 702. AD converters 705 and 707 can be used to digitally encode any number of sets or photodiodes, where each array of photodiodes includes at least one photodiode. As used herein, a photosensor can include any number of photodiodes.

[0081] In some embodiments, the photodiode network is comprised of multiple photodiodes sharing a common high-bandwidth output and a relatively low-bandwidth output, which can propagate over high-speed transmission path 710 and low-speed transmission path 714, respectively. Corresponding trigger switch circuit blocks 703 and 704 mutually exclusively propagate high-speed nuclear pulses for digital encoding by A / D converter 705 and provide highly encoded signals that can be used to determine the timing of the nuclear pulse's origin, while multiplexer 706 and A / D converter 707 provide low-bandwidth encoding for use in energy discrimination of the same nuclear pulse. Generally, higher-bandwidth signals are referred to herein as "high-speed" or "timing" signals, and lower-bandwidth signals are referred to as "slow" or "energy" signals.

[0082] The high-speed transmission path 710, labeled Tx3 (or Tx4 in the corresponding block 704), is configured to allow a delay sufficient to allow the comparator 709 and trigger switch components (e.g., flip-flop 711, inductor 712, and output diode 713) to conduct charge before the upcoming nuclear pulse signal, qualified by threshold 715 and comparator 709, arrives at the input of output diode 713. Similarly, the low-speed transmission path 714, labeled Tx1 (or Tx2 in the corresponding block 702), is configured to allow a delay sufficient to allow the trigger switch circuit block 703 (or 704) to activate control logic 716 and the multiplexer 706 to conduct charge to fully encode the nuclear pulse signal via the A / D converter 707. The high-speed transmission path 710 and the low-speed transmission path 714 are configured to generate a signal delay sufficient to allow encoding of the entire nuclear scintillation pulse envelope. Additionally, the high speed channel 710 and the low speed channel 714 are selected to have sufficient spectral bandwidth for their respective signal content.

[0083] 8A is a two-port network lumped element model 800 of a conductive transmission line, according to one embodiment. The lumped elements 805 of the transmission line are modeled by discrete elements, as described below. Resistive element 801 represents the series resistance of the transmission line. Inductive element 802 represents the series inductance or reactance of the transmission line. Resistive element 803 represents the shunt conductance of the transmission line, and capacitive element 804 represents the shunt capacitance (admittance) of the transmission line.

[0084] 8B shows a distributed model 810 of lumped elements 805A-805N per unit length of a transmission line, according to one embodiment. Each lumped element 805A-805N is embodied by a two-port network lumped element 805 of FIG. 8A. A relatively short transmission line can be embodied as a single two-port lumped element 805. However, a longer transmission line is more accurately modeled as a distributed model of lumped elements 805A-805N per unit length, where the discrete element values ​​of each lumped element 805A-805N may be the same or different from one another, as described herein and understood in the art.

[0085] Based on the solution of the telegrapher's equation derived by Oliver Heaviside from James Maxwell's equations, the transfer function H(ω,x), commonly called the propagation function, of a transmission line is given by the following equation:

[0086]

number

[0087] where: x = length of transmission line, L = series inductance, R = series resistance, C = shunt capacitance, G = shunt conductance, where R, L, G, and C are values ​​per unit length.

[0088] The propagation coefficient γ(ω) is calculated as the negative natural logarithm of the propagation function.

[0089]

number

[0090] where the real component of γ represents the attenuation per unit length in nepers and the imaginary component represents the phase lag per unit length in radians. For frequencies where the impedance or admittance of the reactive component is significantly greater than the respective passive resistance or conductance, the propagation function is approximately:

[0091]

number

[0092] This is a linear phase delay, and the propagation velocity per unit is given by its reciprocal.

[0093]

number

[0094] This is equal to the rate per unit of material.

[0095]

number

[0096] where: ε r = relative permittivity, μ r = relative permeability, Propagation delay is given by the inverse of the signal speed of the line.

[0097] The impedance of a transmission line is given by the square root of the ratio of the series line resistance and impedance to the shunt conductance and admittance.

[0098]

number

[0099] And at frequencies where the per-unit reactive impedance and admittance significantly exceed the per-unit passive resistance and conductance, the characteristic impedance is simply the square root of the ratio of inductance to capacitance for all frequencies thereafter, and thus appears theoretically to be purely resistive.

[0100] For conductive transmission lines, because the overall propagation delay of the line is much shorter than the rise or fall time of the input signal, this mode of operation is commonly referred to as the lumped element region and can be modeled by the two-port network lumped element model 800. The lumped element region is considered when:

[0101]

number

[0102] Here, x is the length of the transmission path, and γ (gamma) is the propagation coefficient.

[0103] Transmission lines operating at lengths beyond lumped element boundaries enter the RC region at relatively low frequencies (e.g., below 1 MHz). In this region of operation, inductance reactance becomes negligible, but series resistance remains significant, especially when considered in combination with the overall capacitance of the line. In this mode, the bandwidth of the line is essentially dominated by the single-pole response, limiting the signal bandwidth to:

[0104]

number

[0105] where f 3db represents the corner of the low-pass frequency characteristic.

[0106] Beyond this frequency range and when longer than the concentrated element region, the transmission line operates in the LC region (e.g., 1 MHz < f < 10 MHz), where the series inductive reactance dominates the line resistance and the capacitive admittance dominates the line shunt conductance. On a more practical basis, non-linear effects occur at frequencies beyond the LC region, which includes the alternating current (AC) skin effect. This attenuates the signal on the line and limits the bandwidth. Further, in commonly used materials and structures, dielectric losses increase at frequencies generally above approximately 1 GHz.

[0107] 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 theoretically, a properly terminated transmission line appears as merely a resistance to a signal transmitter at frequencies above a moderate frequency. In reality, losses occur at higher frequencies due to the skin effect and dielectric losses. These losses effectively limit the bandwidth and result in group delay. Thus, it is understood that the delay and bandwidth of a conductive transmission line are inversely related. This inverse correlation poses a challenge to the design and implementation of the high-speed transmission line 710. Here, a delay exceeding 2 nanoseconds is desirable to ensure the subsequent propagation delay resulting from the implementation of the elements 709, 711, 712, 713.

[0108] The output diode 713 is activated prior to a qualified propagation signal (e.g., when a photo sensor signal exceeds the threshold V th 715) reaching the input of the output diode 713 from the high-speed transmission line 710 (or another high-speed delay element). At the same time, the bandwidth of the high-speed transmission line 710 (or another high-speed delay element) is high enough to hold the high-bandwidth signal with a fast rise from the photo sensor 708 required for timing resolution. Similarly, the low-speed transmission line 714 has sufficient bandwidth while the trigger switch circuit block 703, control logic 716, and multiplexer 706 also provide sufficient propagation delay to operate and enable conduction in time for the corresponding energy signal to be fully digitally encoded by the AD converter 707.

[0109] FIG. 9 shows a time-domain graph 900 of two nuclear pulse signals, according to one embodiment. Each of the illustrated silicon photomultiplier photosensor signals 901 and 902 results from the deposition of nuclear particles on a coupled scintillator. Signal 901 is representative of a capacitively isolated (i.e., high-pass filtered) silicon photomultiplier avalanche diode signal, referred to herein as the fast or timing signal of the detector module. Signal 902 is representative of a low-pass filtered version of the same avalanche diode signal, referred to herein as the slow or energy signal of the detector module. The decay constant of slow signal 902 is proportional to the decay constant (τ low-pass >5τ scint ), the maximum peak of signal 902 approximates the total charge Q collected from the photosensor divided by the capacitance C of the circuit. In another embodiment, a circuit or processing unit can be used to integrate the unfiltered photosensor signal to obtain a proportional energy measurement.

[0110] The detector module can use the fast timing signal 901 to determine the timing of nuclear scintillation events (also referred to herein as original or initial). The fast and slow signals 901 and 902 are more closely reproduced by signals 903 and 904, respectively, which represent white, or broadband, thermal noise, which has an approximately Gaussian amplitude distribution. When using voltage discrimination, indicated by a voltage level threshold 907, this noise, like other noise sources, also introduces timing uncertainty. The use of the voltage level threshold 907 to discriminate the timing of a signal is referred to as "timing kickoff."

[0111] Using voltage discrimination on signal 904 (corresponding to slow signal 902) results in timing uncertainty. Gaussian noise in slow signal 902 results in a Gaussian timing jitter histogram with timing uncertainty graphed as 906. This effect is noticeable because timing pickoff of a slow-rising signal (i.e., a shallow slope) is more susceptible to the vertical amplitude effect of noise, referred to herein as slope-to-noise consideration. However, fast signal 903 (corresponding to fast signal 901) is less susceptible when considering its fast-rising signal (i.e., a steep slope). Comparing fast signal 903 to a voltage level threshold 907 (e.g., threshold) results in a narrower timing distribution graphed as 905. Therefore, the impact on time resolution due to noise in fast signals 901 and 903 is reduced or improved compared to the impact on time resolution due to noise in slow signals 902 and 904.

[0112] Accordingly, various embodiments of the presently described systems and methods utilize a timing circuit to analyze the high-speed, capacitively isolated output signals of one or more photodiodes to determine timing information (e.g., timing-kickoff values) of scintillation events. Similarly, various embodiments of the presently described systems and methods utilize a separate energy circuit to determine energy information (e.g., total energy values) of scintillation events. It is understood that some circuit components may be shared between the timing circuit and the energy circuit. 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 an energy circuit and a timing circuit and / or may share an analog-to-digital converter. Each energy circuit and timing circuit may include multiple channels of discrete electronic components, each channel dedicated to one or more scintillation crystals, and the channels share the same analog-to-digital converter.

[0113] FIG. 10 shows a graph 1000 of the energy spectral density of the two nuclear pulse signals of FIG. 9 in the frequency domain, according to one embodiment. The energy spectral density of signal 901 of FIG. 9 is illustrated by graph line 1010. The energy spectral density of signal 902 of FIG. 9 is illustrated by graph line 1020. As can be seen, the frequency spectrum of low-speed signal 902 is generally cut off around 400 MHz, with trend line 1040 being approximately the level of the noise floor above that. However, for signal 901, the corresponding high-frequency spectrum of graph line 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 a signal impinging on the circuit can be derived by the following relationship, which can be shown to be proportional to the time constant of the circuit:

[0114]

number

[0115] where: t r = 10-90% rise time of the step signal, and f 3db = 3db bandwidth (BW) of the circuit.

[0116] In various embodiments, high frequency content is preserved to generate a timing signal that is steeper than would otherwise be achievable. Given the timing signal's slope-to-noise ratio and slope limitations due to the transmission path's bandwidth, some embodiments of the system optimally preserve wideband timing channels to reduce timing jitter (e.g., uncertainty) and improve timing pickoff.

[0117] 7 , high-speed transmission line 710 may be selected to have the smallest or shortest practical length to reduce high-frequency attenuation while still generating sufficient delay necessary for the remaining circuit elements in the box for block 703 (e.g., comparator 709, flip-flop 711, inductor 712, and diode 713) to activate in time for A / D converter 705 to encode the capacitively isolated “high-speed signal” from photodiode 708. As a result, the propagation delays of comparator 709, flip-flop 711, bias tee inductor 712, and output diode 713 are selected to operate as quickly as possible so that the propagation delay of high-speed transmission line 710 is short (e.g., so that high-speed transmission line 710 is physically short). Bias tee inductor 712 may be designed for high-speed switching utilizing a low-loss, low-relative-permeability core, such as an air core, and may exhibit a low inductance, e.g., less than 100 nanohenries, to facilitate fast biasing of diode 713. In contrast, the slow rising energy signal from photodiode 708 may employ a much longer slow transmission path 714, resulting in more propagation delay, given that the high frequency content is significantly lower than that of the high speed timing signal. For example, the delay of the slow transmission path 714 may be more than twice the delay of the high speed transmission path 710. In some embodiments, the delay of the slow transmission path 714 may be five times or even longer than the delay of the high speed transmission path 710.

[0118] Given the importance of the high-frequency spectrum of timing signals, specialized high-frequency, fast-switching diodes, such as PIN diodes, that utilize intrinsic semiconductor regions with high levels of implantation can be utilized to reduce the delay required for the high-speed transmission line 710. For high-frequency, small signals, forward-biased PIN diodes advantageously operate as current-controlled variable resistors, providing resistances of less than 100 ohms at 1 GHz, and in some cases, small currents, such as approximately 10 milliamps or less than 10 ohms. A relatively large cross-sectional intrinsic region reduces the forward conduction (i.e., "on") resistance of the PIN diode, thereby improving (i.e., reducing) insertion loss, but undesirably increases junction capacitance in the reverse-biased (i.e., "off") state. Increasing the thickness of the intrinsic region between adjacent P and N regions reduces capacitance while also increasing reverse-bias breakdown. However, thicker intrinsic regions undesirably increase switching times, thereby requiring correspondingly longer delays from the high-speed transmission line 710. In various embodiments, low junction capacitance improves the operation of circuit blocks 703 and 704 in an untriggered state (e.g., for optimal operation) by isolating photosensor noise from circuit blocks 701 and 702 from feeding through to a common node present at the input of AD converter 705.

[0119] According to various embodiments, these tradeoffs are considered and carefully accounted for in the design and selection of the high-speed transmission line 710 (or other delay element) and output diode 713 to achieve acceptable or optimized timing resolution. The use of high-energy bandgap semiconductors with energies greater than 1.3 eV (at 300°K), such as those provided by gallium arsenide, increases the reverse bias breakdown voltage over silicon, thereby allowing for thinner intrinsic regions to provide faster switching times. Furthermore, gallium arsenide provides carrier lifetimes typically nominally less than 10 nanoseconds, and in some cases nominally faster than 5 nanoseconds, at low currents (e.g., 1-20 milliamps), allowing for rapid removal of charge from the intrinsic region.

[0120] In some cases, the turn-on (i.e., forward conduction mode) time of a PIN diode can be faster than the turn-off (i.e., reverse bias configuration) time due to differences in charge carrier injection versus removal. Such PIN diode characteristics are acceptable in some embodiments given the unique nature of the PET signal. The tolerance for this slow turn-off time is provided by the relatively rare nature of PET single events in a typical block base, whereby each event typically requires a long time to decay (e.g., approximately 200 nanoseconds for lutetium oxyorthosilicate-based scintillators), thus allowing several orders of magnitude of difference in turn-off duration relative to turn-on to be tolerated in this unique case (i.e., fast turn-on and slow turn-off for PET single events). Desirable characteristics for a PIN diode used to implement output diode 713 include low insertion loss (e.g., less than 1 dB) and high turn-off isolation (e.g., greater than 15 dB) for frequencies typically encountered from gamma radiation scintillation photodiode signals, as depicted by FIG. 10, and up to several gigahertz (e.g., 4 GHz) or higher. High turn-off isolation is achieved at these frequencies with a relatively low junction capacitance, e.g., 150 femtofarads or less. The diode cathode can optionally be negatively offset to reduce the threshold voltage at which the diode operates in the on-state (e.g., for forward active / conducting). Gallium arsenide (GaAs) or aluminum gallium arsenide (AlGaAs) diodes, with their fast conductivity, low resistance, and ultra-low capacitance, may be used to achieve target fast switching times for the target bandwidth of transmission line 710. The target rise time of the high-speed signal from photodiode 708 can be, for example, less than 1 nanosecond under expected PET scanner noise levels to achieve precise timing. The target bandwidth of the transmission line 710 can be, for example, 350 MHz or more over the entire length.

[0121] The AD converters 705 and 707 may be sampled at or above the Nyquist rate (i.e., half the period of the highest bandwidth of the respective input signals) to encode the timing and energy signals, respectively. The sampling rate may be selected according to the Nyquist-Shannon sampling theorem. Alternatively, the conversion rate of the AD converters may be selected in terms of uncertainty propagation, utilizing one or more data points of the signal decay curve to achieve an optimal timing pickoff that minimizes the uncertainty propagation given by the following equation:

[0122]

number

[0123] In this case, each data point has an uncertainty of σ that is also correlated to other data points by a correlation coefficient ρ. The system can utilize a formula to reduce or minimize the timing kickoff uncertainty. The correlation coefficient is further 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 concepts can be achieved using other circuits, components, and techniques to achieve the same or similar results. For example, the low-speed energy AD converter 707 could be eliminated and the signal from the high-speed AD converter 705 could be digitally integrated to determine the signal's total energy. Furthermore, the trigger circuit 703 could be implemented using alternative components, such as a two-stage comparator with an RC delay circuit between them, since the scintillator's time constant, and therefore the signal envelope, is uniquely known. Furthermore, the switch element 712 (shown as an inductive element) and the output diode 713 could alternatively be implemented with transistors (e.g., with a common base or common gate) and operated in an open-drain configuration. Furthermore, the high-speed transmission path 710 and the low-speed transmission path 714 could each be implemented optically through the use of optical fibers or other light guides. The photodiode 708 can also be configured differently, with a pull-up resistor connected to the cathode to capacitively decouple high-speed signals from it, while optionally retaining an anode resistor for low-frequency energy signal detection. As a further alternative, the anode of the photodiode can be connected directly to the input of a low-input impedance transimpedance amplifier.

[0125] FIG. 11 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 an array of photosensors associated with each elongated scintillation crystal. The array of photosensors may include any number of photodiodes arranged, for example, in a photodiode array (e.g., one or more single-photon avalanche photodiodes or silicon photomultiplier tubes). The illustrated diagram includes two electronic processing channels 1110 and 1120. The first electronic processing channel 1110 is associated with a first array of photodiodes, including the illustrated first and second arrays 1101 and 1102 of photodiodes. The second electronic processing channel 1120 is associated with a second array of photodiodes, including the illustrated third and fourth photodiode arrays 1111 and 1112.

[0126] According to various embodiments, the processing circuitry associated with a detector module includes two processing channels to enable simultaneous processing of signals from adjacent scintillation crystals. However, it is understood that the processing circuitry of a detector module may include any number of channels sharing any number of analog-to-digital converters. Each channel of the N-channel processing circuitry may be associated with any number of adjacent or non-adjacent scintillation crystals within a block of scintillation crystals of a 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 (e.g., every other) of the elongated scintillation crystals in the detector module may be associated with a first electronic processing channel 1110. More specifically, the photodiodes associated with half of the elongated scintillation crystals in the detector module are connected to the first electronic processing channel 1110 and share a single high-speed timing A / D converter 1105 and a single slower speed or energy A / D converter 1107.

[0128] Thus, the first electronic processing channel 1110 may be associated with any number of photodiode arrays, including the illustrated first photodiode array 1101 and second photodiode array 1102. The photodiode arrays 1101 and 1102 may have their own triggered switch circuits 1103 and 1104 and / or may share a single triggered switch circuit. The triggered switch circuits 1103 and 1104 (or a shared single triggered switch circuit) may be embodied, for example, as described in conjunction with FIG. 7 and / or in accordance with various alternative embodiments described herein. The multiplexer 1106 is used to select between the various photodiode arrays, including the photodiode arrays 1101 and 1102, for energy signal encoding by the energy A / D converter 1107.

[0129] The photodiode arrays in the other half of the elongated scintillation crystal in the detector module can be associated with a second electronic processing channel 1120. These photodiode arrays share a single high-speed timing A / D converter 1115 and a single low-speed or energy A / D converter 1117. The second electronic processing channel 1120 can be associated with any number of photodiode arrays, including the illustrated third photodiode array 1111 and fourth photodiode array 1112. The photodiode arrays 1111 and 1112 may have their own triggered switch circuits 1113 and 1114 and / or may share a single triggered switch circuit. The triggered switch circuits 1113 and 1114 (or a shared single triggered switch circuit) may be embodied, for example, as described in conjunction with FIG. 7 and / or in accordance with various alternative embodiments described herein. Multiplexer 1116 is used to select between various photodiode arrays, including photodiode arrays 1111 and 1112, for energy signal encoding by energy A / D converter 1117. Digital control logic, as described in conjunction with FIG. 7, or variations thereof, may be used to control the diode bias in multiplexers 1106 and 1116 and / or trigger switch circuits 1103, 1104, 1113, and 1114.

[0130] As described in more detail herein with respect to other figures, some (or all) of the electronic processing channels 1110 and 1120 for a particular scintillation crystal in a detector module may be integrated into an intermediate substrate. The intermediate substrate may be disposed behind the photosensor on a side of the elongated scintillation crystal. In some embodiments, each elongated scintillation crystal is associated with a separate intermediate substrate. In other embodiments, multiple elongated scintillation crystals may share a single intermediate substrate. The intermediate substrate associated with multiple elongated scintillation crystals may include, for example, biasing circuits and signal shaping circuits of circuit block 701 according to any of the various embodiments described herein.

[0131] In other embodiments, additional portions of the high-speed diode circuit network 700 may be incorporated into an intermediate substrate associated with one or more of the elongated scintillation crystals. The intermediate substrate may include and / or be connected to a timing circuit and an energy circuit according to any of the various embodiments described herein. The energy circuit may include, for example, a first AD converter, a multiplexer that selects between the photodiode output signals for conversion by the first AD converter, a signal delay path connecting the photodiodes to the multiplexer, and digital control logic that controls the multiplexer selection of the photodiode output signals. The timing circuit may include, for example, a second AD converter that performs high-speed digital encoding of each photodiode output signal, a trigger component such as a flip-flop that selectively triggers encoding by the second AD converter, and a comparator that compares each photodiode output signal to a threshold value.

[0132] FIG. 12 is a diagram illustrating blocks of scintillation crystals 1200 arranged in a two-dimensional array, according to one embodiment. The illustrated diagram shows the distal end faces of the elongated scintillation crystals. As shown in the enlarged view of the elongated scintillation crystals in the upper left, each elongated scintillation crystal includes a reflective material as a layer or coating applied to the distal end face 1230, the transversely oriented side faces 1231 and 1233, and the first axially oriented face 1232. A reflective material is also disposed on the proximal end 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] An array of photosensors 1201 (e.g., one or more arrays of photodiodes) is disposed along an opposing, second, axially-facing surface 1234, which is not coated with a reflective material. Thus, scintillation photons from a scintillation event within one of the elongated scintillation crystals are internally reflected within the elongated scintillation crystal until received by a laterally-positioned photosensor. The reflective material prevents light sharing between adjacent crystals in the two-dimensional array of elongated scintillation crystals. The simplified illustration includes an array of 16 scintillation crystals. However, it will be understood that the two-dimensional array of elongated scintillation crystals may include M elongated scintillation crystals axially of the detector ring and N elongated scintillation crystals transversely to the detector ring axis, where M and N are each integer values.

[0134] The elongated scintillation crystals in the two-dimensional array are seamlessly arranged such that each elongated scintillation crystal, its reflective layer / coating, and its photosensor(s) are adjacent to one another. Alternating shading is used for adjacent elongated scintillation crystals to represent the electronics processing channel associated with each elongated scintillation crystal in the two-dimensional array. In the illustrated embodiment, elongated scintillation crystals depicted with a dotted fill pattern may utilize a first electronics processing channel (e.g., electronics processing channel 1110 in FIG. 11 ). Elongated scintillation crystals depicted with a cross-hatched fill pattern may utilize a second electronics processing channel (e.g., electronics processing channel 1120 in FIG. 11 ). In this configuration, inter-crystal scattered radiation of gamma photons may partially deposit energy in one crystal and the rest of its energy in an adjacent crystal, but may be processed without being wasted due to dead time in one processing channel per block.

[0135] The reflective material may be thin relative to the size of each elongated scintillation crystal (e.g., as a thin panel, a deposited layer, a painted film, a coating, etc.), while the thickness of the photodiode(s) may be substantially thick. This thickness due to the reflective material and sensors consumes space between the crystals, thereby reducing the scintillator area (and volume) compared to a block detector of the same shape without a side-facing photodetector or reflective coating. The percentage ratio of the scintillator area on the emitting face (the illustrated distal end face) to the defined block area is referred to herein as the fill factor of the block. The lower the fill factor, the less sensitive the detector. Minimizing the thickness of the reflective material and / or the thickness of laterally positioned (e.g., side-mounted, side-integrated, and / or other laterally positioned) photosensors results in a higher fill factor.

[0136] Photosensor thicknesses of 400 μm or less are achieved by electrically coupling the photosensor, which may consist of a bare silicon sensor, a scintillation crystal utilizing microlead frame packaging (also known as flat no-lead) or wafer-scale packaging, optionally employing through-silicon vias, to a flexible polyimide printed circuit board or similar material. Bonding or coupling of the photosensor to the side of the elongated scintillation crystal may be achieved through the use of an optical couplant, such as silicone grease or transparent epoxy, or grease, to aid light transmission efficiency and may also act as an adhesive to aid manufacturing and / or improve durability. Such bonding between the scintillation crystal, couplant material, and / or photosensor can be optically optimized by selecting materials within a narrow refractive index range to maximize the critical angle of reflection, thereby providing efficient optical coupling with low reflection. Lutetium oxyorthosilicate (LSO) has a refractive index of approximately 1.82. Bismuth germanate (BGO) has a refractive index of approximately 2.15, and perovskites have a refractive index of approximately 2.0–2.3.

[0137] The surface areas of the elongated scintillation crystals and / or photosensors may be polished, roughened, or otherwise treated to enhance light transmission. Bonding between the silicon packaging and the intermediate substrate and / or printed circuit board may be achieved through the use of solder balls, conductive amalgam paste, conductive epoxy, ultrasonic welding, or the like. Using these techniques, fill factors of 85% or greater, and even 88% or greater, can be achieved for various sizes of scintillation crystals having elongated prismatic shapes.

[0138] In some cases, a semiconductor wafer support material, such as glass, is utilized during photoreceptor fabrication. The support material, such as glass, may be bonded to the light-receiving surface of the wafer. Such glass, or other transparent or translucent material, may also be desirable for its optical properties used during exposure. The glass or translucent material can be thinned to achieve an optimal (increased) fill factor, for example, if the transparent or translucent optical material is thinner than the finished semiconductor photoreceptor. Alternatively, a scintillator material can be utilized as the wafer support material in place of glass or other transparent or translucent material.

[0139] Some scintillation crystals with sideways photosensors, each with one or more sideways photodiodes, are well suited for PET. For example, scintillators such as BGO and LSO typically exhibit a double-exponential response, modeling the finite time required for the scintillator to first form an emissive state and then for the fluorescence to decay. The overall shape of the response is typically modeled as:

[0140]

number

[0141] Here, the time constant T r corresponds to a population of generally fast emitting states, and T f generally models the slow decay of light.

[0142] The light output (i.e., the number of photons emitted, N) is directly proportional to the energy absorbed by the scintillator. For example, LSO emits approximately 25,000 photons per megaelectronvolt (MeV) absorbed, and has a decay time constant (T f ) is approximately 47 nanoseconds (ns). The statistical uncertainty (sample variance) of the light output is approximately equal to the square root of the number of photons measured. Because the double exponential response results in a specific peak time, increasing light output results in a steeper signal front (i.e., rise time), regardless of amplitude. Correspondingly, the timing uncertainty, or jitter, is generally inversely proportional to the signal amplitude of the scintillator's photosensor(s), as depicted in graph 900.

[0143]

number

[0144] This reduces the timing jitter t jitter decreases inversely proportional to the square root of the measured photon number N.

[0145] In time-of-flight PET, reducing the timing uncertainty reduces image noise, thereby improving the image signal-to-noise ratio (SNR):

[0146]

number

[0147] where: D = effective diameter of the imaged object, c = speed of light, SNR non-ToF = signal-to-noise ratio without time-of-flight applied.

[0148] The SNR of a PET scanner is proportional to the square root of the noise effective counts (NEC). Therefore, an improvement in SNR is equivalent to an improvement in coincidences. Therefore, as mentioned earlier, the sensitivity of a PET scanner increases in proportion to the square of the improvement in SNR. However, a loss of fill factor reduces sensitivity. Because a PET image is constructed by lines of coincident response (i.e., from two detectors), a loss of fill factor worsens in proportion to the square of the fill factor. Thus, for example, a fill factor of 90% reduces PET sensitivity by 81%. Therefore, for the lateral readout, there is a design optimum for the volume displaced by the lateral photosensors.

[0149] In various embodiments, multiple rings of detector modules are used to form a detection cylinder. The multiple rings of detector modules forming the cylinder can be described as having an axial depth d of the cylinder (e.g., bore). The axial depth is, in many embodiments, much less than the diameter D of the rings, particularly for human body tomography. Typically, blocks of detector modules are butted axially sequentially to define a constant number of rings throughout the depth of the cylinder.

[0150] Bonding side-facing or side-facing photodiodes to multiple crystals circumferentially or laterally increases the diameter of the cylinder. The increased diameter opens up the bore, making it more susceptible to radiation from outside the field of view, increasing random and scattered radiation and introducing noise into the acquired image. Furthermore, increasing the bore diameter reduces the radiation acceptance angle within the scanner—i.e., the angle at which the response line emerges along the axial length of the bore—and thus reduces sensitivity. The geometric efficiency of a scanner is related to the coverage of the solid angle provided by the detector for a given point source within the field of view.

[0151]

number

[0152] In contrast, orienting the thickness of side-facing or laterally positioned photodiodes along the axial length of the bore (i.e., along the axially-facing side of each elongated scintillation crystal) increases the bore depth but leaves the cylindrical diameter unaffected. Lengthening the bore with inactive (i.e., non-scintillator) material reduces the fill factor, resulting in a loss of sensitivity. However, this reduction in fill factor occurs equally whether the photosensors are positioned on the axially-facing side of the elongated scintillation crystal or on the axially-facing side of the elongated scintillation crystal. Despite the loss of fill factor sensitivity per unit length, increasing the bore length increases sensitivity resulting from the increased acceptance angle. Therefore, locating photosensors on the axially-facing side of the elongated scintillation crystal reduces noise due to random or scattered radiation and improves sensitivity due to the improved acceptance angle compared to locating photosensors on lateral axially-facing sides (e.g., sequentially along the circumference of the bore).

[0153] For reference, elongated scintillation crystal 1250 is axially adjacent to elongated scintillation crystal 1251 and is axially adjacent to (i.e., axially adjacent to) elongated scintillation crystal 1253. Meanwhile, elongated scintillation crystal 1250 is axially adjacent to (i.e., axially adjacent to) elongated scintillation crystal 1252 and is axially adjacent to (i.e., axially adjacent to) elongated scintillation crystal 1254.

[0154] FIG. 13A is a perspective view of a single scintillation crystal pixel 1300 having an array of photosensors 1301-1306 for side readout, according to one embodiment. The elongated scintillation crystal 1350 has a rectangular prism shape and is covered (i.e., surrounded or coated) on five sides by a reflective material (e.g., a coating or thin layer) and is covered on one side by an array of photosensors 1301-1306. The elongated scintillation crystal 1350 is described as having four sides, two axially oriented sides, and two lateral oriented sides. The distal end face 1312 is described as the emission face and is oriented radially to the detector ring to receive gamma radiation photons from annihilation events. The proximal end face 1311 is at the opposite end of the length of the elongated scintillation crystal 1350. The array of photosensors 1301-1306 is disposed on the axially oriented side of the elongated scintillation crystal 1350. Each array of photosensors 1301, 1302, 1303, 1304, 1305, and 1306 may include a plurality (e.g., a two-dimensional array) of photodiodes. For example, photosensor array 1301 may be comprised of a two-dimensional array of single-photon avalanche photodiodes operating in Geiger mode, and photosensor array 1301 is a silicon photomultiplier.

[0155] As an example, gamma radiation (e.g., gamma photons) may impinge on elongated scintillation crystal 1350 at location 1307 near distal end face 1312. In existing embodiments in which a single photosensor is located at the proximal end, the scintillation light must travel from location 1307 to the proximal end of elongated scintillation crystal 1350, potentially experiencing absorption and / or multiple reflections along the length of elongated scintillation crystal 1350 and traveling at various angles, resulting in a distribution of photon arrival times at the proximal detection end face. Such travel results in attenuation of the scintillation light and further results in a distribution of photon arrival times that slows the signal rise time and overall subsequent light detection (e.g., as discussed in connection with FIG. 9 ), resulting in timing jitter.

[0156] In contrast, detection of scintillation light along the sides of elongated scintillation crystal 1350 by a laterally positioned array of photosensors 1301-1306 results in higher light collection over a shorter light travel path. The faster detection provided by the laterally positioned array of photosensors 1301-1306 reduces the attenuation and range of the photon arrival time distribution, thereby increasing the photosensor output signal and slope.

[0157] The laterally positioned (e.g., laterally mounted, laterally integrated, or otherwise laterally positioned) array of photosensors 1301-1306 further aids in determining the depth of interaction, or scintillation occurrence (e.g., locating location 1307), which provides the scanner system with the ability to reduce parallax errors, as previously described, and also provides the scanner system with the ability to achieve improved time-of-flight resolution by reducing spatiotemporal uncertainty in the points of absorption / scintillation of gamma radiation photons. For example, the points of scintillation resulting from gamma photons at location 1307 being absorbed within the distal portion of elongated scintillation crystal 1350 result in a gradient of illumination (or photon impingement) on the array of photosensors 1306-1301, in order of magnitude (or luminous flux). By measuring the magnitude of the output of each array of photosensors 1301-1306, the area of ​​gamma photon absorption can be determined.

[0158] As a second example, gamma photons 1308 generated at an oblique angle relative to elongated scintillation crystal 1350 result in scintillation as shown, with greater illumination of the array of photosensors 1304 and 1305 than of the array of photosensors 1306, resulting in a gradient of light extending in both directions (i.e., proximally and distally) that allows ratiometric determination of gamma absorption points within elongated scintillation crystal 1350. As yet another example, gamma photons 1309 are absorbed within the proximal end of elongated scintillation crystal 1350, resulting in maximum coverage of light within the array of photosensors 1302 relative to all other arrays of photosensors 1301 and 1303-1306. From this absorption point, a gradient of light extends both proximally and distally of the elongated scintillation crystal 1350, resulting from absorption of light along the lateral length of the elongated scintillation crystal 1350, and again, by comparing the output amplitude and / or timing of each photosensor, a ratiometric determination of the absorption point can be determined. One method for determining the absorption region is via the following method for a set of lateral photosensors, e.g., S={1301, 1302, ... 1306}.

[0159]

number

[0160] where: AF i = Amplitude ratio of photosensor i, AP i = peak amplitude of photosensor i, |S| denotes the cardinality of set S, and I=argmax(AF[J]), where I is the index number of the sensor with the largest amplitude, which represents the geometric area of ​​the scintillation (or partial scintillation in the case of intercrystalline scattering).

[0161] The amplitude peak, AP, may be, for example, the peak of signal 902, which represents the approximate total charge collected by the photosensor, as described above. Other methods may optionally be employed to alternatively obtain AP, such as integrating the high-speed timing signal 901 by digital signal processing or taking the peak of signal 901. Furthermore, only a specific amplitude range may optionally be considered valid, such that scattered or pile-up energy is rejected (e.g., by using low-level discriminators (LLDs) and high-level discriminators (ULDs)). Alternatively, scattered energy deposited on two or more crystals (i.e., less than the LLD value) may be utilized to interpolate the location of the scintillation as occurring between two or more crystals where the total energy between the two or more crystals sums between the LLD and ULD values, and thus, overall, is considered to have resulted from a single gamma photon with inter-crystal scattering. Furthermore, the scattered energy measured between various crystals may be utilized to reconstruct the geometric line of response of a single gamma photon where the total energy sums between the LLD and ULD (e.g., sums to approximately 511 keV). A gamma photon resulting from inter-crystal scattering can be determined as terminating at its last partial scintillation location, its first partial scintillation location, or any location in between.

[0162] In some embodiments, interaction depth is determined by measuring the arrival time of photons at the various arrays of photosensors 1301-1306. As an example, for a gamma photon 1309 absorbed within the proximal end face 1311 of the elongated scintillation crystal 1350, the earliest detection of scintillation light occurs at the nearest array 1302 of photosensors, followed by 1301 and 1303, then 1304, 1305, and finally 1306. Localization of the scintillation event can be determined by measuring the time of first light detection by each photosensor (or photosensor array) in the array (or arrays) of photosensors 1301-1306 based on a known velocity (e.g., the speed of light through a medium) within the corresponding elongated scintillation crystal 1350.

[0163] FIG. 13B shows an exploded view of a single scintillation crystal pixel 1300 having an array of photosensors 1301-1306 and panels of reflective material 1321-1325 separated from an elongated scintillation crystal 1350, according to one embodiment.

[0164] 13C is a perspective view of an elongated scintillation crystal 1350 having an array of photosensors 1301-1306 for lateral photosensor readout, along with an intermediate substrate 1314 and a connector array 1315, according to one embodiment. The intermediate substrate 1314 is connected to or optionally integrated with the array of photosensors 1301-1306 along a first axially oriented side of the elongated scintillation crystal 1350. The intermediate substrate 1314 is operable to transmit and / or modify electrical signals from the array of photosensors 1301-1306 of the elongated scintillation crystal 1350 toward the proximal end face 1311 (e.g., toward the connector array 1315).

[0165] Electrical signals from the photodiode semiconductor dies forming the photosensor array are directly connected to or integrated into a very thin electrical intermediate substrate 1314. The electrical intermediate substrate 1314 can be composed of, for example, silicon, glass, fiberglass, or organic substrates to form homogeneous or heterogeneous integrated components. In various embodiments, the intermediate substrate 1314 can include integral connections to the individual photodiodes in the photosensor array. In various embodiments, the intermediate substrate 1314 can include, for example, the circuit components illustrated in circuit block 701 of FIG. 7. The intermediate substrate 1314 enables a very high fill factor of the crystal array. The intermediate substrate 1314 can utilize any combination of metal, polysilicon, through-silicon vias, conductive pads, wire bonds, or similar techniques for interconnection. The intermediate substrate 1314 can optionally utilize solder balls, microbumps, conductive amalgam paste, conductive epoxy, ultrasonic welding, etc. The intermediate substrate 1314 can connect, actively and / or passively modify, and / or network, and / or transmit the signals of the multiple photodiodes to the proximal end of the elongated scintillation crystal 1350. The intermediate substrate may be embodied as an integrated circuit. Connection to the intermediate substrate 1314 at the proximal end face 1311 may be via a connector array 1315, which may include, for example, solder balls such as a ball grid array (BGA), or rigid metal conductors.

[0166] In some embodiments, the BGA connections on the intermediate substrate 1314 can optionally form connections to a flex circuit board, allowing for routing of intermediate substrate signals outside of the elongated scintillation crystal array volume, etc. In this way, the intermediate substrate 1314 differs from (optional) flex circuit boards such as flex circuit boards made from polyimide or thick printed circuit boards (e.g., typically 1.6 mm thick or greater) made from fiberglass.

[0167] In some embodiments, intermediate substrate 1314 receives power or signals from external circuitry via connector array 1315. Intermediate substrate 1314 may be passive or active, and may have the ability to multiplex, amplify, or otherwise process signals, including, for example, conversion to optical signals. The active components for such processing may be fabricated directly within intermediate substrate 1314, as in the case of a silicon intermediate substrate, or may be bonded onto and directly connected to intermediate substrate 1314.

[0168] In some embodiments, intermediate substrate 1314 provides very thin mechanical support for the array of photosensors. Intermediate substrate 1314 also allows for the routing of power and signals to and from each photodiode. In this manner, power and signals are routed through intermediate substrate 1314 as electrical components form a serial chain including the photodiodes of photosensor array 1301-1306, intermediate substrate 1314, and finally electrical connector array 1315.

[0169] FIG. 14A is a diagram of another embodiment of a sub-divided scintillation crystal pixel 1400 having an array of photosensors 1401-1406 for lateral photosensor readout, according to one embodiment. As shown, the elongated scintillation crystal 1450 has the shape of a rectangular prism. The elongated scintillation crystal 1450 is divided along its length into multiple divided volumes with reflective layers disposed between adjacent faces of the divided volumes. In the illustrated embodiment, the elongated scintillation crystal 1450 is divided into six cubic volumes 1451-1456, each separated by a reflective layer 1420 (see FIG. 14B). In this manner, each divided cubic volume 1451-1456 contains scintillation light within its separate volume and shares light exclusively with its corresponding array of photosensors 1401-1406. Little or no scintillation light is shared between the cubic volumes 1451-1456.

[0170] For example, scintillation light generated by incident gamma photons within cubic volume 1454 is internally reflected within cubic volume 1454 and received (e.g., detected) by array of photosensors 1403. Similarly, scintillation light generated within cubic volume 1451 is detected exclusively or nearly exclusively by array of photosensors 1406. One advantage of the embodiment shown in Figures 14A-C is that the greatest amount of scintillation light is identifiable as originating from a particular region (i.e., cubic volume). Furthermore, complete collection of light occurs within a particular cubic volume, shortening the photon travel length and therefore reducing variability in the timing of photon arrival at the photosensor(s).

[0171] FIG. 14B shows an exploded view of a subdivided scintillation crystal pixel 1400 having an array of photosensors 1401-1406 and a panel of reflective material 1420 separating each subdivided cubic volume 1451-1456, according to one embodiment. The elongated scintillation crystal may be subdivided into more than six subdivided volumes or into fewer than six subdivided volumes. Furthermore, each subdivided volume need not be cubic. For example, the elongated scintillation crystal may be subdivided into only two volumes, each of which would still be in the shape of an elongated rectangular prism. In other embodiments, the elongated scintillation crystal may not have square end faces, and therefore the subdivided volumes are not necessarily cubic volumes.

[0172] FIG. 14C is a diagram of a sub-divided scintillation pixel having an array of photosensors 1401-1406 for lateral photosensor readout, along with an intermediate substrate 1414 and a connector array 1415, according to one embodiment. The intermediate substrate 1414 may operate and / or be configured according to any of the variations and embodiments described in conjunction with the intermediate substrate 1314 of FIG. 13C. Similarly, the connector array 1415 may operate and / or be configured according to any of the variations and embodiments described in conjunction with the connector array 1315 of FIG. 13C. As with all figures, the sizes, proportions, shapes, and dimensions of the various components and elements are drawn to scale and are not necessarily drawn in proportion to one another. For example, the connector array 1415 may be much smaller and / or have a different shape than shown. Additionally, the intermediate substrate 1414 and / or the connector array 1415 may be closer to or extend further from the proximal end face 1411 of the elongated scintillation crystal 1450.

[0173] An additional benefit of side photosensing is achieved when otherwise undesirable, highly attenuated scintillation crystals are employed, which may otherwise be incompatible with edge readout methods. By arranging photosensors on the side of a rectangular prism, light localized to the scintillation can be efficiently captured. Furthermore, for scintillators with lower density and stopping power, longer crystals may be required to achieve similar conversion efficiencies (or stopping power) as shorter, denser crystals. Side photosensing offers an even greater advantage over edge readout, which creates a long optical path and reduces timing resolution, as the photon arrival time range is significantly reduced via side photosensing. Therefore, side photosensing (or readout) potentially allows the utilization of more economical or higher-performance scintillators that would otherwise be incompatible or undesirable for edge readout configurations.

[0174] FIG. 15A is a diagram of a detector module block of scintillation crystals 1500 having side-reading photosensors 1501-1504 and a thermal management hot plate 1575, according to one embodiment. The use of side photosensing via an array of photosensors 1501-1504 allows for the placement of the hot plate 1575 on the proximal end face 1511 of the block of scintillation crystals 1500 for heat conduction. The distal end face 1512 of the block of scintillation crystals 1500 is oriented to face radially into the bore of the scanner to receive gamma radiation therefrom. The hot plate 1575 conceptually represents a thermal conductor placed in contact with the proximal end face 1511 of each elongated scintillation crystal in the block of scintillation crystals 1500. The five reflective materials on each face of the elongated scintillation crystals may be thermally conductive in some embodiments.

[0175] In this manner, the hot plate 1575 can conduct heat through the reflective material on the proximal end face 1511 of each elongated scintillation crystal to cool the elongated scintillation crystal. The reflective material (e.g., a coating or thin layer of material) on the proximal end face 1511 of each elongated scintillation crystal may be thermally coupled to the reflective material on the other face of each elongated scintillation crystal. In such an embodiment, the elongated scintillation crystals (and the array of laterally disposed photosensors) may be cooled from all sides as heat is conducted through the reflective material to the hot plate 1575.

[0176] In some embodiments, the hot plate 1575 also functions as a reflector. For example, the hot plate 1575 can simultaneously serve as a reflector by polishing, sandblasting, or other finishing processes to provide specular or diffuse reflection. In some embodiments, the hot plate 1575 may be connected to the proximal end face 1511 of the block of elongated scintillation crystals via a thermally conductive paste, or via welding each elongated scintillation crystal to the reflector, or both. Examples of thermally conductive and reflective materials that can be used to form the reflector (or reflective coating) and / or the hot 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 hot plate 1575 can be slotted or otherwise have holes or electrical conduits to support signal entry and exit from the crystal block volume to support lateral readout. In some embodiments, the hot plate is formed as part of the reflector material and 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 formed to wrap around the distal end 1512 and the sides of the block of scintillation crystal 1500 .

[0177] In the illustrated embodiment, the two-dimensional array of elongated scintillation crystals includes a 4x4 array of a total of 16 elongated scintillation crystals. It is understood that any number of elongated scintillation crystals in a one- or two-dimensional array may be used to form a block of scintillation crystals in a detector module. In the illustrated embodiment, each elongated scintillation crystal may be composed of six discrete arrays of photosensors, each including multiple photodiode sensors. In an alternative embodiment, a single elongated array of photosensors may be disposed on the axially facing side of each elongated scintillation crystal.

[0178] The use of a thermally conductive material in direct or indirect contact with the elongated scintillation crystal and the adjacent photosensor may stabilize and / or increase the scintillator's luminescence and photosensor gain through cooling. Thermal fluctuations in the elongated scintillation crystal and / or photosensor 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 for installation in a controlled room. In certain embodiments, the PET scanner may utilize water at or below ambient temperature for use in the scanner. In other examples, an air-to-air heat exchanger may be used to remove convective electron heat from within the scanner to the cooler ambient environment.

[0179] In other embodiments, an air-to-liquid heat exchanger may be utilized to transfer heat via liquid conduction (e.g., via water) from a thermally connected plate (such as hot plate 1575) to the external ambient air (i.e., outside the scanner housing). In still other embodiments, the PET scanner may have access to a liquid-to-liquid heat exchanger for transferring heat. In many cooling configurations, heat conduction or convection is only applied generally or broadly within the scanner housing, without direct or significant heat conduction (e.g., between laterally adjacent crystal faces within individual detector modules) to or from the arrays of elongated scintillation crystals within the individual detector modules.

[0180] Furthermore, in many embodiments, the PET scanner may not have a cooling system configured to cool the system below room temperature. One reason that a PET scanner may be configured to operate at or near ambient temperature is due to the risk of condensation from very cold active cooling surfaces damaging components therein. Therefore, the scanner thermal control system may monitor dew point conditions and their temperature thresholds to maintain the cooling surfaces above condensation temperatures.

[0181] However, some embodiments of the presently described systems and methods utilize a cooling system intended to directly cool the elongated scintillation crystals and / or photosensors to improve thermal regulation and output resolution (e.g., scintillator light and photosensor signal). Additionally, some embodiments of the presently described systems and methods cool the elongated scintillation crystals to below ambient temperature to increase scintillator light output and photosensor gain. By insulating cooled surfaces having temperatures below the dew point such that exterior surfaces exposed to ambient conditions are above the dew point threshold, the elongated scintillation crystals and / or photosensors within the detector modules of a PET scanner can be operated at temperatures substantially below ambient room temperature. In some examples, each gamma radiation detector module is sealed with an insulating layer to encapsulate the elongated scintillation crystals, the array of photosensors, and the reflective material. In some embodiments, wiring passes through the sealed enclosure.

[0182] FIG. 15B is a block diagram of a scintillation crystal 1590 having photosensors 1501, 1502, 1503, and 1504, intermediate substrates 1514, 1515, 1516, and 1517, and connector arrays 1524, 1525, 1526, and 1527, according to one embodiment. In various embodiments, the detector module receives gamma photons at its distal end face 1512. The gamma photons are scintillated within one (or more) of the elongated scintillation crystals, generating low-frequency scintillation photons. The scintillation photons are received by side photosensor(s) associated with the elongated scintillation crystals. Electrical signals generated by the side photosensors are transmitted by an electrical intermediate substrate to a connector array (e.g., a ball grid array). In the illustrated embodiment, each of the 16 elongated scintillation crystals is associated with an intermediate substrate disposed laterally with six arrays of photosensors. Thus, scintillation crystal block 1590 includes 16 elongated scintillation crystals, 16 intermediate substrates, and an array of 96 photosensors (e.g., 96 silicon photomultipliers, each containing multiple photodiodes), not all of which are numbered.

[0183] In some embodiments, a single laterally disposed intermediate substrate may be associated with multiple elongated scintillation crystals. For example, the intermediate substrate may be shared by two adjacent scintillation crystals. As another example, each row of four scintillation crystals may share a single intermediate substrate (e.g., spanning all four sides of the scintillation crystals). In such an example, a block 1590 of scintillation crystals would include 16 elongated scintillation crystals, four intermediate substrates, and an array of 96 photosensors (each of which may include multiple individual photodiodes). Furthermore, a single laterally disposed intermediate substrate may integrate photosensors on two opposing sides, thereby optically interlocking with two or more crystals on each side of the intermediate substrate.

[0184] In various embodiments, a PET scanning system includes one or more rings of detector modules arranged with photosensors and electrical intermediate substrates oriented axially relative to the rings. In such embodiments, the axial width of each ring of detector modules includes the sum of the widths of the elongated scintillation crystals in the axial direction, the sum of the widths of the axially aligned photosensor arrays on each side of the elongated scintillation crystals, the sum of the widths of the axially aligned reflectors or reflective coatings, and the sum of the widths of the axially aligned intermediate substrates connected to the photosensors. The illustrated embodiment is simplified to include only four elongated scintillation crystals in the axial direction, with each elongated scintillation crystal having a square end (e.g., each elongated scintillation crystal has the same height and width).

[0185] Thus, the width of the block 1590 of scintillation crystals in the axial direction is equal to four times the sum of the widths of the intermediate substrate, photosensor, and 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 lateral 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.

[0186] 15C is a perspective view of the detector block of scintillation crystals 1590 of FIG. 15B with hot plates 1541, 1542, 1543, and 1545 added to the proximal end face 1511 of each row of elongated scintillation crystals in the block of scintillation crystals 1590. Hot plates 1541, 1542, 1543, and 1545 may be configured and / or operate according to any of the various embodiments and variations described in conjunction with hot plate 1575 of FIG. 15A. The use of multiple hot plates 1541, 1542, 1543, and 1545 facilitates routing of signals through an intermediate substrate to the connector array and / or facilitates connection to the connector array.

[0187] 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 for connecting to a connector array 1520 on an intermediate substrate 1514 associated with each elongated scintillation crystal 1550. The intermediate substrate 1514 operates to transmit electrical signals from the multiple arrays of photosensors 1501 on each elongated scintillation crystal 1550 to the respective connector array 1520. A hot plate 1540 is provided to facilitate thermal cooling of the elongated scintillation crystals 1550, the photosensors 1501, and / or the electronic and conductive components within the intermediate substrate 1514.

[0188] In the illustrated embodiment, each scintillation crystal 1550 is associated with a separate intermediate substrate 1514. In some embodiments, a single intermediate substrate may be associated with multiple scintillation crystals. For example, a single intermediate substrate may have a width that spans the sides of multiple scintillation crystals. A single intermediate substrate may be associated with, for example, four or five scintillation crystals.

[0189] 7 and 12, a first channel of processing circuitry 1580 includes energy and timing circuits for processing signals from elongated scintillation crystals 1550 with dot shading. A second channel of processing circuitry 1580 includes energy and timing circuits for processing signals from elongated scintillation crystals 1550 with crosshatch shading. A PET scanner can include multiple detector modules, each including a block or blocks of scintillation crystals 1590 with processing circuitry 1580. Image processing circuitry in the PET scanner can receive timing and energy signals from the dual channel processing circuitry 1580 of each detector module to reconstruct an image, as understood in the art.

[0190] 16 is a perspective view of a block 1600 of elongated scintillation crystals having double-sided intermediate substrates 1614 and 1615 connected to photosensor arrays of adjacent elongated scintillation crystals, according to one embodiment. As described in connection with other embodiments, hot plates 1641, 1643, and 1645 are spaced apart so as to contact proximal end faces 1611 of the block of elongated scintillation crystals 1600. The particular dimensions of hot plates 1641, 1643, and 1645 may be selected to facilitate routing of signals through intermediate substrates 1614 and 1615 via connector arrays 1620 and 1622.

[0191] In the illustrated example, block 1600 of elongated scintillation crystals includes 16 elongated scintillation crystals, including numbered elongated scintillation crystals 1650, 1651, 1652, 1653, and 1654, and eight intermediate substrates, including numbered intermediate substrates 1614 and 1615. It is understood that a block of scintillation crystals may include any number of elongated scintillation crystals in a one-dimensional or two-dimensional array having an even or odd number of elongated scintillation crystals. Each elongated scintillation crystal has a reflective material or coating on five sides such that any optical radiation generated by a scintillation event within any elongated scintillation crystal is reflected for detection by its associated array of photosensors. In the illustrated example, each intermediate substrate has an array of photosensors on opposite sides of its axial inner diameter.

[0192] As shown, intermediate substrate 1614 has a first array of photosensors 1601 on a first side for detecting optical radiation (e.g., from scintillation events therein) within elongated scintillation crystal 1650. A second array of photosensors 1602 on a second side of intermediate substrate 1614 detects optical radiation within elongated scintillation crystal 1651. Electrical signals from the first and second arrays of photosensors 1601 and 1602 are transmitted, processed, and / or partially processed through intermediate substrate 1614. Intermediate substrate 1614 conveys the processed or partially processed electrical signals from photosensors 1601 and 1602 to a connector array (not visible) at proximal end 1611 of block 1600.

[0193] Revisiting the function of the double-sided intermediate substrate, intermediate substrate 1615 has a first array of photosensors 1603 on a first side for detecting optical radiation in elongated scintillation crystal 1652. A second array of photosensors 1604 on a second side of intermediate substrate 1615 detects optical radiation in elongated scintillation crystal 1653. Electrical signals from the first and second arrays of photosensors 1603 and 1604 are transmitted through intermediate substrate 1615 and processed and / or partially processed. Intermediate substrate 1615 conveys the processed or partially processed electrical signals from photosensors 1603 and 1604 to connector array 1620 at proximal end 1611 of block 1600. In the illustrated embodiment, each of the first and second arrays of photosensors 1603 and 1604 includes six discrete arrays of photosensors. It is understood that each photosensor may include more or fewer discrete photosensors.

[0194] In the illustrated embodiment, block 1600 includes an even number of scintillation crystals in the direction of the axial bore, such that an even number of double-sided intermediate substrates are used. In some embodiments, the block of elongated scintillation crystals includes an odd number of scintillation crystals in the direction of the axial bore. In such embodiments, an odd number of intermediate substrates may be utilized, with the intermediate substrates on one end of the block being single-sided intermediate substrates each serving a single elongated scintillation crystal.

[0195] FIG. 17 illustrates a ring 1700 of detector modules 1790 (e.g., blocks) having axially oriented photosensors on the sides of elongated scintillation crystals, according to one embodiment. Each of the detector modules 1790, while simplified, is intended to represent any of the various embodiments of detector modules described herein. For example, although not shown in each example, each detector module 1790 can include an intermediate substrate on the axially aligned side of the photosensor, which in turn is disposed on the axially aligned side of each elongated scintillation crystal. Each detector module 1790 is positioned with its distal end face 1712 facing into 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 bore) and a transverse direction around the circumference of the ring 1700.

[0196] In the illustrated embodiment, the elongated scintillation crystals of each detector module 1790 or "block" are arranged seamlessly in the axial direction of the detector ring with no substantial gaps between them, such that two adjacent elongated scintillation crystals are separated from each other by (i) an array of photosensors on a first axially facing side of one of the adjacent elongated scintillation crystals, (ii) a reflective material on a second axially facing side of the other adjacent elongated scintillation crystal, and (iii) optionally, an intermediate substrate connected to the array of photosensors on the first axially facing side of each elongated scintillation crystal.

[0197] 18 shows a ring 1800 of detector modules 1890 having transversely oriented photosensors on sides of elongated scintillation crystals, according to one embodiment. Again, each of the detector modules 1890 in the ring 1800, while simplified, 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 intermediate substrate on the axially-aligned side of the photosensor, which in turn is disposed on the axially-aligned side of a respective elongated scintillation crystal.

[0198] The transverse axial orientation of the intermediate substrate and photosensors results in an axial width of ring 1800 that is smaller than axial width 1720 of ring 1700 of FIG. 17, even though the same number of detector modules are utilized to form the ring. However, the axial orientation of the intermediate substrate and photosensors results in an inner diameter 1810 of ring 1800 that is larger than inner diameter 1710 of ring 1700 of FIG. 17, even though the same number of detector modules are utilized. As described herein, the axial orientation of FIG. 17 results in a smaller inner diameter 1710 for a given number of detector modules. For reasons described herein, a smaller diameter may be advantageous in some applications.

[0199] 19 is a graph 1900 of sensitivity gain and fill factor loss of side photosensor readout for various crystal geometries using the edge readout as a baseline, according to one embodiment. As shown, the side sensitivity (compared to the edge readout), plotted by the dashed line (labeled Side Gain), increases approximately linearly with the length of the elongated scintillation crystal relative to the width of the edge of the elongated scintillation crystal (for an edge size of 2 mm). 2 , 3mm 2 , and 4 mm 2(Plotted for squares in ). Conversely, as the length-to-width aspect ratio increases (i.e., the crystal becomes thinner), the photosensor thickness (370 μm in the plot) increases, resulting in an increased fill factor loss, as shown by the dotted line (labeled Fill Factor Loss), and a decrease in lateral readout gain. Nevertheless, the sensitivity gain obtained by lateral photosensor readout outweighs the fill factor loss, as shown by the solid line in the plotted ratio (labeled Coupling Gain).

[0200] FIG. 20 is a graph 2000 of the overall lateral sensitivity gain (compared to edge readout) for rectangular prism scintillation crystals of various lengths (e.g., 10-30 mm long) with square end faces, according to various embodiments. As shown, longer / higher aspect ratios indicate higher side readout gain. Fill factor loss is greater for thinner crystals (e.g., less than about 2 mm), while short, wide crystals experience the least gain due to reduced lateral photosensor surface area relative to edge readout.

[0201] In the context of the information conveyed in FIGS. 19 and 20 , various embodiments of the presently described systems and methods include a detector module having a two-dimensional or three-dimensional array of scintillation crystals with dimensions that target a relatively higher gain than the gain of the end readout. In some embodiments, the dimensions of the elongated scintillation crystals are selected as a function of reducing or minimizing fill factor loss and increasing the total gain. For example, a two-dimensional array of elongated scintillation crystals within a detector module can utilize scintillation crystals with dimensions and length-to-width ratios that provide a fill factor greater than 85%. In various embodiments, each elongated scintillation crystal has a length between its distal and proximal end faces of between 10 and 30 millimeters, a length-to-width ratio in the axial direction of between 3 and 15, and a length-to-width ratio in the transverse direction of between 3 and 15. In various embodiments, the end faces of the elongated scintillation crystals are square, so that the axial width and the transverse width are equal.

[0202] FIG. 21 is a graph 2100 of the optical dynamic range of a silicon photomultiplier photosensor having 1600 single-photon avalanche photodiode microcells per square millimeter of rectangular prism scintillation crystals of various dimensions, according to one embodiment. The dynamic range is proportional to the lateral surface area of ​​the crystal. In various embodiments, each elongated scintillation crystal is associated with an array of photosensors. The photosensor array can be composed of one or more silicon photomultiplier photosensors arranged on the axially-facing side of the elongated scintillation crystal.

[0203] 22 is a graph 2200 of light output versus temperature for the scintillator LYSO:Ce, according to one embodiment. According to various embodiments, the elongated scintillation crystals may be maintained at a temperature below ambient temperature to achieve higher light yield output. For example, in some embodiments, an increase in light output of approximately 4% is achieved by operating the elongated scintillation crystals in each detector module at −40° C. instead of 21° C., which is approximately room temperature. Of note, the voltage variance per Hertz bandwidth of thermally induced electronic noise of an array of photosensors, referred to in the art as Johnson-Nyquist noise, is given by:

[0204]

number

[0205] where: k B = Boltzmann's constant (Joules / Kelvin), T = absolute temperature of the resistor (Kelvin), and R = quenching resistance of the photosensor or other equivalent resistance.

[0206] Therefore, when the temperature is reduced, all other factors being constant, there is a reduction in thermal noise given by:

[0207]

number

[0208] where: T1 = upper operating temperature (in Kelvin), T2 = lower operating temperature (Kelvin).

[0209] Thus, for example, lowering the temperature from 21°C to -40°C reduces the root-mean-square (RMS) noise by approximately 11% in total. For example, increasing the light output by approximately 4% improves the SNR by approximately 17%. In addition to the thermal noise reduction of silicon photomultiplier photosensors at low temperatures, there is also an improvement in photoelectric gain. According to various embodiments, detector modules with side-sensitive photons using subambient-cooled scintillation crystals provide improved light collection efficiency and include thermal management components (such as hot plates) to efficiently reduce temperature for greater scintillator light output, increased photosensor transmission, and reduced noise.

[0210] In some embodiments, condensation on cooled surfaces in a PET scanning environment is eliminated or reduced by controlling the ambient environment and / or insulating the cold surfaces from warmer, moist air. At a surface temperature of -40°C, condensation is approximately 119.3 mg / m under standard conditions. 3 (or 79 ppm) absolute humidity can result in 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, such as nitrogen. The water vapor concentration in the sealed, positively pressurized chamber can be reduced to less than 1 ppm. In some embodiments, a vacuum pump is used to evacuate the air in the sealed chamber, reducing the pressure to approximately 10 mbar or less. Dry nitrogen gas is used to purge the chamber. The insulated chamber can then be pressurized above atmospheric pressure using dry nitrogen gas.

[0211] This disclosure has been made with reference to various exemplary embodiments, including the best mode. However, those skilled in the art will recognize that changes and modifications can be made to the exemplary embodiments without departing from the scope of the disclosure. While the principles of the disclosure have been illustrated in various embodiments, many changes in structure, arrangement, proportions, elements, materials, and components may be made to adapt to particular environments and / or operating requirements without departing from the principles and scope of the disclosure. These and other changes or modifications are intended to be included within the scope of the disclosure.

[0212] Clause 1. A gamma radiation detector module comprises an array of scintillation crystals disposed on a detector, each scintillation crystal including a proximal end face, a distal end face oriented into the detector to receive gamma photons, and four sides including a first side, a second side, a third side, and a fourth side; an array of photosensors disposed along the first side of each scintillation crystal to detect scintillation photons; and reflective material disposed on the proximal end face, the distal end face, and the second, third, and fourth sides of each scintillation crystal to internally reflect the scintillation photons.

[0213] Clause 2. A gamma radiation detector module as described in clause 1, wherein each scintillation crystal comprises an elongated scintillation crystal, the array of elongated scintillation crystals configured to be arranged on a detector ring, the four sides of each scintillation crystal including two axially oriented sides and two cross-axially oriented sides, and a distal end face oriented radially into the detector ring to receive gamma photons.

[0214] Clause 3. A gamma radiation detector module as described in clause 2, wherein the first side is axially oriented and the photosensor is disposed along the axially oriented side.

[0215] Clause 4. A gamma radiation detector module as described in clause 1, wherein the array of elongated scintillation crystals is configured to be disposed on a detector panel that operates in cooperation with an opposing detector panel.

[0216] Clause 5. A gamma radiation detector module according to clause 1, wherein each scintillation crystal is in the shape of an elongated polygonal prism having N sides, where N is an integer value.

[0217] Clause 6. The gamma radiation detector module of clause 1, wherein each scintillation crystal comprises a cubic-shaped scintillation crystal.

[0218] Clause 7. The gamma radiation detector module of clause 1, wherein each scintillation crystal comprises an elongated scintillation crystal in the shape of an elongated rectangular parallelepiped.

[0219] Clause 8. The gamma radiation detector module of clause 7, wherein a distal end face of each elongated scintillation crystal is square.

[0220] Clause 9. A gamma radiation detector module as described in clause 8, wherein each elongated scintillation crystal has a length between its distal end face and its proximal end face of between 10 millimeters and 30 millimeters, and the ratio of the length of each elongated scintillation crystal to its width is between three and ten (3-10).

[0221] Clause 10. The gamma radiation detector module of clause 3, wherein the array of elongated scintillation crystals comprises a one-dimensional array of elongated scintillation crystals arranged along the axial direction of the detector ring.

[0222] Clause 11. A gamma radiation detector module as described in clause 10, wherein the elongated scintillation crystals are closely spaced in the axial direction of the detector ring, and a given elongated scintillation crystal is separated from an axially adjacent elongated scintillation crystal by (i) a photosensor on a first axially facing side of the given elongated scintillation crystal and (ii) a reflector on a second axially facing side of the axially adjacent elongated scintillation crystal.

[0223] Clause 12. A gamma radiation detector module as described in clause 3, wherein the array of elongated scintillation crystals comprises a two-dimensional array of elongated scintillation crystals, the two-dimensional array of elongated scintillation crystals comprising M elongated scintillation crystals in an axial direction of the detector ring and N elongated scintillation crystals in a transaxial direction of the detector ring, wherein M and N are each integer values.

[0224] Clause 13. A gamma radiation detector module as described in Clause 12, wherein the elongated scintillation crystals are arranged adjacent to each other without gaps in the axial and cross-axial directions of the detector ring, and a given elongated scintillation crystal is separated from an axially adjacent elongated scintillation crystal by (i) a photosensor on a first axially facing side of the given elongated scintillation crystal and (ii) a reflective material on a second axially facing side of the axially adjacent elongated scintillation crystal, and wherein the given elongated scintillation crystal and an elongated scintillation crystal adjacent to the cross-axial direction are separated from each other in the cross-axial direction by the reflective material.

[0225] Clause 14. A gamma radiation detector module as described in clause 13, wherein the thickness of the photosensor and reflective material between adjacent elongated scintillation crystals in the axial direction of the detector ring is less than 500 micrometers, and the thickness of the reflective material between adjacent elongated scintillation crystals in the cross-axial direction of the detector ring is less than 100 micrometers.

[0226] Clause 15. A gamma radiation detector module as described in clause 14, wherein each elongated scintillation crystal has a length between its distal end face and its proximal end face of between 10 millimeters and 30 millimeters, each elongated scintillation crystal has an axial length to width ratio of between 3 and 10, and each elongated scintillation crystal has a cross-axial length to width ratio of between 3 and 10.

[0227] Clause 16. A gamma radiation detector module according to clause 13, wherein the fill factor of the area of ​​the distal end face of the elongated scintillation crystal relative to the total area is greater than 85%.

[0228] Clause 17. The gamma radiation detector module of clause 1, wherein the reflective material comprises a thermally conductive reflective material.

[0229] 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.

[0230] Clause 19. The gamma radiation detector module of clause 17, further comprising a thermally conductive plate coupled to a proximal end surface of at least a portion of the scintillation crystal.

[0231] Clause 20. A gamma radiation detector module as described in clause 1, further comprising a sealed housing enclosing the scintillation crystals, photosensors, and reflective material, and a routing guide wire for routing an electrical signal from the photosensor of each scintillation crystal through the sealed housing from its first side toward its proximal end face.

[0232] Clause 21. The gamma radiation detector module of clause 1, further comprising a routing guide wire for routing an electrical signal from the photosensor of each scintillation crystal from its first side toward its proximal end face.

[0233] Clause 22. A gamma radiation detector module as described in clause 3, further comprising an intermediate substrate connected to a photosensor along a first axially facing side of each elongated scintillation crystal, the intermediate substrate operable to transmit an electrical signal from the photosensor of each elongated scintillation crystal toward the proximal end face of each elongated scintillation crystal.

[0234] Clause 23. A gamma radiation detector module as described in clause 22, wherein the array of elongated scintillation crystals comprises a two-dimensional array of elongated scintillation crystals, the two-dimensional array of elongated scintillation crystals comprising M elongated scintillation crystals in an axial direction of the detector ring and N elongated scintillation crystals in a transaxial direction of the detector ring, wherein M and N are each integer values.

[0235] Clause 24. A gamma radiation detector module as described in Clause 23, wherein the elongated scintillation crystals are arranged adjacent to each other without gaps in the axial and cross-axial directions of the detector ring, and a given elongated scintillation crystal is separated from an axially adjacent elongated scintillation crystal by (i) a photosensor on a first axially facing side of the given elongated scintillation crystal and (ii) a reflector on a second axially facing side of the axially adjacent elongated scintillation crystal, and the given elongated scintillation crystal and an elongated scintillation crystal adjacent to the cross-axial direction are separated from each other in the cross-axial direction by the reflector.

[0236] Clause 25. A gamma radiation detector module as described in clause 1, wherein the array of photosensors in each scintillation crystal includes at least one two-dimensional array of individual photodiodes.

[0237] Clause 26. A gamma radiation detector module as described in clause 25, wherein each photodiode comprises a single photon avalanche diode (SPAD) operating in Geiger mode.

[0238] Clause 27. A gamma radiation detector module as described in clause 1, wherein the array of photosensors for each scintillation crystal includes a plurality of separate photosensors for providing separate photosensor measurements corresponding to separate depth ranges between the proximal end face and the distal end face of each scintillation crystal.

[0239] Clause 28. A gamma radiation detector module as described in clause 27, wherein the array of photosensors in each scintillation crystal comprises a plurality of two-dimensional arrays of individual photodiodes.

[0240] Clause 29. A gamma radiation detector module as described in clause 28, wherein each photodiode comprises a single photon avalanche diode (SPAD) operating in Geiger mode.

[0241] Clause 30. A gamma radiation detector module as described in clause 29, wherein each photosensor includes a silicon photomultiplier (SiPM), and each scintillation crystal is associated with a plurality of silicon photomultipliers extending along a first side thereof.

[0242] Clause 31. The gamma radiation detector module of clause 30, further comprising a processing circuit including: a plurality of energy circuits, each energy circuit configured to generate an energy signal as a sum of energy detected by one or more silicon photomultiplier tubes; and a plurality of timing signal circuits, each timing signal circuit configured to generate a separate timing signal for scintillation photons detected by one or more silicon photomultiplier tubes.

[0243] Clause 32. A gamma radiation detector module as described in clause 31, wherein the processing circuit is a dual-channel high-speed circuit, each channel including circuitry for digitally encoding one or more timing and energy signals of the silicon photomultiplier tube.

[0244] 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 adjacent scintillation crystals.

[0245] Clause 34. A gamma radiation detector module as described in clause 31, wherein said processing circuitry includes a different timing signal circuit for each set of three silicon photomultiplier tubes.

[0246] Clause 35. A gamma radiation detector module as described in clause 31, wherein the processing circuitry includes a separate energy circuit for each silicon photomultiplier tube.

[0247] Clause 36. A gamma radiation detector module as described in clause 31, wherein said processing circuitry includes a separate energy circuit for each set of six silicon photomultiplier tubes.

[0248] Clause 37. A gamma radiation detector module according to clause 27, wherein each scintillation crystal is divided along its length into a plurality of divided volumes.

[0249] Clause 38. A gamma radiation detector module according to clause 37, wherein each divided volume of each scintillation crystal has a different crystal time constant.

[0250] Clause 39. A gamma radiation detector module as described in clause 38, wherein processing circuitry associated with the detector module determines from which segmented volume a scintillation photon originated based on the time constant of the detected scintillation photon.

[0251] Clause 40. A gamma radiation detector module as described in clause 27, wherein each scintillation crystal is divided along its length into a plurality of segmented volumes having a reflective layer disposed between adjacent faces of the segmented volumes, each segmented volume being associated with a separate array of photosensors, such that scintillation photons generated within the segmented volume are reflected within the segmented volume for detection by a corresponding array of photosensors.

[0252] Clause 41. A positron emission tomography (PET) scanning system comprising: a plurality of gamma radiation detector modules arranged to form a detector ring, each detector module being an array of elongated scintillation crystals, each elongated scintillation crystal including a proximal end face, two axially oriented side faces, two cross-axially oriented side faces, and a distal end face oriented radially of the detector ring for receiving gamma photons; an array of photosensors disposed along a first of the axially oriented side faces of each elongated scintillation crystal for detecting scintillation photons; a reflective material disposed on a second of the proximal end face, distal end face, radially oriented side face, and axially oriented side face of each elongated scintillation crystal for internally reflecting the scintillation photons; a cooling system for cooling the detector modules; and an imaging system connected to the detector modules for generating images based on electronic output from the detector modules.

[0253] Clause 42. A PET scanning system as described in Clause 41, wherein the detector ring is formed as a plurality of axially aligned rings of detector modules, each of the axially aligned rings including at least four detector modules.

[0254] Clause 43. The PET scanning system of clause 42, wherein the detector ring has a diameter 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.

[0255] Clause 44. A PET scanning system according to clause 41, wherein each of the elongated scintillation crystals is in the shape of an elongated N-sided polygonal prism, where N is an integer value.

[0256] Clause 45. A PET scanning system according to clause 41, wherein each of the elongated scintillation crystals is in the shape of an elongated rectangular prism.

[0257] Clause 46. The PET scanning system of clause 45, wherein the distal end face of each elongated scintillation crystal is square.

[0258] Clause 47. A PET scanning system as described in clause 46, wherein each elongated scintillation crystal has a length between its distal end face and its proximal end face of between 10 millimeters and 30 millimeters, and the ratio of the length of each elongated scintillation crystal to its width is between three and ten (3-10).

[0259] Clause 48. A PET scanning system according to clause 41, wherein the array of elongated scintillation crystals in each detector module comprises a one-dimensional array of elongated scintillation crystals arranged along the axial direction of the detector ring.

[0260] Clause 49. A PET scanning system as described in Clause 48, wherein the elongated scintillation crystals of each detector module are closely spaced in the axial direction of the detector ring such that they are separated from axially adjacent elongated scintillation crystals by (i) an array of photosensors on a first axially facing side of a given elongated scintillation crystal and (ii) a reflective material on a second axially facing side of an axially adjacent elongated scintillation crystal.

[0261] Clause 50. A PET scanning system as described in Clause 41, wherein the array of elongated scintillation crystals in each detector module comprises a two-dimensional array of elongated scintillation crystals, and the two-dimensional array of elongated scintillation crystals in each detector module comprises M elongated scintillation crystals in the axial direction of the detector ring and N elongated scintillation crystals in the transaxial direction of the detector ring, and M and N are each integer values.

[0262] Clause 51. The PET scanning system of clause 50, wherein the elongated scintillation crystals of each detector module are arranged adjacent to one another with no gaps in the axial and cross-axial directions of the detector ring, and a given elongated scintillation crystal is separated from an axially adjacent elongated scintillation crystal by (i) an array of photosensors on a first axially facing side of the given elongated scintillation crystal and (ii) a reflective material on a second axially facing side of the axially adjacent elongated scintillation crystal, such that the given elongated scintillation crystal and its cross-axially adjacent elongated scintillation crystals are separated from one another in the cross-axial direction by the reflective material.

[0263] Clause 52. A PET scanning system as described in Clause 51, wherein the thickness of the array of photosensors and reflective material between adjacent elongated scintillation crystals in each detector module in the axial direction of the detector ring is less than 500 micrometers, and the thickness of the reflective material between adjacent elongated scintillation crystals in each detector module in the cross-axial direction of the detector ring is less than 100 micrometers.

[0264] Clause 53. A PET scanning system as described in clause 52, wherein each elongated scintillation crystal has a length between its distal end face and its proximal end face of between 10 millimeters and 30 millimeters, the ratio of length to width in the axial direction of each elongated scintillation crystal is between 3 and 10, and the ratio of length to width in the cross-axial direction of each elongated scintillation crystal is between 3 and 10.

[0265] Clause 54. A PET scanning system according to clause 51, wherein the fill factor of the area of ​​the distal end face of the elongated scintillation crystal in each detector module relative to the total area of ​​each detector module is greater than 85%.

[0266] Clause 55. The PET scanning system of clause 41, wherein the reflective material comprises a thermally conductive reflective material.

[0267] 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.

[0268] Clause 57. The PET scanning system of clause 56, further comprising a heat conductor plate coupled to a proximal end surface of at least some of the elongated scintillation crystals in the array of elongated scintillation crystals of each detector module.

[0269] Clause 58. A PET scanning system as described in Clause 41, wherein each detector module further comprises a sealed housing enclosing the elongated scintillation crystal, the array of photosensors, and the reflective material, and a routing guide wire for routing electrical signals from the array of photosensors of each elongated scintillation crystal from its first side toward its proximal end face and through the sealed housing to the imaging system.

[0270] Clause 59. The PET scanning system of clause 41, further comprising a routing guide wire for routing an electrical signal from the photosensor of each elongated scintillation crystal of each detector module from its first side toward its proximal end face.

[0271] Clause 60. A PET scanning system as described in Clause 41, further comprising an intermediate substrate connected to the array of photosensors along a first axially facing side of each elongated scintillation crystal of each detector module, the intermediate substrate operable to transmit electrical signals from the array of photosensors of each elongated scintillation crystal toward the proximal end face of each elongated scintillation crystal.

[0272] Clause 61. A PET scanning system as described in Clause 60, wherein the array of elongated scintillation crystals in each detector module comprises a two-dimensional array of elongated scintillation crystals, and the two-dimensional array of elongated scintillation crystals in each detector module comprises M elongated scintillation crystals in the axial direction of the detector ring and N elongated scintillation crystals in the transaxial direction of the detector ring, and M and N are each integer values.

[0273] Clause 62. The PET scanning system of clause 61, wherein the elongated scintillation crystals of each detector module are arranged adjacent to each other with no gaps in the axial and cross-axial directions of the detector ring, and a given elongated scintillation crystal and an axially adjacent elongated scintillation crystal are separated from each other in the axial direction by (i) an array of photosensors on a first axially oriented side of the given elongated scintillation crystal, (ii) an intermediate substrate of the given elongated scintillation crystal, and (iii) a reflective material on a second axially oriented side of the axially adjacent elongated scintillation crystal, and the given elongated scintillation crystal and an axially adjacent elongated scintillation crystal are separated from each other in the cross-axial direction by the reflective material. Clause 63. A PET scanning system according to clause 41, wherein the array of photosensors of each elongated scintillation crystal comprises a two-dimensional array of individual photodiodes.

[0274] Clause 64. A PET scanning system as described in Clause 41, wherein the array of photosensors of each elongated scintillation crystal includes a plurality of photosensors for providing distinct photosensor measurements corresponding to distinct depth ranges between the proximal end face and the distal end face of each elongated scintillation crystal.

[0275] Clause 65. A PET scanning system according to clause 64, wherein the array of photosensors of each elongated scintillation crystal comprises a two-dimensional array of individual photodiodes.

[0276] Clause 66. The PET scanning system of clause 65, wherein each photodiode comprises a single photon avalanche diode (SPAD).

[0277] Clause 67. A PET scanning system as described in Clause 66, wherein each photosensor includes a silicon photomultiplier (SiPM), and each elongated scintillation crystal is associated with a plurality of silicon photomultipliers extending along its first axially oriented side.

[0278] Clause 68. The PET scanning system of clause 67, further comprising: an energy circuit for each elongated scintillation crystal of each detector module, the energy circuit for generating an energy signal as a sum of the energies detected by the array of photosensors of each respective elongated scintillation crystal; and a timing signal circuit for each elongated scintillation crystal, the timing signal circuit for generating a separate timing signal for the scintillation photons detected by each individual photosensor of each respective elongated scintillation crystal.

[0279] Clause 69. A PET scanning system as described in clause 64, wherein each elongated scintillation crystal is divided along its length into a plurality of segmented volumes having a reflective layer disposed between adjacent faces of the segmented volumes, each segmented volume being associated with a separate array of photosensors, such that scintillation photons generated within the segmented volume are reflected within the segmented volume for detection by a corresponding array of photosensors.

[0280] Clause 70. The PET scanning system of clause 64, further comprising high speed circuitry connected to the array of photosensors in each elongated scintillation crystal of each detector module, for digitally encoding scintillation events detected in each elongated scintillation crystal as a timing signal for each of the plurality of photosensors associated with each elongated scintillation crystal, and an energy signal summing the total energy detected by the plurality of photosensors associated with each elongated scintillation crystal.

[0281] Clause 71. A PET scanning system as described in clause 70, wherein the high speed circuitry is a dual channel high speed circuitry, each channel including circuitry for digitally encoding timing and energy signals for a plurality of photosensors associated with each elongated scintillation crystal.

[0282] Clause 72. The PET scanning system of clause 71, wherein each elongated scintillation crystal in the array of elongated scintillation crystals in each detector module uses different high speed circuitry than adjacent elongated scintillation crystals.

[0283] Clause 73. A gamma radiation detector module comprising: an array of elongated scintillation crystals arranged on a detector ring, each elongated scintillation crystal including a proximal end face, two axially oriented side faces, two cross-axially oriented side faces, and a distal end face oriented radially into the detector ring to receive gamma photons; an array of photosensors having detection faces for detecting scintillation photons, the detection faces of the photosensor array being disposed along the length of a first of the axially oriented side faces of each elongated scintillation crystal; and an intermediate substrate connected to an output face of the photosensor array along the length of the first axially oriented side face of each elongated scintillation crystal, the intermediate substrate operable to transmit an electrical signal from the output face of the photosensor array of each elongated scintillation crystal towards the proximal end face of each elongated scintillation crystal.

[0284] Clause 74. A gamma radiation detector module as described in clause 73, wherein the array of photosensors in each elongated scintillation crystal includes a two-dimensional array of single photon avalanche diodes (SPADs), and the intermediate substrate includes bias circuitry that causes the two-dimensional array of SPADs in each elongated scintillation crystal to operate in Geiger mode.

[0285] Clause 75. A gamma radiation detector module as described in clause 73, wherein the fill factor of the area of ​​the distal end face of the elongated scintillation crystal relative to the total distal end face area including the distal end face of the array of photosensors of each elongated scintillation crystal and the intermediate substrate is greater than 85%.

[0286] Clause 76. A gamma radiation detector module as described in Clause 73, comprising reflective material disposed on a proximal end face, a distal end face, two cross-axial side faces, and a second of the axial side faces of each elongated scintillation crystal, the reflective material operating to internally reflect scintillation photons generated within each elongated scintillation crystal.

[0287] Clause 77. A gamma radiation detector module as described in clause 76, wherein each elongated scintillation crystal is divided along its length into a plurality of segmented volumes having reflective layers disposed between adjacent faces of the segmented volumes, and each segmented volume of each elongated scintillation crystal is associated with a unique array of photosensors, such that scintillation photons generated within the segmented volume are reflected within the segmented volume for detection by the corresponding array of photosensors.

[0288] Clause 78. A gamma radiation detector module according to clause 76, wherein each of the elongated scintillation crystals is in the shape of an elongated N-sided polygonal prism, where N is an integer value.

[0289] Clause 79. A gamma radiation detector module according to clause 76, wherein each of the elongated scintillation crystals is in the shape of an elongated rectangular parallelepiped.

[0290] Clause 80. A gamma radiation detector module according to clause 79, wherein the distal end face of each elongated scintillation crystal is square.

[0291] Clause 81. A gamma radiation detector module as described in clause 80, wherein each elongated scintillation crystal has a length between its distal end face and its proximal end face of between 10 millimeters and 30 millimeters, and the ratio of the length of each elongated scintillation crystal to its width is between three and ten (3-10).

[0292] Clause 82. A gamma radiation detector module according to clause 76, wherein the array of elongated scintillation crystals comprises a one-dimensional array of elongated scintillation crystals arranged along the axial direction of the detector ring.

[0293] Clause 83. The gamma radiation detector module of clause 82, wherein the elongated scintillation crystals are seamlessly arranged in the axial direction of the detector ring without gaps, such that a given elongated scintillation crystal and an axially adjacent elongated scintillation crystal are separated from each other by (i) an array of photosensors on a first axially oriented side of the given elongated scintillation crystal, (ii) an intermediate substrate of the given elongated scintillation crystal, and (iii) a reflective material on a second axially oriented side of the axially adjacent elongated scintillation crystal; Clause 84. A gamma radiation detector module as described in clause 76, wherein the array of elongated scintillation crystals comprises a two-dimensional array of elongated scintillation crystals, the two-dimensional array of elongated scintillation crystals comprising M elongated scintillation crystals in an axial direction of the detector ring and N elongated scintillation crystals in a transaxial direction of the detector ring, wherein M and N are each integer values.

[0294] Clause 85. A gamma radiation detector module according to clause 84, wherein the elongated scintillation crystals are arranged adjacent to each other without gaps in the axial and cross-axial directions of the detector ring, such that a given elongated scintillation crystal and an axially adjacent elongated scintillation crystal are separated from each other in the axial direction by (i) an array of photosensors on a first axially oriented side of the given elongated scintillation crystal, (ii) an intermediate substrate of the given elongated scintillation crystal, and (iii) a reflective material on an axially oriented second side of the axially adjacent elongated scintillation crystal, such that the given elongated scintillation crystal and an axially adjacent elongated scintillation crystal are separated from each other in the cross-axial direction by the reflective material. Clause 86. A gamma radiation detector module as described in clause 85, wherein the thickness of the array of photosensors and reflective material between adjacent elongated scintillation crystals in the axial direction of the detector ring is less than 500 micrometers, and the thickness of the reflective material between adjacent elongated scintillation crystals in the direction perpendicular to the axial direction of the detector ring is less than 100 micrometers.

[0295] Clause 87. A gamma radiation detector module as described in clause 86, wherein each elongated scintillation crystal has a length between its distal end face and its proximal end face of between 10 millimeters and 30 millimeters, and each elongated scintillation crystal has a length to axial width ratio of between 3 and 10, and each elongated scintillation crystal has a length to axial width ratio of between 3 and 10.

[0296] Clause 88. A gamma radiation detector module according to clause 85, wherein the fill factor of the area of ​​the distal end face of the elongated scintillation crystal relative to the total area is greater than 85%.

[0297] Clause 89. A gamma radiation detector module as described in 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 heat conductor plate coupled to a proximal end surface of at least some of the elongated scintillation crystals in the array of elongated scintillation crystals.

[0300] Clause 92. The gamma radiation detector module of clause 73, further comprising a hermetically sealed housing enclosing the elongated scintillation crystal, the array of photosensors, and the intermediate substrate.

[0301] Clause 93. A gamma radiation detector module as described in clause 73, wherein the array of photosensors of each elongated scintillation crystal includes a plurality of photosensors for providing distinct photosensor measurements corresponding to distinct depth ranges between the proximal end face and the distal end face of each elongated scintillation crystal.

[0302] Clause 94. A gamma radiation detector module according to clause 93, wherein the array of photosensors of each elongated scintillation 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. A gamma radiation detector module as described in clause 95, wherein each photosensor includes a silicon photomultiplier (SiPM), and each elongated scintillation crystal is associated with a plurality of silicon photomultipliers extending along its first axially oriented side.

[0305] Clause 97. A circuit for processing nuclear scintillation events, comprising: a plurality of photosensors that detect scintillation photons generated by a scintillation event in a scintillation crystal and output an electrical signal; an energy circuit that generates a digital signal representative of the total energy detected by one or more photosensors of the plurality of photosensors during a detection time period; and a timing circuit that generates a digital signal having a distinct timing signal for the scintillation photons detected by the one or more photosensors during the detection time period.

[0306] Clause 98. The circuit of clause 97, wherein the photosensor includes a plurality of single photon avalanche diodes (SPADs) biased to operate in Geiger mode, the bias circuit being integrated in the intermediate substrate, and each of the plurality of photosensors being positioned on a side of the scintillation crystal.

[0307] Clause 99. The circuit of clause 97, wherein the energy circuit comprises an analog-to-digital converter, a multiplexer that selects between the output signals of each of the plurality of photosensors for conversion by the analog-to-digital converter, a signal delay path that connects one or more of the plurality of photosensors to the multiplexer, and digital control logic that controls the multiplexer's selection between the output signals of the plurality of photosensors.

[0308] Clause 100. The circuit of clause 97, wherein the timing circuit includes: an analog-to-digital converter for high-speed digital encoding of the capacitively isolated output of one or more photosensors of the plurality of photosensors; a transmission line connecting the capacitively isolated output of one or more photosensors of the plurality of photosensors to the analog-to-digital converter through an output diode; a flip-flop for selectively triggering the encoding of the capacitively isolated output of one or more of the plurality of photosensors by the analog-to-digital converter by biasing the output diode to an on state; and a comparator for comparing the capacitively isolated output of one or more of the plurality of photosensors with one of a low voltage threshold and a low current threshold, the output of the comparator selectively triggering the flip-flop.

[0309] Clause 101. The circuit of clause 100, wherein the output diode comprises a PIN diode, the PIN diode having at least one of a high energy bandgap semiconductor of 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.

[0310] Clause 102. The circuit of clause 100, wherein the output diode comprises a PIN diode including 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 bias 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 milliamps.

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

[0314] Clause 106. The circuit of clause 97, wherein the energy circuit comprises a first analog-to-digital converter, a multiplexer selecting between output signals of each of the plurality of photosensors for conversion by the first analog-to-digital converter, a signal delay path connecting each of the photosensors to the multiplexer, and digital control logic controlling the selection of the multiplexer between output signals of the plurality of photosensors, and wherein the timing circuit comprises a second analog-to-digital converter for high speed digital encoding of the capacitively isolated output of one or more of the plurality of photosensors, a path connecting the capacitively isolated output derived from one or more of the plurality of photosensors to the analog-to-digital converter through an output diode, a flip-flop selectively triggering encoding of the capacitively isolated output derived from one or more of the plurality of photosensors by the second analog-to-digital converter by biasing the output diode to an on state, and a comparator comparing the output signal of one or more of the plurality of photosensors to a threshold, the output of the comparator selectively triggering the flip-flop to bias the output diode to an on state.

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

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

[0317] Clause 109. The circuit of clause 106, wherein the output diode comprises a PIN diode comprising 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 bias 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 milliamps.

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

[0321] Clause 113. The circuit of clause 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 milliamps, and a forward bias on-resistance of less than 10 ohms at an operating frequency of 1 gigahertz.

[0322] The present disclosure is to be regarded in an illustrative, not a restrictive, sense, and all such modifications are intended to be included within its scope. Similarly, advantages, other benefits, and solutions to problems have been described above with respect to various embodiments. However, the advantages, benefits, solutions to problems, and any element(s) that may cause or make more noticeable any advantage, benefit, or solution are not to be construed as critical, required, or essential features or elements. The present disclosure includes and encompasses at least the following claims: [Brief explanation of the drawings]

[0323] [Figure 1A] 1 is a perspective view of a block diagram of a portion of a positron emission tomography (PET) detector system according to one embodiment. [Figure 1B] FIG. 2 is another diagram of a block diagram of a portion of a PET detector system according to one embodiment. [Figure 2] FIG. 1 illustrates a positron source having multiple lines of response, according to one embodiment. [Figure 3]1 shows a tube-type block detector with four photomultiplier tubes according to one embodiment. [Figure 4] FIG. 1 is a block diagram of a ring of detector modules having multiple response lines, according to one embodiment. [Figure 5A] FIG. 1 is a block diagram of a ring of detector modules with a single off-center line of response (LoR), according to one embodiment. [Figure 5B] FIG. 1 is a block diagram of a ring of detector modules having multiple off-center response lines, according to one embodiment. [Figure 6] FIG. 10 illustrates a response line segmented in time and space according to one embodiment. [Figure 7] FIG. 1 is a simplified circuit diagram of a high speed diode network according to one embodiment. [Figure 8A] 1 is a two-port network lumped element model of a conductive transmission line, according to one embodiment. [Figure 8B] FIG. 1 illustrates a distribution model of lumped elements per unit length, according to one embodiment. [Figure 9] 1 shows a time domain graph of two nuclear pulse signals according to one embodiment. [Figure 10] 10 shows a graph of the energy spectral density of the two nuclear pulse signals of FIG. 9 according to one embodiment. [Figure 11] FIG. 1 is a hierarchical block diagram of a high speed diode network according to one embodiment. [Figure 12] 1 illustrates a detector module having a block of elongated scintillation crystals arranged in a two-dimensional array with an array of photosensors for lateral photosensor readout, according to one embodiment. [Figure 13A] FIG. 1 is a diagram of a single elongated scintillation crystal pixel with multiple arrays of photosensors for lateral photosensor readout, according to one embodiment. [Figure 13B] FIG. 2 is an exploded view of a single elongated scintillation crystal according to one embodiment. [Figure 13C]FIG. 1 is a diagram of a single elongated scintillation crystal pixel having an array of photosensors for side photosensor readout, a middle substrate, and a connector array, according to one embodiment. [Figure 14A] 1 illustrates a subdivided elongated scintillation crystal pixel having an array of photosensors for side photosensor readout, according to one embodiment. [Figure 14B] 1 illustrates an exploded view of a subdivided scintillation pixel according to one embodiment. [Figure 14C] FIG. 1 is a diagram of a sub-divided scintillation pixel having an array of photosensors for side photosensor readout, an intermediate substrate, and a connector array, according to one embodiment. [Figure 15A] 1 is a diagram of a block of elongated scintillation crystal with side-reading photosensors and thermal management plates, according to one embodiment. [Figure 15B] 1 is a diagram of a block of an elongated scintillation crystal with an array of photosensors, an intermediate substrate, and a connector, according to one embodiment. [Figure 15C] 1 is a diagram of a block of elongated scintillation crystal with an array of photosensors, an intermediate substrate, a connector array, and a thermal management plate, according to one embodiment. [Figure 15D] FIG. 2 is a diagram of a block of elongated scintillation crystals connected to a dual-channel processing circuit, according to one embodiment. [Figure 16] FIG. 1 is a diagram of a block of elongated scintillation crystals having a double-sided intermediate substrate connected to photosensor arrays of adjacent elongated scintillation crystals, according to one embodiment. [Figure 17] FIG. 1 illustrates a ring of detector modules having axially oriented photosensor arrays flanking an elongated scintillation crystal, according to one embodiment. [Figure 18] 1 illustrates a ring of detector modules having a transverse array of photosensors on the sides of an elongated scintillation crystal, according to one embodiment. [Figure 19]10 is a graph of sensitivity gain and fill factor loss for side photosensor readout, according to one embodiment. [Figure 20] 1 is a graph of the overall sensitivity gain of rectangular prism scintillation crystals of various lengths with square ends, according to one embodiment. [Figure 21] 1 is a graph of the photonic dynamic range of a silicon photomultiplier photosensor for various dimensions of an elongated rectangular prism scintillation crystal, according to one embodiment. [Figure 22] 1 is a graph of scintillator light output versus temperature, according to one embodiment.

Claims

1. an array of scintillation crystals disposed on the detector, each scintillation crystal including a proximal end face, a distal end face oriented into the detector to receive gamma photons, and four sides including a first side, a second side, a third side, and a fourth side; an array of photosensors disposed along the first side of each scintillation crystal for detecting scintillation photons; reflective material disposed on the proximal end face, the distal end face, and the second, third, and fourth side faces of each scintillation crystal for internally reflecting scintillation photons; 1. A gamma radiation detector module comprising:

2. each scintillation crystal comprises an elongated scintillation crystal; the array of elongated scintillation crystals is configured to be disposed on a detector ring; the four sides of each scintillation crystal include two axially oriented sides and two cross-axially oriented sides; the distal end face faces radially of the detector ring to receive the gamma photons; 2. The gamma radiation detector module of claim 1, wherein the first side faces axially and the photosensor is disposed along the axially facing side.

3. the array of elongated scintillation crystals comprises a two-dimensional array of elongated scintillation crystals; 3. The gamma radiation detector module of claim 2, wherein the two-dimensional array of elongated scintillation crystals includes M elongated scintillation crystals in an axial direction of the detector ring and N elongated scintillation crystals in a transaxial direction of the detector ring, where M and N are each integer values.

4. the elongated scintillation crystals are arranged adjacent to each other without gaps in the axial direction and the cross-axial direction of the detector ring; a given elongated scintillation crystal is separated from an axially adjacent elongated scintillation crystal by (i) a photosensor on a first axially oriented side of said given elongated scintillation crystal and (ii) a reflective material on a second axially oriented side of the axially adjacent elongated scintillation crystal; 4. The gamma radiation detector module of claim 3, wherein the given elongated scintillation crystal and an adjacent elongated scintillation crystal in the cross-axis direction are separated from each other in the cross-axis direction by the reflective material.

5. a thickness of the photosensor and reflector between adjacent elongated scintillation crystals in the axial direction of the detector ring is less than 500 micrometers; 5. The gamma radiation detector module of claim 4, wherein the thickness of the reflective material between adjacent elongated scintillation crystals in the cross-axial direction of the detector ring is less than 100 micrometers.

6. 6. The gamma radiation detector module of claim 1, wherein the reflective material comprises a thermally conductive reflective material.

7. an intermediate substrate connected to the photosensor along a side of each elongated scintillation crystal oriented in the first axis direction; 3. The gamma radiation detector module of claim 2, wherein the intermediate substrate is operable to transmit electrical signals from the photosensor of each elongated scintillation crystal toward the proximal end face of each elongated scintillation crystal.

8. the array of elongated scintillation crystals comprises a two-dimensional array of elongated scintillation crystals; 8. The gamma radiation detector module of claim 7, wherein the two-dimensional array of elongated scintillation crystals includes M elongated scintillation crystals in an axial direction of the detector ring and N elongated scintillation crystals in a transaxial direction of the detector ring, where M and N are each integer values.

9. the elongated scintillation crystals are arranged adjacent to each other without gaps in the axial direction and the cross-axial direction of the detector ring; a given elongated scintillation crystal is separated from an axially adjacent elongated scintillation crystal by (i) a photosensor on the first axially oriented side of the given elongated scintillation crystal and (ii) a reflective material on the second axially oriented side of the axially adjacent elongated scintillation crystal; 9. The gamma radiation detector module of claim 8, wherein the given elongated scintillation crystal and an axially adjacent elongated scintillation crystal are separated from each other in the cross-axial direction by the reflective material.

10. 6. A gamma radiation detector module according to claim 1, wherein the array of photosensors in each scintillation crystal comprises at least one two-dimensional array of individual photodiodes.

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

12. 6. A gamma radiation detector module as described in any one of claims 1 to 5, wherein the array of photodiodes of each scintillation crystal includes a plurality of separate photosensors for providing separate photosensor measurements corresponding to separate depth ranges between the proximal end face and the distal end face of each scintillation crystal.

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

14. 14. The gamma radiation detector module of claim 13, wherein each photosensor includes a silicon photomultiplier (SiPM), and each scintillation crystal is associated with a plurality of silicon photomultipliers extending along the first side thereof.

15. further comprising processing circuitry; the processing circuitry a plurality of energy circuits, each configured to generate an energy signal as a sum of energies detected by one or more silicon photomultiplier tubes; and a plurality of timing signal circuits, each timing signal circuit configured to generate a different timing signal for scintillation photons detected by the one or more silicon photomultiplier tubes.

16. the processing circuit is a dual-channel high-speed circuit, each channel including circuitry for digitally encoding one or more of the timing signals and the energy signals of the silicon photomultiplier tube; 16. The gamma radiation detector module of claim 15, wherein each scintillation crystal in the array of scintillation crystals uses different high speed circuitry than adjacent scintillation crystals.

17. 1. A positron emission tomography (PET) scanning system comprising a plurality of gamma radiation detector modules arranged to form a detector ring, Each detector module is an array of elongated scintillation crystals, each elongated scintillation crystal including a proximal end face, two axially oriented side faces, two cross-axially oriented side faces, and a distal end face oriented in a radial direction of the detector ring for receiving gamma photons; an array of photosensors disposed along a first one of the axially facing sides of each elongated scintillation crystal for detecting scintillation photons; a reflective material disposed on a second one of the proximal end face, the distal end face, the radially oriented side face, and the axially oriented side face of each elongated scintillation crystal for internally reflecting scintillation photons; a cooling system for cooling the detector module; an image processing system coupled to the detector module and configured to generate an image based on electronic output from the detector module.

18. an intermediate substrate connected to the array of photosensors along a side of each elongated scintillation crystal of each detector module facing in the first axis direction; 18. The PET scanning system of claim 17, wherein the intermediate substrate is operable to transmit electrical signals from the array of photosensors of each elongated scintillation crystal toward a proximal end face of each elongated scintillation crystal.

19. 20. The PET scanning system of claim 18, wherein the array of elongated scintillation crystals in each detector module comprises a two-dimensional array of elongated scintillation crystals, the two-dimensional array of elongated scintillation crystals in each detector module comprising M elongated scintillation crystals in an axial direction of the detector ring and N elongated scintillation crystals in a cross-axial direction of the detector ring, where M and N are each integer values.

20. the elongated scintillation crystals of each detector module are arranged adjacent to each other without any gaps in the axial direction and the cross-axial direction of the detector ring; a given elongated scintillation crystal and an axially adjacent elongated scintillation crystal are axially separated from one another by (i) the array of photosensors on a first axially facing side of the given elongated scintillation crystal, (ii) the intermediate substrate of the given elongated scintillation crystal, and (iii) the reflective material on the second axially facing side of the axially adjacent elongated scintillation crystal; 20. The PET scanning system of claim 19, wherein the given elongated scintillation crystal and an adjacent elongated scintillation crystal in the cross-axis direction are separated from each other in the cross-axis direction by the reflective material.

21. an array of elongated scintillation crystals disposed on a detector ring, each elongated scintillation crystal including a proximal end face, two axially oriented side faces, two cross-axially oriented side faces, and a distal end face oriented radially into the detector ring to receive gamma photons; an array of photosensors having detection surfaces for detecting scintillation photons, the detection surfaces of the array of photosensors being disposed along the length of a first surface of an axially oriented side of each elongated scintillation crystal; an intermediate substrate connected to an output face of the array of photosensors along the length of a side of each elongated scintillation crystal facing the first axis, the intermediate substrate operable to transmit an electrical signal from the output face of the array of photosensors of each elongated scintillation crystal toward the proximal end face of each elongated scintillation crystal; 1. A gamma radiation detector module comprising:

22. the array of photosensors of each elongated scintillation crystal comprises a two-dimensional array of single photon avalanche diodes (SPADs); 22. The gamma radiation detector module of claim 21, wherein the intermediate substrate includes bias circuitry for operating the two-dimensional array of SPADs of respective elongated scintillation crystals in Geiger mode.

23. 23. A gamma radiation detector module as described in claim 21 or 22, comprising a reflective material disposed on the proximal end face, the distal end face, the two cross-axially oriented side faces, and a second one of the axially oriented side faces of each elongated scintillation crystal, the reflective material operating to internally reflect scintillation photons generated within each elongated scintillation crystal.

24. the array of elongated scintillation crystals includes a one-dimensional array of elongated scintillation crystals arranged along an axial direction of the detector ring; 24. The gamma radiation detector module of claim 23, wherein the elongated scintillation crystals are closely spaced in the axial direction of the detector ring, and a given elongated scintillation crystal and an axially adjacent elongated scintillation crystal are separated from each other by (i) an array of photosensors on a side of the given elongated scintillation crystal facing in the first axial direction, (ii) the intermediate substrate of the given elongated scintillation crystal, and (iii) the reflective material on a side of the axially adjacent elongated scintillation crystal facing in the second axial direction.

25. 24. The gamma radiation detector module of claim 23, wherein the array of elongated scintillation crystals comprises a two-dimensional array of elongated scintillation crystals, the two-dimensional array of elongated scintillation crystals comprising M elongated scintillation crystals in an axial direction of the detector ring and N elongated scintillation crystals in a cross-axial direction of the detector ring, where M and N are each integer values.

26. The elongated scintillation crystals are arranged adjacent to each other without any gaps in the axial direction and the cross-axial direction of the detector ring, a given elongated scintillation crystal and an axially adjacent elongated scintillation crystal are axially separated from one another by (i) an array of photosensors on a first axially oriented side of the given elongated scintillation crystal, (ii) the intermediate substrate of the given elongated scintillation crystal, and (iii) the reflective material on the second axially oriented side of the axially adjacent elongated scintillation crystal; 26. The gamma radiation detector module of claim 25, wherein the given elongated scintillation crystal and an adjacent elongated scintillation crystal in the cross-axis direction are separated from each other in the cross-axis direction by the reflective material.

27. 1. A circuit for processing nuclear scintillation events, comprising: a plurality of photosensors that detect scintillation photons generated by the scintillation phenomenon in the scintillation crystal and output an electrical signal; an energy circuit for generating a digital signal representative of the total energy detected by one or more photosensors of the plurality of photosensors during a detection period; a timing circuit for generating a digital signal having different timing signals for the scintillation photons detected by the one or more photosensors during the detection period.

28. 28. The circuit of claim 27, wherein the photosensor includes a plurality of single photon avalanche diodes (SPADs) biased to operate in Geiger mode, the bias circuit being integrated into an intermediate substrate, and each of the plurality of photosensors being disposed on a side of the scintillation crystal.

29. The energy circuit an analog-to-digital converter; a multiplexer that selects the output signals of each of the plurality of photosensors for conversion by the analog-to-digital converter; a signal delay line connecting one or more of the plurality of photosensors to the multiplexer; and digital control logic that controls the multiplexer selection between the output signals of the plurality of photosensors.

30. the timing circuit an analog-to-digital converter for high-speed digital encoding of the capacitively isolated output of one or more photosensors of the plurality of photosensors; a transmission line connecting the capacitively isolated output of one or more photosensors among the plurality of photosensors to the analog-to-digital converter via an output diode; a flip-flop that selectively triggers encoding of the capacitively isolated output of one or more of the plurality of photosensors by the analog-to-digital converter by biasing the output diode to an on state; a comparator that compares the capacitively isolated output of one or more photosensors of the plurality of photosensors to one of a low voltage threshold and a low current threshold, the output of the comparator selectively triggering the flip-flop.

31. the output diode includes a PIN diode; The PIN diode is 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.

32. The energy circuit a first AD converter; a multiplexer that selects output signals from each of the plurality of photosensors for conversion by the first analog-to-digital converter; a signal delay line connecting each photosensor to the multiplexer; digital control logic for controlling the multiplexer selection between the output signals of the plurality of photosensors; The timing circuit a second analog-to-digital converter for high-speed digital encoding of the capacitively isolated output of one or more photosensors of the plurality of photosensors; a transmission line connecting a capacitively isolated output derived from one or more of the plurality of photosensors to the AD converter via an output diode; a flip-flop that selectively triggers encoding of the capacitively isolated output derived from one or more of the plurality of photosensors by the second analog-to-digital converter by biasing the output diode to an on state; a comparator that compares an output signal of one or more photosensors of the plurality of photosensors to a threshold, the output of the comparator selectively triggering the flip-flop to bias the output diode to an on state.

33. 33. The circuit of claim 32, wherein the comparator comprises a voltage comparator that compares a voltage level of an output signal of one or more photosensors of the plurality of photosensors to a voltage threshold.

34. 33. The circuit of claim 32, wherein the comparator comprises a current comparator that compares a current level of an output signal of one or more photosensors of the plurality of photosensors to a current threshold.

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

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

37. 33. 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 milliamps.

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

39. 33. The circuit of claim 32, wherein the output diode comprises a PIN diode 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 milliamps, and a forward bias on-resistance of less than 10 ohms at an operating frequency of 1 gigahertz.

Images