High-resolution depth-encoded PET detector with pseudo-prismatic optical guide array

Abstract

A particle detection device and a method for manufacturing the same are provided. The particle detection device includes a scintillator array including a plurality of scintillator crystals, a plurality of detectors provided at a bottom end of the scintillator array, and a plurality of pseudoprisms provided at a top end of the scintillator array. The pseudoprisms are configured to redirect particles between the top ends of the crystals of the scintillator array. The bottom ends of a first group of crystals of the scintillator array are configured to direct particles to a first detector of the plurality of detectors, and the bottom ends of a second group of crystals of the scintillator array are configured to direct particles to a second detector substantially adjacent to the first detector.

Description

[Technical Field]

[0001] Priority This application claims the benefits of U.S. Provisional Patent Applications No. 62 / 806,035 and No. 62 / 915,676, filed with the U.S. Patent and Trademark Office on 15 February 2019 and 16 October 2019, respectively. The contents of each provisional application are incorporated herein by reference in their entirety.

[0002] government support This invention was made with the support of the U.S. Government under Grant No. EB024849, awarded by the National Institutes of Health. The U.S. Government has certain rights with respect to this invention.

[0003] This invention generally relates to the field of radiographic imaging, and more specifically to positron emission tomography (PET). [Background technology]

[0004] Molecular imaging using PET is a powerful technique primarily used for the diagnosis, treatment selection, treatment monitoring, and research of cancer[1] and neuropsychiatric disorders[2]. Despite its high molecular specificity, quantitative nature, and clinical applicability, PET has been unable to achieve its full potential as a go-to molecular imaging modality, mainly due to its relatively poor spatial resolution, currently 3–6 mm[3,4]. This type of spatial resolution makes it impossible to measure target densities in nodules related to etiology and pathophysiology, as well as in many human and rodent brain regions.

[0005] To mitigate parallax errors (line of response mispositioning) for long scintillator crystals, depth-encoding PET detector modules have been developed [5]. These enable small-diameter PET rings with lower component costs per detector ring, large solid angle coverage for increased sensitivity, and a low contribution of annihilation gamma-ray noncollinearity to spatial resolution when using crystals with small cross-sectional areas [4, 6]. Furthermore, depth-of-interaction (DOI) information can be used to perform deconvolution of optical photon transport in long crystals, thus improving temporal resolution [7, 8]. Depth-encoding detectors based on dual-ended readout achieve the best continuous DOI resolution of <2 mm [9, 10]. High-resolution PET systems such as Clear-PEM, specifically for mammography, have been developed using dual-ended DOI readout detectors

[11] . However, these systems are too expensive to commercialize due to the larger number of readout electronic components compared to standard single-ended readout PET scanners. Recently developed high-resolution versions of these detectors exhibit relatively inferior energy and temporal resolution because glass light guides are used at the crystal-readout interface to achieve accurate crystal identification

[12] . To obtain DOI information, alternative single-ended readout detector modules have been proposed, such as multi-layer phoswich blocks [13, 14], retroreflectors for modules with monolithic scintillators

[15] , and other custom reflector designs [16, 17]. However, in all these designs, there are trade-offs between depth coding, cost, scintillator-readout coupling ratio, accuracy of crystal identification, and energy and temporal resolution.An ideal depth coding detector module to mitigate these trade-offs is a single-ended readout depth coding detector module in which the crystal array is directly coupled to silicon photomultiplier (SiPM) pixels without intermediate glass optical guides, minimizing the sharing of scintillation photons moving downward across multiple pixels and maintaining good temporal resolution. Furthermore, to maintain good energy and DOI resolution and mimic the behavior of a dual-ended depth coding readout detector, upward-moving photons that do not contribute to temporal information should have their direction changed towards the nearest SiPM by bending their photon path by 180°.

[0006] In response to this, recent studies have explored detector modules consisting of a depolished polycrystalline scintillator array, with one end coupled to SiPM pixels in a 4:1 ratio and the other end coupled to a uniform glass light guide, in order to develop a practical and cost-effective high-resolution time-of-flight (TOF) PET scanner and to achieve continuous DOI localization using single-ended readout [8, 18, 19]. See Frazao et al., U.S. Patent No. 10,203,419, which is incorporated herein by reference. These detector modules each measure 1.53 × 1.53 × 15 mm 3 Crystals and 3×3mm 2Energy-weighted averaging is used for crystal identification to achieve 9% energy and DOI resolution and 3 mm full width at half maximum (FWHM) using SiPM pixels [8]. However, these arrays have poor crystal identification along their edges and corners due to the lack of light-sharing neighbors

[19] , and this is a challenge that must be addressed, as edge and corner pixels account for 75% and 44% of 4×4 and 8×8 SiPM readout chips, respectively. Furthermore, intercrystal light sharing is inefficient when using a uniform glass light guide. This is because many photons traveling upward are reflected back to the primary pillar, and the remaining photons are isotropically shared between neighboring crystals with a Gaussian intensity distribution. The problem with isotropic light sharing is that low-intensity signals are distributed across many SiPMs, and their integrity is severely affected by dark counts, resulting in degraded energy and DOI resolution. [Overview of the project]

[0007] To overcome the shortcomings of conventional systems, this specification provides a particle detector based on a prismatoid PET (Prism-PET) detector module and a method for operating the particle detector.

[0008] Accordingly, aspects of the present invention solve the above problems and drawbacks and provide the advantages described below. One aspect of the present invention provides a particle detection device comprising a scintillator array containing a plurality of scintillator crystals, a plurality of detectors provided at the bottom of the scintillator array, and a plurality of pseudoprisms provided at the top of the scintillator array. Each of the pseudoprisms is configured to change the orientation of a particle between the top ends of the crystals of the scintillator array. The bottom end of a first group of crystals of the scintillator array is configured to guide a particle to a first detector among the plurality of detectors, and the bottom end of a second group of crystals of the scintillator array is configured to guide a particle to a second detector substantially adjacent to the first detector.

[0009] One aspect of the present disclosure provides a particle detector comprising a scintillator array including a plurality of scintillator crystals, a plurality of detectors provided at the bottom of the scintillator array, a plurality of pseudoprisms provided at the top of the scintillator array, and at least one processor that operably communicates with the plurality of detectors. The at least one processor includes a plurality of supervised machine learning algorithms configured to perform 3D gamma-ray localization of at least one interaction site in at least one of the plurality of scintillator crystals.

[0010] The above-described aspects, features, and advantages of certain embodiments of the present invention, as well as other aspects, features, and advantages, will become more apparent from the following detailed description in conjunction with the accompanying drawings. [Brief explanation of the drawing]

[0011] [Figure 1] This figure shows the distribution of light using a uniform glass light guide module and a uniform glass light guide module. [Figure 2] This figure shows a uniform distribution due to light sharing within a conventional optical guide. [Figure 3] This figure shows a Prism-PET module and the distribution of light using the Prism-PET module according to embodiments of the present disclosure. [Figure 4] This figure shows a pseudo-angular columnar optical guide array based on the embodiment of this disclosure. [Figure 5] This figure shows the arrangement of a 4:1 coupled Prism-PET module according to an embodiment of the present disclosure. [Figure 6] This is a perspective view providing details of a pseudo-prism array based on embodiments of the present disclosure. [Figure 7] This figure shows the arrangement of a 9:1 coupled Prism-PET module according to an embodiment of the present disclosure. [Figure 8]A perspective view of an optical guide array according to an embodiment of the present disclosure. [Figure 9] A diagram showing a 4-to-1 detector readout with uniform glass. [Figure 10] A diagram showing the detector readout of a 4-to-1 coupled Prism-PET module according to an embodiment of the present disclosure. [Figure 11] A diagram showing the detector readout of a 9-to-1 coupled Prism-PET module according to an embodiment of the present disclosure. [Figure 12] A diagram showing a crystal discrimination histogram based on centroiding and measurement energy histograms with and without DOI filter processing according to an embodiment of the present disclosure. [Figure 13] A diagram showing the DOI resolution of a uniform glass optical guide module and a 4-to-1 coupled Prism-PET module according to an embodiment of the present disclosure. [Figure 14(a)-(d)] A graph of the measured DOI resolution of a 4-to-1 coupled Prism-PET module according to an embodiment of the present disclosure. [Figure 15(a)-(d)] A diagram showing the sensitivity graphs and dimensions of several different PET scanners according to an embodiment of the present disclosure. [Figure 16(a)-(f)] A diagram showing the theoretically optical distribution of side-by-side Compton interactions within a 4-to-1 coupled Prism-PET module according to an embodiment of the present disclosure. [Figure 17] A diagram showing the photoelectric and Compton interaction measurements within a 4-to-1 coupled Prism-PET module according to an embodiment of the present disclosure.

Mode for Carrying Out the Invention

[0012] A detailed description of certain embodiments of the present invention is made with reference to the accompanying drawings. To avoid obscuring the present invention with unnecessary details, descriptions of related functions or structures known in the art are omitted when describing the present invention for clarity in understanding the concepts of the present invention.

[0013] This specification discloses a single-ended readout depth coding detector module that utilizes a dedicated pattern of segmented pseudoprismatic optical guides. Among the features of the Prism-PET detector modules disclosed in various embodiments, at least three distinct pseudoprismatic designs are utilized, namely center pseudoprismatics, edge pseudoprismatics, and corner pseudoprismatics, each having a predetermined different design to mitigate edge and corner artifacts and thus achieve uniform crystal recognition performance.

[0014] To generate deterministic anisotropic intercrystal light sharing patterns and maximize the signal-to-background ratio on those SiPMs, thereby improving both energy resolution and DOI resolution, intercrystal light sharing is limited to crystals belonging to the nearest SiPM.

[0015] This segmentation pattern improves crystal identification by separating adjacent crystals that would otherwise have similar readout patterns. The shape of each pseudoprism is interchangeable, and embodiments of the pseudoprism are substantially formed as at least one of the following: at least one prism, at least one antiprism, at least one frustum, at least one traingle, at least one cupola, at least one parallelepiped, at least one wedge, at least one pyramid, at least one truncated pyramid, at least one portion of a sphere, at least one cuboid, and at least one pyramid. For ease of reference, the right triangular prism is discussed herein. The right triangular prism improves the intercrystal light sharing ratio and therefore improves both crystal identification and DOI resolution.

[0016] As shown in Figures 1-8, when an optical photon is incident on the hypotenuse of a right-angled triangular prism, the optical photon experiences a 180° deflection, and due to the crystal-prismatic coupling scheme offset from the crystal-pixel coupling, the optical photon is efficiently guided to an adjacent crystal coupled to a different readout pixel.

[0017] Figure 1 shows the light distribution using a uniform glass optical guide module and a uniform glass optical guide module simulated with TracePro. Figure 2 shows the uniform distribution due to light sharing within a conventional optical guide.

[0018] Figure 3 shows the light distribution using a uniform glass Prism-PET and a uniform glass Prism-PET simulated with TracePro. Figure 4 shows a pseudo-prismatic optical guide array of an embodiment of the present disclosure. Figure 4 shows that for the Prism-PET of an embodiment of the present disclosure, light sharing is limited to an array of 16 × 16 crystals 120 coupled to the same pseudo-prismatic column, thereby improving the inter-crystal light sharing ratio.

[0019] Figure 5 shows the arrangement of a 4:1 coupled Prism-PET module according to an embodiment of the present disclosure. The lower left corner of Figure 5 is a plan view showing the relative arrangement of the 2x2 crystals of the scintillator array, which includes multiple scintillator crystals, and the SiPM pixels 140 of the multiple detectors provided at the bottom of the scintillator array. The upper right corner of Figure 5 shows three distinct pseudoprism designs utilized in the embodiment of Figure 5, the center pseudoprism 162, the edge pseudoprism 168, and the corner pseudoprism 166, which have different designs to mitigate edge and corner artifacts and thus achieve uniform crystal recognition performance. As shown in Figure 6, these three distinct pseudoprisms are provided at the top 122 of the scintillator array in a predetermined arrangement and are configured to change the orientation of particles between the top vertices of the crystals of the scintillator array.

[0020] For comparison, different Prism-PET detector modules were fabricated. The first Prism-PET detector 142 was similar to the module previously considered in references [8, 20], with one side coupled 4:1 to an 8×8 SiPM readout array 140 and the other side (the radiation receiving side) coupled to a uniform glass optical guide, measuring 1.4 × 1.4 × 20 mm. 3 The first Prism-PET detector consists of a 16×16 array of lutetium yttrium orthosilicate (LYSO) crystals. The second Prism-PET detector 144 consists of the same crystal and readout geometry, but to optimize the light sharing pattern, the conventional single uniform glass light guide has been replaced with a pseudo-prismatic light guide array that has a unique design and layout of prismatics at the corners, edges, and center of the detector module (Figures 5-8). The third Prism-PET detector has a pseudo-prismatic light guide array and uses the same SiPM array as the other detectors, but to achieve a 9:1 coupling, it uses a 0.96×0.96×20mm array. 3An array of approximately 24 × 24 LYSO crystals was used (Figures 7 and 8). In both Prism-PET detector modules, the scintillator crystals were coupled in equal proportions to the readout pixels and right-angled triangular prisms.

[0021] The coupling scheme of the prisms is offset from the coupling scheme of the readout pixels so that each crystal is coupled only to other crystals belonging to different readout pixels (Figure 5). When optical photons are incident on the pseudoprisms after gamma-ray interaction within the crystals, the orientation of the photons (i.e., particle 300) is efficiently redirected to adjacent crystals due to the right-angled triangular prism geometry, which improves the light sharing ratio between pixels (Figure 2). The geometry of each pseudoprism is position-dependent, and to optimize crystal separation, this geometry is predetermined to separate adjacent crystals along the edges and corners that would otherwise have similar readout patterns. In certain embodiments of this disclosure, a first group comprises four crystals, a second group comprises four crystals, and the first and second groups share two adjacent crystals among their four crystals. In certain embodiments, only the shared crystals are configured to guide particles to both the first and second detectors.

[0022] As shown in Figures 7 and 8, a first pseudoprism 162 of a plurality of pseudoprisms is configured to redirect particles between the apex ends 122 of a group of nine crystals in the scintillator array 120. In embodiments of the present disclosure, a center crystal 139 of a group of nine crystals is configured to guide particles to four adjacent detectors 142, 144, 146, and 148, a second pseudoprism 164 of the plurality of pseudoprisms is configured to redirect particles between the apex ends of another group of nine crystals in the scintillator array, the first pseudoprism 162 is substantially adjacent to the second pseudoprism 164, and the group of nine crystals 132 is substantially adjacent to another group of nine crystals 134. In embodiments of the present disclosure, a corner pseudoprism of the plurality of pseudoprisms is configured to redirect particles between the apex ends of a group of five crystals in the scintillator array. In embodiments of this disclosure, an edge pseudoprism 168 among a plurality of pseudoprisms is configured to change the orientation of a particle between the apex ends of a group of five crystals in a scintillator array.

[0023] This bonding scheme limits intercrystal light sharing to adjacent SiPMs, improving crystal identification. The refractive index n between the scintillator columns, prisms, and bonding adhesive can be matched to further improve light sharing and, consequently, enhance DOI resolution and crystal identification. All prisms were fabricated using SF10 glass with n=1.767 (instead of BK7 with n=1.53, which is the material for the uniform glass light guide) and bonded to the scintillator array using NOA170 adhesive with n=1.7. Barium sulfate (BaSO4) is used as a reflector material between the crystals and prisms due to its high spatial performance, which does not degrade energy or time resolution

[21] . In this case, the SiPM saturation effect, which is known to distort energy resolution in the positive direction and have a negative effect on DOI resolution, was not considered

[22] .

[0024] Figure 6 shows a perspective view providing details of a pseudo-prism array based on an embodiment of the present disclosure.

[0025] Figure 6 shows a perspective view of the pseudoprismatic array of the 4:1 coupled Prism-PET module, cross-sectional views of the pseudoprismatics and their corresponding crystals, and individual figures of the corner, edge, and center pseudoprismatics. The bottom end of the first group of crystals in the scintillator array shown in Figure 6 is configured to guide particles to a first detector among multiple detectors, and the bottom end of the second group of crystals in the scintillator array is configured to guide particles to a second detector substantially adjacent to the first detector.

[0026] Figure 7 shows the arrangement of a 9:1 coupled Prism-PET module according to an embodiment of the present disclosure. The inset in Figure 7 shows a predetermined readout pattern for each crystal belonging to a single pseudoprismatic optical guide within the 9:1 coupled module.

[0027] Figure 8 shows a perspective view of the optical guide array, the pseudoprismatic crystal array, and a cross-sectional view of the pseudoprismatic crystal of a 9:1 coupled Prism-PET module according to an embodiment of the present disclosure.

[0028] Experimental measurements are used to demonstrate the advantages in terms of crystal identification, energy resolution, and DOI resolution. This includes how Prism-PET enables crystal-to-readout coupling of up to 9:1, which can be used to significantly improve spatial resolution without increasing the number of readout channels (Figures 7 and 8).

[0029] Figure 9 shows a 4:1 detector readout with uniform glass. Figure 10 shows a detector readout of a 4:1 coupled Prism-PET module according to an embodiment of the present disclosure. Figure 11 shows a detector readout of a 9:1 coupled Prism-PET module according to an embodiment of the present disclosure.

[0030] These detector modules consisted of LYSO crystal arrays manufactured in X-Lum (Shanghai, China), coupled (4:1 or 9:1) to an 8x8 SiPM array (Hamamatsu S13361-3050AE-08). Data acquisition was performed using PETsys Electronics' TOFPET2 application-specific integrated circuit (ASIC) and FEB / D v2 readout board. Flood data was acquired on 4:1 and 9:1 coupled detector modules with pseudo-prismatic light guide arrays by uniformly exposing the modules with a 3 MBq Na-22 sodium point source (effective diameter 5 mm). Flood histograms were created using 10,000,000 events from the 4:1 modules and 22,500,000 events from the 9:1 modules (to acquire an equal number of events per crystal).

[0031] Figure 12 shows the Gaussian histograms and filtered energy spectra of the 4:1 uniform glass in Figure 9, the 4:1 Prism-PET module in Figure 10, and the 9:1 coupled Prism-PET module in Figure 11. The upper half of Figure 12 shows 1D Gaussian histograms illustrating the x-direction crystal separation from the corner, edge, and center readout pixels of the modules in Figures 9-11. The lower half of Figure 12 shows the filtered energy spectra from the center crystal of Figures 9-11 with and without DOI correction (13%, 9%, and 10%).

[0032] DOI performance was measured experimentally on a per-crystal basis using a method similar to that described in reference

[18] . Using lead collimation, modules were exposed to a Na-22 source at five different crystal depths (2, 6, 10, 14, and 18 mm). The source was placed in a lead cylinder with a 1 mm pinhole. The pinhole was a single 1.4 × 1.4 × 20 mm on one side of the DOI-aligned module and on the other side of the reference module. 3The crystals were aligned. Using simultaneous events between these two modules, scattering events were eliminated, and only events along the intended line of response were accepted. For all crystals, a histogram of the DOI estimation parameter

[18] w was computed and plotted. The w histogram was then transformed into DOI space using linear regression to determine the slope between w and the ground truth DOI, which should be at the center of each Gaussian peak. The DOI resolution of the crystal was computed using the width of the Gaussian peak transformed into DOI space (Figure 14). The DOI resolution is depth-dependent and equal to the FWHM of the Gaussian histogram. The overall DOI resolution for a crystal was computed as the average of the DOI resolutions across the measured depths

[18] . The DOI resolution of each module was computed using a typical center crystal for each module.

[0033] The spatial performance of the Prism-PET module of this disclosure is characterized by comparison with a standard uniform glass optical guide module using the flood histogram of a fabricated module (Figures 9-11). The glass optical guide module has the disadvantage of edge and corner effects, resulting in inferior position-dependent crystal separation. Prism-PET enables superior crystal separation without edge and corner artifacts across the entire detector array. This was not achieved with previous 4:1 coupled detector modules with single-ended TOF-DOI readout [8, 19, 20]. Similar results were shown with a 9:1 coupled Prism-PET module (Figure 11) at 3.2 × 3.2 mm. 2 The TOF-DOI PET detector module with SiPM pixels exhibits homogeneous crystal separation of less than 1 millimeter. Creating a 1D event positioning histogram (in the x-direction) confirms that the Prism-PET of this disclosure has uniform crystal separation performance at the center, edges, and corners. Furthermore, the Prism-PET achieves energy resolutions of 14% and 16% in 4:1 and 9:1 coupled modules with DOI correction, while uniform optical guiding only achieves 20% energy resolution (graph below, Figure 12).

[0034] Figure 13 shows the DOI of the modules of Figures 10 and 11 based on embodiments of the present disclosure. The histograms in Figures 13(a)–13(c) show the calculated DOI resolution at the interaction position in the depth direction of the center crystal for a 4:1 coupled detector module with uniform glass (Figure 13(a)) and a 4:1 coupled detector module with a pseudoprismatic optical guide (Figure 13(b)). Figure 13(c) shows a comparison of DOI resolution based on the optical guide used, which shows that the Prism-PET detector modules of the present disclosure achieve a twofold improvement in DOI resolution compared to the uniform glass optical guide, as experimentally measured for a single center crystal of each module. The measured DOI resolution of the glass optical guide was 5 mm in FWHM, showing strong agreement with previously reported results

[19] . The Prism-PET module achieved an FWHM DOI localization of 2.5 mm, which was the best resolution reported to date using single-ended readout. The increased depth dependence of the w parameter is due to 1) a controlled deterministic light-sharing pattern within the pseudoprism, 2) increased light transfer from the scintillator to the optical guide due to refractive index matching, and 3) an improved 180° deflection of the path of upward-moving optical photons due to the right-angled triangular prism geometry, all of which improve light-sharing between crystals coupled to the same pseudoprism. Using DOI information, both temporal and energy resolution can be improved, with temporal resolution being improved by performing deconvolution of depth-specific photon transport inside the scintillator, and energy resolution being improved by constructing depth-specific photoelectric peaks [8, 18]. Embodiments of this disclosure achieved energy resolutions of 9% and 10% in 4:1 and 9:1 coupled Prism-PET modules and 13% in a uniform optical guide after applying DOI-based corrections. The DOI and energy resolution values ​​improved and deteriorated slightly after SiPM saturation correction, respectively

[22] , and it should be noted that the reported values ​​actually represent the relative performance of the Prism-PET of this disclosure compared to a uniform optical guide module, rather than absolute performance.

[0035] Figures 14(a) to 14(d) show DOI resolution graphs based on embodiments of the present disclosure, where the conversion from a DOI-specific w histogram to a DOI histogram indicates the DOI resolution of a single crystal at each depth. Figure 14(a) shows histograms of the DOI estimation parameter w acquired at 2, 6, 10, 14, and 18 mm. Figure 14(b) shows the fit between w and DOI by linear regression. Figure 14(c) shows the DOI histogram generated by obtaining the w histogram from Figure 14(a) and multiplying it by the slope of the linear fit from Figure 14(b). Figure 14(d) shows the DOI resolution at each acquisition depth based on the width of the Gaussian curve in Figure 14(c).

[0036] Perhaps the most important parameter to consider when building a PET system is gamma-ray detection sensitivity, which is directly related to the signal-to-noise ratio (SNR) and therefore determines patient processing, delivered dose, and image quality. Monte Carlo simulations using very advanced software such as GATE are the most reliable method for modeling and calculating system-level sensitivity. However, relative improvements in sensitivity and comparisons between systems can be performed analytically by calculating sensitivity gain based on (a) geometric sensitivity and (b) coincidence time resolution (CTR) with respect to time-of-flight readout (TOF), in equation (1), the sensitivity gain is equal to the square of the SNR gain

[24] .

[0037]

number

[0038] In the above equation, D is the diameter of the imaged object, and Δx is the length of the reconstructed line segment along the line-off response, which is directly proportional to CTR(Δt) in equation (2).

number

[0039] An example of a dedicated brain PET scanner that can be built using the Prism-PET detector module would be a cylindrical ring with an axial length of 50 cm and a diameter of 25 cm.

[0040] Figures 15(a) to 15(d) show sensitivity graphs based on embodiments of the present disclosure. Figure 15(a) shows the dimensions and geometric ranges of an example of the Siemens Biograph Vision, Explorer Total-Body PET scanner, and Prism-PET brain scanner. Figure 15(b) shows the geometric sensitivity of a point source placed at the center of each scanner shown in Figure 15(a). Figure 15(c) shows the relative sensitivity gain as a function of simultaneous time resolution. Figure 15(d) shows the effective sensitivity gain calculated as the product of the geometric efficiency (shown in Figure 15(b)) and the TOF sensitivity gain (shown in Figure 15(c)).

[0041] Figure 15(a) shows the dimensions of a brain Prism-PET scanner according to embodiments of the present disclosure in comparison to the dimensions of exemplary whole-body (Siemens Biograph Visions) and total-body (Expolorer) PET scanners. Having a small ring diameter and a large axial field of view greatly improves geometric efficiency (Figure 15(b)), but in exchange for a significant increase in parallax error and partial volumetric effects, which can be mitigated by performing depth-direction interaction location (DOI) readout

[26] . As a result, only small-diameter scanners for specific organs should be constructed that have a detector module with DOI localization capability, such as the inventor's Prism-PET module.

[0042] The CTR for TOF readout can also be recovered by using DOI readout and performing deconvolution of the DOI dependency (i.e., the difference in optical photon path lengths) to the simultaneous timing [8]. Because the inventor's module has a better DOI resolution (2.5 mm vs. 3 mm), assuming the same reported CTR (about 150 ps) shown herein, which is a safe lower bound estimate, Prism-PET enables a TOF sensitivity gain of nearly 10-fold based on equation (1) when imaging an object with a D of about 20 cm, such as a human brain (Figure 15(c)). The TOF sensitivity gain for imaging a human brain is slightly lower for Siemens Biograph Vision, which achieves a CTR of about 220 ps

[25] , and much lower for Explorer, which has a CTR > 400 ps

[23] (Figure 15(c)).

[0043] Figure 15(d) shows the overall effective sensitivity gain for imaging the human brain, considering both geometric efficiency and TOF sensitivity gain. Based on the above calculations, the Prism-PET scanner of the embodiments of this disclosure can improve sensitivity by 3-fold and 4-fold, respectively, compared to the Siemens Biograph Vision and Explorer scanners.

[0044] Figures 16(a) to 16(f) illustrate the Compton interaction according to embodiments of the present disclosure.

[0045] With respect to Compton interactions, the Prism-PET of this disclosure enables Compton scattering energy resolution (and therefore localization) due to its deterministic light-sharing pattern. Assume that the Prism-PET has a light guide and a 16x16 array of lutetium LYSO crystals coupled 4:1 to an 8x8 array of silicon photomultiplier tube (SiPM) pixels. Based on the approximation that each 511 keV gamma ray generates a signal on four different pixels through light sharing, the light-sharing ratio between all crystals belonging to the same pseudoprism can be measured directly from flood data using photoelectric events. Using this information, the energies of the primary interaction (i.e., recoil electrons) and secondary interaction sites (i.e., scattered gamma rays) are resolved. After obtaining the resolved energies, the two independently absorbed events within the scintillation block can be localized, and the scattering angles and DOIs can be determined. Due to the transition from random light sharing to a deterministic pattern in photoelectric events, side-by-side Compton scattering events can be identified with respect to the Prism-PET module of this disclosure (Figures 16 and 17).

[0046] Figures 16(a) to (c) show examples of Compton energy resolution in a polycrystalline scintillator array having Prism-PET according to the present disclosure. Figure 16(d) provides examples of the optical sharing fraction ratio between labeled pixel 1 and adjacent pixels in Figures 16(a) to (c). In one case, both pixels (2 and 4) are adjacent to pixel 1, resulting in equal optical sharing fractions, while in the other case, pixel 3 is diagonally opposite pixel 1, resulting in a smaller optical sharing fraction. (E), (F) Energy and DOI errors of Compton interaction resolution for Prism-PET.

[0047] The classical Compton energy decomposition can be performed as follows: The total absorbed energies EA and EB by components A and B (scattered and recoiled electrons) are given by equation (3) as the sum of the energies of all four SiPMs. [Number]

[0048] In the above equation, E A1 and E B1 are the maximum deposited energy in the SiPM coupled to the crystal pixel where the interaction occurred, and E A2,3,4 and E B2,3,4 are the deposited energy in the adjacent columns due to light leakage at the bottom (from the SiPM side) and light leakage at the top through the prism mirror light guide. The experimental results in Supplementary Figure 16(a) correspond to four known parameters E 1-4 shown by the energy detected by each of the four pixels after side-by-side Compton scattering events, and the total gamma particle energy accumulated is shown by Equation (4). E γ = E A + E B (4)

[0049] Note that the energies of the components of the Compton scattering event, namely E A1-4 and E B1-4 are unknown. Writing these equations based on the measured energy gives Equation (5). E1 = EA1 + EB4 E2 = EA2 + EB1 E3 = EA3 + EB2 E4 = EA4 + EB3 (5)

[0050] This gives four equations and eight unknowns. However, the deposited energy in adjacent columns is correlated. Consider the inset in Figure 16(d). The maximum deposited energy occurred in the top left SiPM. Assuming that the shared fraction with three adjacent crystals depends on the proximity of the adjacent crystals to the crystal where the interaction occurred, and using the Pythagorean theorem to form a right triangle with the centers of the three adjacent crystals as vertices, Equation (6) is obtained.

[0051]

number

[0052] In the above equation, for example, d 12 This is the distance between the centers of the primary SiPM1 and the adjacent SiPM2. Substituting equation 6 into equation 5 yields equation (7). E1 = EA1 + EB2 E2 = EA2 + EB1 E3 = 0.7EA2 + EB2 E4=E A2 +0.7E B2 (7)

[0053] This involves four equations and four unknowns. It should be noted that, in practice, due to some unavoidable small misalignment between the prismatic mirror light guide and the scintillator column, the shared fractions have a spatial deviation from the ideal case shown in Equation 6. However, as shown in Figure 16(d), they can be obtained empirically across the array by analyzing the shared fractions from individual photoelectric events obtained using flood histogram experiments. Figures 16(b) and (c) show two decomposed elements of measured side-by-side Compton scattering events based on the above analysis.

[0054] If the inventor's module has DOI localization, the DOI variable can be expressed as equation (8). wA = EA1 / EA w B =E B1 / E B (8)

[0055] As shown in Figures 15(e) to (f), {E} is based on 200,000 experimental gamma events. A1 ,E B1} and {w A ,w BThe inventor's estimate of the error percentage is approximately 10%. This error can be further reduced by using a convolutional neural network as the estimator. This is particularly useful because millions of gamma events can be collected as a training dataset using flood histogram experiments.

[0056] Figures 15(a)–(d) illustrate how a Compton event, in which recoil electrons and scattered gamma rays are completely absorbed by adjacent scintillators of two different SiPMs, can be decomposed into its constituent elements. Using classical Compton decomposition, the DOI variable w was calculated with a full width at half maximum (FWHM) error of 11% (Figure 16(e)). Furthermore, the FWHM energy error resulting from the Compton decomposition was 15% (Figure 16(f)).

[0057] Figure 17 illustrates the photoelectric and Compton interactions according to embodiments of the present disclosure, with a graph of the random light-sharing pattern of a glass light guide superimposed on a graph of the deterministic light-sharing pattern of the Prism-PET according to embodiments of the present disclosure. Figure 17 shows experimental results for several examples of Compton events absorbed in adjacent crystals of the "Prism-PET module of the present disclosure" versus the "module with a plate glass light guide".

[0058] The light sharing pattern within a glass optical guide module is random, which makes it difficult (and often impossible) to decompose the detected energy into the constituent energies of scattered photons and recoil electrons. In a Prism-PET module, due to its right-angled triangular prism geometry, the light sharing pattern is deterministic, which makes it practical to decompose events into their constituent energies based on known light sharing ratios between crystals.

[0059] Accordingly, a particle detector is provided, comprising a scintillator array containing multiple scintillator crystals, multiple detectors provided at the bottom of the scintillator array, multiple pseudoprisms provided at the top of the scintillator array, and at least one processor that operably communicates with the multiple detectors. This at least one processor includes multiple supervised machine learning algorithms, including convolutional and regression networks, configured to perform 3D gamma-ray localization of at least one interaction site in at least one scintillator crystal among the multiple scintillator crystals. This at least one processor is configured to recover at least one Compton event scattered among the multiple scintillator crystals and to localize that at least one Compton event at the scintillator level based on 3D gamma-ray localization. This at least one processor is further configured to determine the scattering angle based on the at least one Compton event and DOI information. This at least one processor is further configured to localize at least one Compton event based on the resolved energies of at least two interactions absorbed within multiple scintillator crystals, the resolved energies based on at least one light-sharing pattern, and the at least one light-sharing pattern based on the positions of the multiple scintillator crystals relative to multiple detectors and multiple pseudoprisms.

[0060] According to embodiments of the present disclosure, at least one optical sharing pattern is mapped based on the optical sharing ratio between the same pseudoprism scintillator crystals, the optical sharing ratio is based on a predetermined geometry of at least one pseudoprism among a plurality of pseudoprisms, this mapping is based on a measured photoelectric event, the resolved energy of at least one primary interaction and at least one secondary interaction, the at least one primary interaction being based on electron recoil and the at least one secondary interaction being based on gamma-ray scattering, and the optical sharing pattern is deterministic. [Industrial applicability]

[0061] Therefore, a cost-effective and practical method is provided for achieving high spatial and DOI resolution in polycrystalline single-ended readout detector modules without introducing edge and corner artifacts. Using embodiments of the present disclosure, depth coding in whole-body and total-body clinical PET scanners

[23] can be enabled without increasing costs (the cost of the pseudo-prismatic optical guide array accounts for less than 10% of the total cost of each Prism-PET module) and without increasing power consumption, and simultaneously (e.g., 2 × 2 × 20 mm) 3 Crystals 6 x 6 mm 2By coupling 9:1 to the readout pixels, spatial resolution can be improved; by intercrystal Compton scatter recovery, sensitivity can be improved; and by DOI correction for timing jitter, temporal resolution can be improved. For small-ring diameter brain imaging, a 9:1 coupling ratio enables spatial resolution of less than 1 millimeter, and expanding the axial field of view to approximately twice that of a whole-body PET scanner enables the same geometric sensitivity gain as the Explorer total-body PET scanner (Figure 15) [8, 23-26]. Furthermore, having a DOI resolution of 2.5 mm significantly reduces parallax errors and potentially allows for the achievement of a simultaneous temporal resolution of approximately 100 ps with DOI correction [8]. This would enable even higher sensitivity and spatial resolution [24-26]. These advantages provide a practical, cost-effective, and power-efficient method that achieves both high spatial resolution and high sensitivity at relatively low doses for quantitative in vivo function and molecular imaging of many organs of the human body, including important brain structures such as the raphe nuclei, cholinergic basal forebrain ganglia, locus coeruleus, and hypothalamic nuclei, which were not resolvable by existing PET scanners, all of which are thought to play critically important roles in basic physiology as well as in the pathophysiology of common neurodegenerative and psychiatric disorders [26-30]. The ability to visualize and quantify these targets and similar targets has the potential to revolutionize molecular imaging in both clinical and research fields, providing a tool that is currently unavailable for early diagnosis and basic research in oncology and brain disorders.

[0062] Another advantage of the embodiments of this disclosure is the ability to more accurately identify the initial interaction site of Compton scattering events, which further improves spatial resolution and sensitivity (Figures 16-17). Traditionally, Compton detection has been performed using multiple detector layers, but a recent paper outlined criteria for localizing and resolving Compton interactions using single-ended readout, citing high-resolution DOI readout as a key feature for Compton scattering recovery

[31] . Uniform optical guides are not optimal for this task because the SiPM patterns of individual events are random. In contrast, the inventors' Prism-PET module generates a deterministic optical sharing pattern regardless of the interaction location inside the primary scintillator column (Figures 1-4 and 14). In particular, Prism-PET enables the resolution of side-by-side scattered photon and recoil electron events, which are the most likely to occur and the most difficult to analyze, into their constituent energies, spatial locations, and DOIs. As crystal size decreases, scattered photons are more likely to be absorbed by crystals other than the primary interaction site, so Compton scattering recovery is particularly important for maintaining high sensitivity in detector modules with small scintillator crystals

[32] .

[0063] Embodiments of this disclosure provide a Prism-PET detector module that is a true single-ended analogue of a dual-ended depth-coded readout, improving light sharing using an efficient reflector that bends light by 180°. A 2.5 mm FWHM DOI resolution is achieved, and up to 9:1 scintillator-SiPM coupling is achieved for high spatial resolution, while simultaneously directly coupling the crystal array to the SiPM pixels to minimize light leakage and maintain the high photon detection efficiency required for good temporal resolution. The top-side reflector consists of an optimized pattern of segmented pseudo-prismatic light guides, which efficiently redirects the direction of the scintillation photon path from the primary crystal to the selected nearest SiPM, thus very closely mimicking the operation of a dual-ended readout detector. This produces an anisotropic and deterministic signal pattern that can be used to decompose side-by-side Compton scattering events into their constituent energies and DOI information for scattering recovery purposes. Therefore, high and uniform spatial resolution is achieved (9:1 bonding of approximately 1 mm crystals; no edge and corner artifacts due to improved light sharing; spatial blur due to Compton scattered photons is reduced by scattering recovery), high sensitivity is achieved (20 mm thick detector and intercrystal Compton scattering recovery), and good energy and temporal resolution is achieved in a compact system (especially after applying DOI correction) (DOI coding eliminates parallax errors and allows for smaller ring diameters). This unique combination of features enabled the development of a cost-effective and compact TOF-DOI-Compton PET scanner based on the Prism-PET module for organ-specific functional and molecular imaging of small animals and humans.

[0064] While the present invention has been shown and described with reference to certain aspects of the invention, those skilled in the art will understand that various modifications of form and detail can be made to those aspects without departing from the spirit and scope of the invention as defined by the appended claims and their equivalents. The description of the claims set forth below shall not be construed as means plus functional elements without the express use of “means to” or “steps to.”

Claims

1. a scintillator array including a plurality of scintillator crystals; a plurality of detectors provided at a bottom end of the scintillator array; a plurality of pseudoprisms provided at the apex of the scintillator array; Equipped with each pseudoprism of the plurality of pseudoprisms configured to redirect particles between apexes of scintillator crystals of the scintillator array; a bottom end of a first group of scintillator crystals of the scintillator array configured to direct particles to a first detector of the plurality of detectors; a bottom end of a second group of scintillator crystals of the scintillator array configured to direct particles to a second detector substantially adjacent to the first detector; Particle detection equipment.

2. 2. The apparatus of claim 1, wherein each pseudoprism is substantially shaped as at least one of: at least one prism, at least one antiprism, at least one frustum, at least one triangle, at least one cupola, at least one parallelepiped, at least one wedge, at least one pyramid, at least one truncated pyramid, and at least a portion of a sphere.

3. The apparatus of claim 1 , wherein the first group includes four crystals and the second group includes four crystals.

4. The apparatus of claim 3 , wherein the first group and the second group share two adjacent crystals of the four crystals.

5. The apparatus of claim 4 , wherein the shared crystal is configured to direct particles to both the first detector and the second detector.

6. 2. The apparatus of claim 1, wherein a first pseudoprism of the plurality of pseudoprisms is configured to redirect particles between the apices of a group of nine crystals of the scintillator array.

7. 7. The apparatus of claim 6, wherein a center crystal of the group of nine crystals is configured to direct particles to four adjacent detectors.

8. 7. The apparatus of claim 6, wherein a second pseudoprism of the plurality of pseudoprisms is configured to redirect particles between the apices of another group of nine crystals of the scintillator array.

9. 9. The apparatus of claim 8, wherein the first pseudoprism is substantially adjacent to the second pseudoprism and the group of nine crystals is substantially adjacent to the other group of nine crystals.

10. The apparatus of claim 1 , wherein a corner pseudoprism of the plurality of pseudoprisms is configured to redirect particles between the apexes of a group of five crystals of the scintillator array.

11. The apparatus of claim 1 , wherein edge pseudoprisms of the plurality of pseudoprisms are configured to redirect particles between the apexes of a group of five crystals of the scintillator array.

12. a scintillator array including a plurality of scintillator crystals; a plurality of detectors provided at a bottom end of the scintillator array; a plurality of pseudoprisms provided on the apex of the scintillator array; at least one processor in operative communication with the plurality of detectors; Equipped with the at least one processor includes a plurality of supervised machine learning algorithms configured to perform three-dimensional (3D) gamma-ray localization of at least one interaction site within at least one scintillator crystal of the plurality of scintillator crystals; Particle detector.

13. 13. The detector of claim 12, wherein the at least one processor is further configured to recover at least one Compton event scattered between the plurality of scintillator crystals and localize the at least one Compton event at a scintillator level based on 3D gamma ray localization.

14. 13. The detector of claim 12, wherein the at least one processor is further configured to determine a scattering angle based on at least one Compton event and depth of interaction (DOI) information.

15. 13. The detector of claim 12, wherein the at least one processor is further configured to localize at least one Compton event based on resolved energies of at least two interactions absorbed in the plurality of scintillator crystals.

16. The detector of claim 15 , wherein the resolved energy is based on at least one light sharing pattern.

17. The detector of claim 15 , wherein the at least one light sharing pattern is based on the positions of the plurality of scintillator crystals relative to the plurality of detectors and the plurality of pseudoprisms.

18. 17. The detector of claim 16, wherein the at least one light-sharing pattern is mapped based on a light-sharing ratio between scintillator crystals of the same pseudoprismatic shape.

19. 20. The detector of claim 18, wherein the light sharing ratio is based on a predetermined geometry of at least one pseudoprism of the plurality of pseudoprisms.

20. the mapping is based on measured photoelectric events, resolved energies of at least one primary interaction, and at least one secondary interaction; the at least one primary interaction is based on electron recoil and the at least one secondary interaction is based on gamma ray scattering; 20. The detector of claim 18.

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