High resolution depth encoding pet detector having prismatoid light guide array

Abstract

SOLUTION: To provide a particle detector including a scintillator array including a plurality of scintillator crystals, a plurality of detectors provided in a bottom end of the scintillator array, and a plurality of prismatoids provided in a top end of the scintillator array. The prismatoid of the plurality of prismatoids is configured to change a direction of particles between the top ends of the crystals of the scintillator array. A first group of bottom ends of the crystals of the scintillator array are configured to guide the particles to the first detector of the plurality of detectors, and a second group of bottom ends of the crystals of the scintillator array are configured to guide the particles to the second detector substantially adjacent to the first detector.EFFECT: Provided is a practical, cost-effective and power-efficient method that achieves both of high spatial resolution and high sensitivity with a relatively low dose for a quantitative in vivo function and molecule imaging for a number of organs of a human body.SELECTED DRAWING: Figure 7

Description

[Technical Field]

[0001] Priority This application claims the benefit of U.S. Provisional Patent Applications Nos. 62 / 806,035 and 62 / 915,676, filed with the U.S. Patent and Trademark Office on February 15, 2019 and October 16, 2019, respectively. The contents of each provisional application are incorporated herein by reference in their entirety.

[0002] government support This invention was made with United States government support under Grant No. EB024849 awarded by the National Institutes of Health. The United States government has certain rights in this invention.

[0003] The present invention relates generally to the field of radiological imaging, and in particular 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 not achieved its full potential as a go-to molecular imaging modality, primarily due to its relatively poor spatial resolution, currently 3–6 mm [3, 4]. This type of spatial resolution does not allow for the measurement of target densities in nodules and many human and rodent brain regions relevant to pathogenesis and pathophysiology.

[0005] To mitigate parallax errors (mispositioning of the line of response) for long scintillator crystals, depth-encoding PET detector modules have been developed [5]. This allows for smaller diameter PET rings with lower component costs per detector ring, larger solid angle coverage for increased sensitivity, and a lower contribution of annihilation gamma-ray noncollinearity to spatial resolution when using crystals with small cross-sections [4, 6]. Furthermore, depth of interaction (DOI) information can be used to perform deconvolution of optical photon transport within 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 the Clear-PEM dedicated to mammography, have been developed using dual-ended DOI readout detectors

[11] . However, these systems require a larger number of readout electronics than standard single-ended readout PET scanners, making them too expensive for commercialization. Recently developed high-resolution versions of these detectors exhibit relatively poor energy and time resolution due to the use of glass light guides at the crystal-readout interface, which are necessary to achieve accurate crystal identification

[12] . Alternative single-ended readout detector modules have been proposed to obtain DOI information, 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 of these designs, trade-offs exist between depth encoding, cost, scintillator-readout coupling ratio, crystal identification accuracy, energy resolution, and time resolution.To mitigate these tradeoffs, an ideal depth-encoding detector module is one with single-ended readout, in which the crystal array is directly coupled to silicon photomultiplier (SiPM) pixels without an intermediate glass light guide, to minimize the sharing of downward-traveling scintillation photons across multiple pixels and maintain good time resolution. Furthermore, to maintain good energy and DOI resolution and mimic the behavior of a dual-ended depth-encoding readout detector, upward-traveling photons that do not contribute time information should change direction toward the nearest SiPM by bending the photon's path 180°.

[0006] Accordingly, recent studies to develop practical, cost-effective, high-resolution time-of-flight (TOF) PET scanners and to achieve continuous DOI localization using single-ended readout have considered detector modules consisting of depolished polycrystalline scintillator arrays coupled 4:1 to SiPM pixels on one side and to a uniform glass light guide on the other side [8, 18, 19]. See U.S. Patent No. 10,203,419 to Frazao et al., the contents of which are incorporated herein by reference. These detector modules each have a size of 1.53 × 1.53 × 15 mm. 3 crystals and 3 x 3 mm 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] . This is a challenge that must be overcome, as edge and corner pixels account for 75% and 44% of the 4x4 and 8x8 SiPM readout chips, respectively. Furthermore, inter-crystal light sharing is inefficient when using a uniform glass light guide. This is because many upward-traveling photons are reflected back into the first column, and the remaining photons are shared isotropically between neighboring crystals with a Gaussian intensity distribution. The problem with isotropic light sharing is the distribution of low-intensity signals across many SiPMs, the integrity of which is severely affected by dark counts, resulting in poor energy and DOI resolution. Summary of the Invention

[0007] To overcome the shortcomings of conventional systems, a particle detector and method of operating the same based on a prismatoid PET (Prism-PET) detector module is provided herein.

[0008]

[0006] Accordingly, aspects of the present invention address the above-mentioned problems and drawbacks and provide the advantages described below. One aspect of the present invention provides a particle detection device including 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 an apex end of the scintillator array. Each pseudoprism is configured to redirect particles between the apex 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.

[0009] One aspect of the present disclosure provides a particle detector including 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 an apex end of the scintillator array, and at least one processor in operative communication with the plurality of detectors, the at least one processor including a plurality of supervised machine learning algorithms configured to perform 3D gamma ray localization of at least one interaction site within at least one scintillator crystal of the plurality of scintillator crystals.

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

[0011] [Figure 1] 10A and 10B show the distribution of light using a uniform glass light guide module and a uniform glass light guide module. [Figure 2] FIG. 1 illustrates uniform distribution of light sharing in a conventional light guide. [Figure 3] FIG. 1 illustrates a Prism-PET module and the distribution of light using the Prism-PET module, according to an embodiment of the present disclosure. [Figure 4] FIG. 1 illustrates a pseudo-prismatic light guide array according to an embodiment of the present disclosure. [Figure 5] FIG. 1 illustrates the layout of a 4:1 coupled Prism-PET module according to an embodiment of the present disclosure. [Figure 6] FIG. 1 is a perspective view providing details of a pseudo-prismatic array according to an embodiment of the present disclosure. [Figure 7] FIG. 1 illustrates the layout of a 9:1 coupled Prism-PET module, according to an embodiment of the present disclosure. [Figure 8]FIG. 1 is a perspective view of a light guide array according to an embodiment of the present disclosure. [Figure 9] FIG. 10 shows a 4:1 detector readout with a uniform glass. [Figure 10] FIG. 1 illustrates detector readout of a 4:1 coupled Prism-PET module, according to an embodiment of the present disclosure. [Figure 11] FIG. 1 illustrates detector readout of a 9:1 coupled Prism-PET module, according to an embodiment of the present disclosure. [Figure 12] FIG. 10 shows a centroiding-based crystal identification histogram and a measured energy histogram with and without DOI filtering, according to an embodiment of the present disclosure. [Figure 13A] 10A-10C illustrate the DOI resolution of a solid glass light guide module and a 4:1 coupled Prism-PET module based on an embodiment of the present disclosure. [Figure 13B] 10A-10C illustrate the DOI resolution of a solid glass light guide module and a 4:1 coupled Prism-PET module based on an embodiment of the present disclosure. [Figure 13C] 10A-10C illustrate the DOI resolution of a solid glass light guide module and a 4:1 coupled Prism-PET module based on an embodiment of the present disclosure. [Figure 14(a)] 1 is a measured DOI resolution graph of a 4:1 coupled Prism-PET module, according to an embodiment of the present disclosure. [Figure 14(b)] 1 is a measured DOI resolution graph of a 4:1 coupled Prism-PET module, according to an embodiment of the present disclosure. [Figure 14(c)] 1 is a measured DOI resolution graph of a 4:1 coupled Prism-PET module, according to an embodiment of the present disclosure. [Figure 14(d)] 1 is a measured DOI resolution graph of a 4:1 coupled Prism-PET module, according to an embodiment of the present disclosure. [Figure 15(a)] FIG. 1 shows sensitivity graphs and dimensions for several different PET scanners, according to an embodiment of the present disclosure. [Figure 15(b)] FIG. 1 shows sensitivity graphs and dimensions for several different PET scanners, according to an embodiment of the present disclosure. [Figure 15(c)] FIG. 1 shows sensitivity graphs and dimensions for several different PET scanners, according to an embodiment of the present disclosure. [Figure 15(d)] FIG. 1 shows sensitivity graphs and dimensions for several different PET scanners, according to an embodiment of the present disclosure. [Figure 16(a)] FIG. 1 illustrates the theoretical light distribution of side-by-side Compton interactions within a 4:1 coupled Prism-PET module, according to an embodiment of the present disclosure. [Figure 16(b)] FIG. 1 illustrates the theoretical light distribution of side-by-side Compton interactions within a 4:1 coupled Prism-PET module, according to an embodiment of the present disclosure. [Figure 16(c)] FIG. 1 illustrates the theoretical light distribution of side-by-side Compton interactions within a 4:1 coupled Prism-PET module, according to an embodiment of the present disclosure. [Figure 16(d)] FIG. 1 illustrates the theoretical light distribution of side-by-side Compton interactions within a 4:1 coupled Prism-PET module, according to an embodiment of the present disclosure. [Figure 16(e)] FIG. 1 illustrates the theoretical light distribution of side-by-side Compton interactions within a 4:1 coupled Prism-PET module, according to an embodiment of the present disclosure. [Figure 16(f)] FIG. 1 illustrates the theoretical light distribution of side-by-side Compton interactions within a 4:1 coupled Prism-PET module, according to an embodiment of the present disclosure. [Figure 17]FIG. 1 illustrates photoelectric and Compton interaction measurements within a 4:1 coupled Prism-PET module according to an embodiment of the present disclosure. DETAILED DESCRIPTION OF THE INVENTION

[0012] The following detailed description of certain embodiments of the present invention will be made with reference to the accompanying drawings. In order to avoid obscuring the present invention with unnecessary details, when describing the present invention, descriptions of relevant functions or structures known in the art will be omitted for clarity in understanding the concept of the present invention.

[0013] Disclosed herein is a single-ended readout depth-encoding detector module that utilizes a dedicated pattern of segmented pseudoprismatic light guides. Among the features of the Prism-PET detector module disclosed in various embodiments, at least three distinct pseudoprismatic designs are utilized: center pseudoprismatic, edge pseudoprismatic, and corner pseudoprismatic, each with a different predetermined design for mitigating edge and corner artifacts and thus achieving uniform crystal discrimination performance.

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

[0015] This segmentation pattern improves crystal discrimination by separating adjacent crystals that would otherwise have similar readout patterns. The shape of each pseudoprism is interchangeable, and pseudoprism embodiments are 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, at least a portion of a sphere, at least one cuboid, and at least one pyramid. For ease of reference, right triangular prisms are discussed herein. Right triangular prisms improve inter-crystal light sharing ratios, thus improving both crystal discrimination and DOI resolution.

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

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

[0018] Figure 3 shows the light distribution using a uniform glass Prism-PET and the uniform glass Prism-PET simulated in TracePro. Figure 4 shows a pseudo-prismatic light guide array according to an embodiment of the present disclosure. For the Prism-PET according to an embodiment of the present disclosure, Figure 4 shows that light sharing is limited to an array of 16x16 crystals 120 bonded to the same pseudo-prismatic column, thereby improving the inter-crystal light sharing ratio.

[0019] FIG. 5 illustrates the layout of a 4:1 coupled Prism-PET module according to an embodiment of the present disclosure. The lower left corner of FIG. 5 is a plan view illustrating the relative arrangement of a 2×2 crystal of a scintillator array containing multiple scintillator crystals and multiple detector SiPM pixels 140 provided at the bottom end of the scintillator array. The upper right corner of FIG. 5 illustrates three distinct pseudoprism designs utilized in the embodiment of FIG. 5. Center pseudoprism 162, edge pseudoprism 168, and corner pseudoprism 166 have different designs for reducing edge and corner artifacts and thus achieving uniform crystal discrimination performance. As shown in FIG. 6, these three distinct pseudoprisms are provided at the apex 122 of the scintillator array in a predetermined arrangement and are configured to redirect particles between the apex ends of the crystals of the scintillator array.

[0020] For comparison, different Prism-PET detector modules were fabricated. The first Prism-PET detector 142 was a 1.4 × 1.4 × 20 mm detector, 4:1 coupled to an 8 × 8 SiPM readout array 140 on one side and a uniform glass light guide on the other side (the radiation receiving side), similar to modules previously discussed in the literature [8, 20]. 3 The second Prism-PET detector 144 consists of a 16 x 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 with a unique design and layout of prismatic pillars at the corners, edges, and center of the detector module (Figures 5-8). The third Prism-PET detector, with its pseudo-prismatic light guide array, uses the same SiPM array as the other detectors, but is sized 0.96 x 0.96 x 20 mm to achieve 9:1 coupling. 3An approximately 24x24 array of LYSO crystals was utilized (Figures 7 and 8). In both Prism-PET detector modules, the scintillator crystals are combined in equal proportions into readout pixels and right triangular prisms.

[0021] The coupling scheme of the prisms is offset from that of the readout pixels so that each crystal is coupled only to other crystals belonging to different readout pixels ( FIG. 5 ). When optical photons enter the pseudoprisms after gamma ray interactions within the crystals, the right-angled triangular prism geometry effectively redirects the photons (i.e., particles 300) to adjacent crystals, improving the light-sharing ratio between pixels ( FIG. 2 ). The geometry of each pseudoprism is position-dependent, and to optimize crystal separation, the geometry is predetermined to separate adjacent crystals along edges and corners that would otherwise have similar readout patterns. In certain embodiments of the present disclosure, a first group includes four crystals, a second group includes four crystals, and the first and second groups share adjacent two of the four crystals. In certain embodiments, only the shared crystals are configured to direct particles to both the first and second detectors.

[0022] 7 and 8 , a first pseudoprism 162 of the plurality of pseudoprisms is configured to redirect particles between the apices 122 of a group of nine crystals of the scintillator array 120. In an embodiment of the present disclosure, a center crystal 139 of the group of nine crystals is configured to direct particles to four adjacent detectors 142, 144, 146, 148, and a second pseudoprism 164 of the plurality of pseudoprisms is configured to redirect particles between the apices of another group of nine crystals of the scintillator array, with the first pseudoprism 162 substantially adjacent to the second pseudoprism 164 and the group of nine crystals 132 substantially adjacent to the other group of nine crystals 134. In an embodiment of the present disclosure, a corner pseudoprism of the plurality of pseudoprisms is configured to redirect particles between the apices of five crystals of the scintillator array. In an embodiment of the present disclosure, edge pseudoprisms 168 of the plurality of pseudoprisms are configured to redirect particles between the apexes of a group of five crystals in the scintillator array.

[0023] Because this coupling scheme limits inter-crystal light sharing to that between adjacent SiPMs, which improves crystal discrimination, the refractive indices n between the scintillator pillars, prisms, and bonding adhesive can be matched to further improve light sharing and, consequently, improve DOI resolution and crystal discrimination. All pillars were fabricated using SF10 glass with n = 1.767 (instead of BK7 with n = 1.53, the material used for the uniform glass light guide) and bonded to the scintillator array using NOA170 adhesive with n = 1.7. Barium sulfate (BaSO4) was used as the reflector material between the crystals and the pillars because of its high spatial performance without degrading energy or time resolution

[21] . The SiPM saturation effect, which is known to positively distort energy resolution and negatively impact DOI resolution

[22] , was not considered.

[0024] FIG. 6 shows a perspective view providing details of a pseudo-prismatic array according to an embodiment of the present disclosure.

[0025] A perspective view of the pseudoprism array of the 4:1 coupled Prism-PET module, a cross-sectional view of the pseudoprisms and their corresponding respective crystals, and individual views of the corner, edge, and center pseudoprisms are shown in Figure 6. The bottom ends of a first group of crystals in the scintillator array shown in Figure 6 are configured to direct particles to a first detector of the multiple detectors, and the bottom ends of a second group of crystals in the scintillator array are configured to direct particles to a second detector substantially adjacent to the first detector.

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

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

[0028] Experimental measurements are used to demonstrate advantages in terms of crystal identification, energy resolution, and DOI resolution, including how Prism-PET enables up to 9:1 crystal-to-readout coupling, 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 4:1 coupled Prism-PET module detector readout, and Figure 11 shows a 9:1 coupled Prism-PET module detector readout according to an embodiment of the present disclosure.

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

[0031] Figure 12 shows Gaussian histograms and filtered energy spectra for the 4:1 uniform glass of Figure 9, the 4:1 Prism-PET module of Figure 10, and the 9:1 combined Prism-PET module of Figure 11. The top half of Figure 12 is a 1D Gaussian histogram showing the crystal separation in the x-direction from the corner, edge, and center readout pixels of the modules of Figures 9-11. The bottom half of Figure 12 is the filtered energy spectrum from the center crystal of Figures 9-11 with (13%, 9%, and 10%) and without (20%, 14%, and 16%) DOI correction.

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

[18] . Using lead collimation, the modules were exposed with a Na-22 source at five different crystal depths (2, 6, 10, 14, and 18 mm). The source was placed inside a lead cylinder with a 1 mm pinhole. The pinhole was attached to a single 1.4 × 1.4 × 20 mm pinhole on one side of the DOI alignment module and on the other side of the reference module. 3The DOI resolution of the crystal was calculated using the width of the Gaussian peaks transformed into DOI space (Figure 14). DOI resolution is depth-dependent and equals the FWHM of the Gaussian histogram. The crystal-specific overall DOI resolution was calculated as the average of the DOI resolution across the measured depth

[18] . A typical center crystal for each module was used to calculate the DOI resolution of each module.

[0033] The spatial performance of the disclosed Prism-PET module is characterized in comparison to a standard uniform glass light guide module using flood histograms of the fabricated module (Figures 9-11). Glass light guide modules suffer from edge and corner effects, resulting in poor position-dependent crystal separation. Prism-PET enables excellent crystal separation without edge and corner artifacts across the entire detector array, something not achieved with previous 4:1 coupled detector modules with single-ended TOF-DOI readout [8, 19, 20]. Similar results were also shown with the 9:1 coupled Prism-PET module (Figure 11), measuring 3.2 × 3.2 mm. 2 The TOF-DOI PET detector module with SiPM pixels shows uniform crystal separation of less than 1 millimeter. Creating a 1D event positioning histogram (in the x-direction) confirms that the disclosed Prism-PET has uniform crystal separation performance at the center, edges, and corners. Prism-PET also achieves 14% and 16% energy resolution with the 4:1 and 9:1 coupling modules, respectively, when DOI correction is used, while the uniform light guide only achieves 20% energy resolution (bottom graph of Figure 12).

[0034] Figure 13 shows the DOI of the modules of Figures 10 and 11 according to an embodiment 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 pseudo-prismatic light guide (Figure 13(b)). Figure 13(c) shows a comparison of DOI resolution based on the light guide used, demonstrating that the Prism-PET detector module of the present disclosure achieves a two-fold improvement in DOI resolution over the uniform glass light guide, as experimentally measured for a single center crystal in each module. The measured DOI resolution for the glass light guide was 5 mm FWHM, showing strong agreement with previously reported results

[19] . The Prism-PET module achieved a FWHM DOI localization of 2.5 mm, 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 light guide due to refractive index matching, and 3) an improved 180° deflection of the upward-moving optical photon paths due to the right-angle triangular prism geometry, all of which improve light-sharing between crystals coupled to the same pseudoprism. Using DOI information, both time resolution and energy resolution can be improved: time resolution by performing a deconvolution of depth-specific photon transport inside the scintillator, and energy resolution by constructing depth-specific photopeaks [8, 18]. After applying DOI-based corrections, embodiments of the present disclosure achieved energy resolutions of 9% and 10%, respectively, with 4:1 and 9:1 coupled Prism-PET modules, and 13% with a uniform light guide. It should be noted that the DOI and energy resolution values ​​improved and worsened slightly, respectively, after implementing SiPM saturation correction

[22] , and as a result, the reported values ​​actually represent the relative performance of the disclosed Prism-PET compared to the uniform light guide module, rather than the absolute performance.

[0035] Figures 14(a)-(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 shows 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 a DOI histogram generated by taking the w histogram of Figure 14(a) and multiplying it by the slope of the linear fit of Figure 14(b). Figure 14(d) shows the DOI resolution at each acquisition depth based on the width of the Gaussian curve of 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 throughput, delivered dose, and image quality. Monte Carlo simulations using highly 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 the sensitivity gain based on (a) geometric sensitivity and (b) coincidence time resolution (CTR) relative to the time-of-flight readout (TOF), as shown in Equation (1), where the sensitivity gain is equal to the square of the SNR gain

[24] .

[0037]

number

[0038] where D is the diameter of the object being imaged, 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)-(d) show sensitivity graphs according to embodiments of the present disclosure. Figure 15(a) shows the dimensions and geometric extents of example Siemens Biograph Vision, Explorer Total-Body PET scanners, and Prism-PET brain scanners. 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 temporal 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 based on an embodiment of the present disclosure, compared 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 significantly improves geometric efficiency (Figure 15(b)), but at the expense of significantly increased parallax error and partial volume effects, which can be mitigated by performing depth-of-interaction (DOI) readout

[26] . As a result, only small-diameter scanners for specific organs should be built with detector modules with DOI localization capabilities, such as our Prism-PET module.

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

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

[23] (Figure 15(c)).

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

[0044] 16(a)-(f) illustrate Compton interactions according to an embodiment of the present disclosure.

[0045] Regarding Compton interactions, the disclosed Prism-PET allows for Compton scattering energy resolution (and therefore localization) due to its deterministic light-sharing pattern. Assume a Prism-PET light guide with a 16 x 16 array of lutetium LYSO crystals coupled 4:1 to an 8 x 8 array of silicon photomultiplier (SiPM) pixels. Based on the approximation that each 511 keV gamma ray generates a signal on four different pixels due to light sharing, the light-sharing ratio between all crystals belonging to the same pseudoprism can be measured directly from the flood data using photoelectric events. Using this information, the energies of the primary interaction (i.e., the recoil electron) and the secondary interaction site (i.e., the scattered gamma ray) can be resolved. After the resolved energies are obtained, the two independently absorbed events within the scintillation block can be localized, and the scattering angle and DOI can be determined. Due to the change from random light sharing of photoelectric events to a deterministic pattern, it is possible for the Prism-PET module of the present disclosure to distinguish side-by-side Compton scatter events (FIGS. 16 and 17).

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

[0047] 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 is the maximum deposited energy in the SiPM coupled to the crystal pixel where the interaction occurs, and E A2,3,4 and E B2,3,4 is the accumulated energy in the adjacent pillar due to light leakage at the bottom (from the SiPM side) and at the top via the prismatic mirror light guide. The experimental results in Supplementary Fig. 16(a) show the four known parameters E, which correspond to the energy detected by each of the four pixels after a side-by-side Compton scattering event. 1-4 and the total gamma particle energy deposited is given by equation (4). E γ =E A +E B (4)

[0049] The energy of the components of the Compton scattering event, i.e., E A1-4 and E B1-4 Note that is unknown. Writing these equations in terms of the measured energies gives us 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 stored energy in adjacent pillars is correlated. Consider the inset in Figure 16(d). The maximum stored energy occurred in the upper left SiPM. Assuming that the sharing fraction with three neighboring crystals depends on the proximity of the neighboring crystals to the crystal where the interaction occurred, and using the Pythagorean theorem to form a right triangle using the centers of the three neighboring crystals as vertices, we obtain Equation (6).

[0051]

number

[0052] In the above equation, for example, d 12 is the distance between the centers of the primary SiPM 1 and the adjacent SiPM 2. Substituting Equation 6 into Equation 5, we obtain Equation (7). E1=EA1+EB2 E2=EA2+EB1 E3=0.7EA2+EB2 E4=E A2 +0.7E B2 (7)

[0053] This has four equations and four unknowns. Note that in practice, due to some unavoidable small misalignment between the prismatic mirror light guide and the scintillator pillars, the shared fractions will have spatial deviations 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 the two resolved components of a measured side-by-side Compton scattering event based on the above analysis.

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

[0055] As shown in Figure 15(e)-(f), the {E A1 ,E B1} and {w A ,w BThe percent error of our estimate of} is about 10%. Using a convolutional neural network as an estimator can further reduce this error, especially since millions of gamma events can be collected as a training data set using flood histogram experiments.

[0056] An example of how a Compton event, where the recoil electron and scattered gamma ray are completely absorbed by adjacent scintillators in two different SiPMs, can be decomposed into its component parts is shown in Figures 15(a)-(d). Calculating the DOI variate w using classical Compton decomposition results in a full width at half maximum (FWHM) error of 11% (Figure 16(e)). Furthermore, Compton decomposition results in a FWHM energy error of 15% (Figure 16(f)).

[0057] Figure 17 illustrates photoelectric and Compton interactions according to an embodiment of the present disclosure, showing a graph of the random light sharing pattern of a glass light guide over a graph of the deterministic light sharing pattern of a Prism-PET module of the present disclosure. Figure 17 shows experimental results of several examples of Compton events absorbed in adjacent crystals in a Prism-PET module of the present disclosure versus a module with a flat glass light guide.

[0058] The light-sharing pattern within the glass light guide module is random, which makes it difficult (and in most cases impossible) to resolve the detected energy into the constituent energies of scattered photons and recoil electrons. Due to the right-angled triangular prism geometry, the light-sharing pattern is deterministic in the Prism-PET module, which makes it practical to resolve an event into its constituent energies based on the known light-sharing ratio between the crystals.

[0059] Thus, a particle detector is provided that includes 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 an apex end of the scintillator array, and at least one processor in operative communication with the plurality of detectors. The at least one processor includes a plurality of supervised machine learning algorithms, including convolutional and recurrent networks, configured to perform 3D gamma-ray localization of at least one interaction site within at least one scintillator crystal of the plurality of scintillator crystals. The at least one processor is configured to recover at least one Compton event scattered between the plurality of scintillator crystals and localize the at least one Compton event at the scintillator level based on the 3D gamma-ray localization. The at least one processor is further configured to determine a scattering angle based on the at least one Compton event and the DOI information. The at least one processor is further configured to localize the at least one Compton event based on resolved energy of at least two interactions absorbed within the plurality of scintillator crystals, the resolved energy being based on at least one light sharing pattern, and the at least one light sharing pattern being based on positions of the plurality of scintillator crystals relative to the plurality of detectors and the plurality of pseudoprisms.

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

[0061] Thus, 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 encoding in whole-body and total-body clinical PET scanners

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

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

[31] . Uniform light guides are suboptimal for this task because the SiPM patterns of individual events are random. In contrast, our Prism-PET module produces deterministic light sharing patterns 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, the most probable and most difficult to analyze scattering events, into their constituent energies, spatial locations, and DOIs. Compton scatter recovery is particularly important for maintaining high sensitivity in detector modules with small scintillator crystals because as the crystal size decreases, scattered photons are more likely to be absorbed in crystals different from the primary interaction site

[32] .

[0063] An embodiment of the present disclosure provides the Prism-PET detector module, a true single-ended analog of a dual-ended depth-encoded readout, using an efficient reflector that bends light 180° to improve light sharing. A 2.5 mm FWHM DOI resolution is achieved, with up to 9:1 scintillator-to-SiPM coupling for high spatial resolution, while 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 reflector consists of an optimized pattern of segmented pseudo-prismatic light guides, which efficiently redirects the scintillation photon path from the primary crystal to the selected nearest-neighbor SiPM, thus closely mimicking the operation of a dual-ended readout detector. This produces an anisotropic and deterministic signal pattern that can be used to resolve side-by-side Compton scattering events into their constituent energies and DOI information for scatter retrieval purposes. Thus, high and uniform spatial resolution is achieved (9:1 coupling of approximately 1 mm crystals, no edge and corner artifacts due to improved light sharing, and spatial blur due to Compton scattered photons is reduced by scatter recovery), high sensitivity is achieved (20 mm thick detector and inter-crystal Compton scatter recovery), and good energy and time resolution is achieved in a compact system (especially after applying DOI correction) (DOI encoding eliminates parallax errors and allows for smaller ring diameters). This unique combination of features allowed us to develop 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 embodiments thereof, those skilled in the art will recognize that various changes in form and detail can be made therein without departing from the spirit and scope of the invention as defined by the appended claims and equivalents thereof. The claims set forth below should not be construed as means-plus-function elements without the express use of "means for" or "steps for." References [1] Mankoff DA, Farwell MD, Clark AS, Pryma DA. Making molecular imaging a clinical tool for precision oncology: a review. JAMA Oncol. 2017;3:695-701. [2] Slifstein M, Abi-Dargham A. Recent developments in molecular brain imaging of neuropsychiatric disorders. Semin Nucl Med. 2017;47:54-63. [3] Van Sluis J, De Jong J, Schaar J, et al. Performance characteristics of the digital biograph vision PET / CT system. J Nucl Med. 2019;60:1031-1036. [4] Moses WW. Fundamental limits of spatial resolution in PET. Nucl Instr and Meth A. 2011;648:S236-S240. [5] Miyaoka RS, Lewellen TK, Yu H, McDaniel DL. Design of a depth of interaction (DOI) PET detector module. IEEE Trans Nucl Sci. 1998;45:1069-1073. [6] Stickel JR, Cherry SR. High-resolution PET detector design: modelling components of intrinsic spatial resolution. Phys Med Biol. 2005;50:179-195. [7] Seifert S, Schaart DR. Improving the time resolution of TOF-PET detectors by double-sided readout. IEEE Trans Nucl Sci. 2015;62:3-11. [8] Pizzichemi M, Polesel A, Stringhini G, et al. On light sharing TOF-PET modules with 3 mm depth of interaction and 157 ps FWHM coincidence time resolution. Phys Med Biol. 2019;64:155008. [9] Yang Y, Wu Y, Qi J, et al. A prototype PET scanner with DOI-encoding detectors. J Nucl Med. 2008;49:1132-40.

[10] Kuang Z, Wang X, Fu X, et al. Dual-ended readout small animal PET detector by using 0.5 mm pixelated LYSO crystal arrays and SiPMs. Nucl Instr and Meth A. 2019;917:1-8.

[11] Abreu MC, Aguiar JD, Almeida FG, et al. Design and evaluation of the clear-PEM scanner for positron emission mammography. IEEE Trans Nucl Sci. 2006;53:71-77.

[12] Du J, Yang Y, Bai X, et al. Characterization of large-area SiPM array for PET applications. IEEE Trans Nucl Sci. 2016;63:8-16.

[13] Schmand M, Eriksson L, Casey ME, et al. Performance results of a new DOI detector block for a high resolution PET-LSO research tomograph HRRT. IEEE Trans Nucl Sci. 1998;45:3000-3006.

[14] Seidel J, Vaquero JJ, Siegel S, et al. Depth identification accuracy of a three layer phoswich PET detector module. IEEE Trans Nucl Sci. 1999;46:485-490.

[15] Gonzalez-Montoro A, Aguilar A, Canizares G, et al. Performance study of a large monolithic LYSO PET detector with accurate photon DOI using retroreflector layers. IEEE Trans Rad Plasma Med Sc. 2017;1:229-237.

[16] Ito M, Lee MS, Lee JS. Continuous depth-of-interaction measurement in a single-layer pixelated crystal array using a single-ended readout. Phys Med Biol. 2013;58:1269-1282.

[17] Kuang Z, Yang Q, Wang X, et al. A depth-encoding PET detector that uses light sharing and single-ended readout with silicon photomultipliers. Phys Med Biol. 2018;63:045009.

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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 said 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. The apparatus of claim 1 , wherein a first pseudoprism of the plurality of pseudoprisms is configured to redirect particles between 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 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 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 at an 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. 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 positions of the plurality of scintillator crystals relative to the plurality of detectors and the plurality of pseudoprisms.

18. 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. 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 the measured photoelectric events, the 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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