Transmission imaging in forward-scattered gamma-ray-based PET scanners with coincidence detection

Abstract

A novel method for acquiring transmission scan data in a PET scanner is disclosed by incorporating one or more fixed gamma ray sources that provide forward scattered gamma photons that can be used as transmission imaging radiation.

Description

[Technical field]

[0001] This application claims the benefit of U.S. Provisional Application No. 63 / 018,654, filed May 1, 2020, the disclosure of which is incorporated herein by reference in its entirety.

[0002] FIELD OF THE DISCLOSURE This disclosure relates to nuclear imaging, and more particularly, to attenuation correction in positron emission tomography (PET). [Background technology]

[0003] PET is a nuclear medicine imaging technique that creates three-dimensional images of the distribution of positron-emitting isotopes in the body. When a radioisotope undergoes positron-emitting decay (also known as positive beta decay), it releases the antimatter equivalent of an electron. As the positron loses energy, it eventually meets an electron and annihilates, producing a pair of annihilation (gamma) photons that usually move in opposite directions. By detecting the time coincidence of a pair of gamma photons, a PET system determines the line along which the annihilation occurred.

[0004] Attenuation correction in PET imaging is a key component in producing artifact-free, quantitative data. Most attenuation correction is based on transmission scans acquired before, during, or after the emission scan. Therefore, obtaining high-quality transmission data is beneficial for achieving accurate attenuation correction.

[0005] Attenuation in PET imaging is the loss of detection of true coincidence events due to absorption in the body and scattering outside the detector field of view (FOV). In PET imaging, two photons must exit the patient at the same time to be detected as a true event. In PET imaging, the loss of detection of true coincidence events due to attenuation can range from 50-95%.

[0006] Loss of counts due to attenuation increases image noise, image artifacts, and image distortion. Without attenuation correction, significant artifacts that may occur in whole-body PET scans include: (1) a significant amount of activity at the surface of the body due to the relative loss of attenuation compared to deeper structures, (2) a distorted appearance of the image due to the difference in attenuation depending on the direction of the radioactivity emanating from a highly radioactive area (e.g., the bladder), and (3) a relative increase in radioactivity due to diffusion in tissues with relatively low attenuation (e.g., the lungs). Therefore, attenuation correction of the data is necessary for accurate measurement of PET scan data.

[0007] In a PET / CT system, x-rays from a CT scan can be used to build an attenuation map of density differences throughout the imaged region, which can be used to correct for the absorption of photons emitted by the decay of fluorodeoxyglucose in an emission scan. However, such an integrated PET / CT system requires the integration of additional CT scanning hardware with the PET imaging hardware. It would therefore be useful to be able to acquire transmission-type scan data with a PET scanner to build an attenuation map without the need for additional transmission scanning hardware, such as a CT scanner. This ability to acquire transmission-type scan data in a PET scanner system to create an attenuation map could also be useful in a PET / MR system, since MR systems do not measure everything in the system that attenuates radiation in a PET scan. For example, PET / MR systems use coils and associated electronics and cabling that attenuate annihilation photons, but by design these are ignored so that they do not appear in the MR image, and their effect on annihilation photon attenuation cannot be considered.

[0008] Several PET scanner designs have been proposed that derive attenuation corrections from background radiation emitted in the detector's own lutetium-based scintillation crystals (e.g., LSO or LYSO scintillation crystals). However, the intensity of LSO background radiation is generally too low to be used in patient scans of normal duration. Therefore, improved PET scanners that can generate transmission imaging data without the use of CT hardware and can create attenuation μ-maps more quickly would be beneficial. Summary of the Invention [Problem to be solved by the invention]

[0009] A PET scanner system is provided that includes a gantry, a plurality of PET detector ring assemblies disposed within the gantry, each detector ring assembly including a plurality of PET detectors arranged in a ring shape about a central opening, the plurality of PET detector ring assemblies being coaxially disposed along a longitudinal axis defined through the central opening, a patient tunnel extending through the central opening of the coaxially disposed PET detector ring assemblies, the plurality of PET detector assemblies being coaxially disposed along a length of the patient tunnel, each of the plurality of PET detectors including a detector and one or more scintillator crystals associated with the detector, and one or more stationary gamma radiation sources disposed in each PET detector ring assembly within the gantry.

[0010] Also provided is a method of using one or more fixed gamma radiation sources in a PET scanner as a transmission scan radiation source for generating scan data used to generate an attenuation map, the PET scanner including a plurality of PET detector ring assemblies, the method including: (a) providing one or more fixed gamma radiation sources in the PET detector ring assemblies, the fixed gamma radiation sources being located outside the detector ring assemblies; (b) detecting gamma photons emitted from the fixed gamma radiation sources, which are forward scattered gamma photons that pass through scintillator crystals in a first set of detector blocks in the PET detector ring assemblies, scatter forward, cross a field of view (FOV) of the PET scanner, and are detected by coincidence electronics of the PET scanner. (c) acquiring list mode data from a blank transmission scan with no radioactivity in the FOV; (d) acquiring list mode data from a transmission scan with a target object in the FOV; (e) generating an attenuation map by comparing the list mode data from the blank transmission scan with the list mode data from the transmission scan with the target object in the FOV; and (f) applying the attenuation map of step (e) to the list mode data from the emission scan to apply attenuation correction to the list mode data from the emission scan.

[0011] Also provided is a method for improving the quality of the transmission scan signal from forward scattered gamma photons by applying time-of-flight (TOF) considerations to transmission image scanning using a fixed gamma source. The method includes: (a) calculating a TOF for a scattered photon, which is a photon from a gamma ray source, that has undergone Compton scattering in a first scintillator crystal of a first detector block to reach a second scintillator crystal of a second detector block based on a distance between the two scintillator crystals; (b) defining a time window, where the time window has a width centered on the calculated TOF; (c) measuring the TOF of an actual scattered gamma photon originating from the first scintillator crystal in the FOV of the PET scanner at a scanned object in the FOV of the PET scanner; (d) comparing the measured TOF derived from (c) above with the calculated TOF to identify measured TOFs that fall within the time window; (e) identifying scattered gamma photons corresponding to those measured TOFs that fall within the time window as transmission source (transmission scanning source) events resulting from Compton scattering in the first scintillator crystal, thereby distinguishing the transmission type data from gamma emission annihilation events and random events in the scanned object; [Brief description of the drawings]

[0012] [Figure 1A] FIG. 1A is a schematic diagram of a PET system. [Figure 1B] FIG. 1B is a schematic diagram of a PET detector ring assembly having a fixed gamma radiation source according to an embodiment of the present disclosure. [Figure 1C] FIG. 1C is a schematic diagram of a PET detector ring assembly having multiple fixed gamma radiation sources according to an embodiment of the present disclosure. [Diagram 2] FIG. 2 is a schematic diagram of a gamma source located behind the PET detector and mounted within a tungsten radiation shield in the "ON" configuration such that gamma rays from the gamma source irradiate the detector scintillation crystal. [Diagram 3]FIG. 3 is a schematic diagram of the gamma source in the "OFF" configuration, positioned behind the PET detector, with a tungsten radiation shield absorbing gamma radiation from the gamma source. [Figure 4A] FIG. 4A is a flow chart diagram of a method according to one embodiment of the present disclosure. [Figure 4B] FIG. 4B is a flow chart diagram of a method according to another embodiment of the present disclosure. [Diagram 5] (a) and (b) are schematic diagrams of the front and side, respectively, of the PET scanner used in the experimental validation, showing the location of different radioactive elements with respect to the measurements. [Figure 6] FIG. 6 is a two-dimensional energy histogram from scans 1, 2, 3, and 4 performed in the experimental validation. [Figure 7] FIG. 7 is a grayscale display showing the distribution of singles rates for 228 detector blocks in the six-ring Biograph Vision PET / CT scanner used in the experimental validation. [Figure 8] FIG. 8 shows the net-true sinograms derived from scans 1, 4, and 5 performed in the experimental validation. [Figure 9] FIG. 9 is a plot showing the ratio of Scs1-blank / Sbg-blank for rebinned sinograms. [Figure 10] FIG. 10 shows a sinogram of the estimated 38 sources. DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

[0013] This description of exemplary embodiments is intended to be read in conjunction with the accompanying drawings, which are to be considered part of the entire written description.

[0014] PET emission data collected in a conventional PET scan contains all the information about the physical effects that the emitted photons undergo before being detected by the detector. To reconstruct the collected emission data, corrections were applied to the emission data in order to reconstruct the true measured emission events. System corrections such as normalization, random coincidence correction, and dead time correction are independent of the objects in the FOV and depend mainly on the system and count rate of the radioactivity irradiating the PET detector inside and outside the FOV. Other corrections are object dependent, such as attenuation correction and scatter estimation, and require attenuation information for the object materials in the scanner's FOV.

[0015] [Typical operation of a PET scanner] 1A illustrates an example of a PET scanner system 200 in which the inventive concepts disclosed herein can be implemented. A human subject 4 to be scanned is shown positioned within a gantry 210 of the PET scanner system 200. The gantry 210 includes a number of PET detector rings 100. Each detector ring includes a number of scintillation crystals 216 and associated detectors 213. An open space in the center of the detector rings 100 defines a patient tunnel T of the gantry 210 and the FOV of the PET scanner. When a PET scan is performed, a positron-emitting radioisotope 6 is introduced into the human subject 4 on a metabolically active molecule that is then carried to the organ of interest by the human subject's bloodstream.

[0016] When a positron emitted from a radioisotope currently present in the human subject encounters an electron, the two are annihilated, resulting in two gamma photons 7 traveling in roughly opposite directions. The annihilation event is identified by the temporal synchronicity of the detection of the two gamma photons by the two oppositely positioned detectors, due to the gamma photons interacting with the detector's scintillation crystals 216. That is, the emission of the gamma photons is detected by each detector 213 substantially simultaneously. When two gamma photons traveling in opposite directions strike corresponding oppositely positioned detectors, creating a coincidence event, the photons identify a line of response (LOR) along which the annihilation event occurred.

[0017] An image of metabolic activity (nuclear medicine image) in the human subject 4 is reconstructed by computer analysis. The PET scanner system 200 includes a system controller 290 connected to and in communication with the detector ring 100. The PET scanner system 200 further includes a data processing unit (event detection unit) 220 that determines and evaluates coincidence events generated by pairs of gamma rays and transfers this information to an image processing unit (computation unit) 230. The detector pairs associated with each LOR generate many coincidence events during a measurement session (i.e., a data acquisition scan). The PET scanner system 200 further includes at least one machine-readable storage medium 250 that is encoded with computer program code. When the computer program code is executed by the system controller 290, the system controller performs various operational functions of the PET scanner system 200.

[0018] [Improvements according to the present disclosure] With reference to FIG. 1B, a novel configuration for a PET scanner system is provided that allows for simultaneous generation of transmission scan data during a PET scan session according to the present disclosure. To provide the PET scanner with the facility to generate simultaneous transmission scan data during a PET scan session, one or more stationary gamma radiation sources 130 are incorporated into each of the PET detector ring assemblies 100 in the gantry 210. Each of the PET detector ring assemblies 100 includes a plurality of detector electronics assemblies 110 arranged in a ring shape around a patient tunnel T. Each detector electronics assembly 110 includes, among other things, a detector 213 (photodetector) and one or more associated scintillator crystals 216. FIG. 1B is schematic, and for ease of illustration, the detector 213 and the associated one or more scintillator crystals 216 are collectively referred to as a PET detector block 214. This designation will be followed throughout the remainder of the disclosure.

[0019] The fixed gamma radiation source 130 is positioned within the PET gantry 210 near and behind the PET detector block 214 and away from the patient tunnel T such that a portion of the gamma photons from the gamma radiation source scatter forward through a first scintillator crystal associated with the first PET detector block 214a and then across the FOV of the PET scanner to a second scintillator crystal associated with a second PET detector block 214b on the opposite side of the PET scanner's FOV. The gamma radiation source 130 is preferably positioned within the PET gantry 210 for practical reasons of shielding and containing gamma radiation within the PET scanner.

[0020] In Figure IB, arrow 11 represents an exemplary gamma photon scattered forward through a first scintillator crystal of a first detector block 214a, and arrow 12 represents the path the scattered gamma photon then takes across the FOV of the PET scanner to a second scintillator crystal of a second detector block 214b. The scattered gamma photon is represented by 12. Because the forward scattered gamma photon 12 passes through the FOV, by identifying a sufficient number of such forward scattered gamma photons via the coincidence electronics of the PET scanner, the scattered gamma photon 12 can be used for transmission imaging in the PET scanner.

[0021] In some preferred embodiments, the one or more fixed gamma radiation sources 130 in each detector ring assembly 100 are positioned behind the detector blocks 214. This configuration allows more of the gamma photons emitted from the gamma radiation source 130 to be forward scattered through the scintillator crystals in the nearby detector block or blocks 214, resulting in an increased radiation dose for transmission imaging.

[0022] 1C, providing and identifying a sufficient number of forward-scattered gamma-ray photons for transmission imaging can be more practically achieved by providing multiple fixed gamma-ray sources 130 in each of the PET detector ring assemblies 100. All of the gamma-ray sources 130 are located behind the PET detector blocks 214, as illustrated in FIG. 1B. A location behind the PET detector blocks 214 means any location on the side of the PET detector blocks 214 facing away from the patient tunnel T. The gamma-ray sources 130 can be positioned anywhere in the gantry, as long as they are behind the PET detector blocks 214. However, in some preferred embodiments, each fixed gamma-ray source 130 is arranged side-by-side between a pair of detector electronics assemblies 100, i.e., between two sets of PET detector blocks 214, as shown in FIGS. 1B and 1C, to maximize detection of forward-scattered photons by the detectors. If the gamma source is directly behind the detector, it corresponds to zero degree scattering, in which case the scattering process leaves very little energy in the scattered detector and cannot be detected. The scattering process cannot be detected until the energy left by the scattering exceeds the detector electronics threshold (the specific threshold varies depending on the PET scanner manufacturer, but as an example, is about 150 keV). Therefore, the most preferred scattering angle is about 40 degrees or more. This can be achieved by placing the fixed gamma source 130 behind the PET detector block 214 and between a pair of detector electronic assemblies 110. If the gamma source is placed behind the PET detector block 214 and between a pair of detector electronic assemblies 110, the two detector blocks 214 are well positioned for scattering to both sides.

[0023] Every fixed gamma radiation source 130 emits some gamma photons that are forward scattered through the FOV and therefore may provide more radiation required for transmission imaging. Figure 1C is a schematic diagram only, and the number of detector electronic assemblies, detector blocks 214, and gamma radiation sources 130 shown do not represent the actual number that would be present in an actual PET scanner system employing embodiments of the present disclosure.

[0024] With this configuration, a scattering event in a first scintillator crystal in the first PET detector block 214a provides a start signal to the coincidence electronics, and detection of a scattered gamma photon 12 by a second scintillator crystal in the second PET detector block 214b provides a stop signal to the coincidence electronics to detect coincidences. Photons from the gamma source undergo Compton scattering in the scintillation crystal of the first PET detector block 214a. In the scintillator, part of the photon's energy is transferred to an electron, which is immediately stopped by interaction with the scintillator, resulting in a flash of light in one scintillator crystal of the first PET detector block 214a, which provides a start signal for the coincidence electronics. The forward scattered gamma photon 12 with reduced energy travels across the FOV and is detected by the second detector block 214b, providing a stop signal for the coincidence electronics. This coincidence detection identifies two scintillation crystals in the PET detector ring assembly 100 and provides a transmission sinogram by counting the gamma photons 12 traveling through the FOV, as the two points in space define a straight line path of the scattered photons across space.

[0025] Thus, generating transmission scan data using radiation from the fixed gamma radiation source 130 without a patient in the FOV produces a blank transmission scan sinogram. By comparing the blank transmission sinogram to a transmission sinogram from a transmission scan having a scanned target volume within the FOV, a mu map (spatial map of photon attenuation coefficients) of the target volume can be constructed for attenuation correction of the emission PET scan data for the target volume.

[0026] A variety of radioisotopes can be used as the gamma ray source 130. Some examples are cesium-137, cobalt-60, and sodium-22. In some preferred embodiments, Cs-137 is used as the gamma ray source 130 material. Cs-137 is preferred as the gamma ray source 130 material because, although its gamma ray energy is somewhat higher than the 511 keV commonly used in PET, it is not so high that the gamma rays are outside the range typically used in many PET detectors. In addition, the half-life of Cs-137 is 30 years, which is advantageous since it does not require source replacement during the expected life of the PET scanner. In addition, while other isotopes have multiple gamma rays with different energies, Cs-137 only has one gamma ray per decay. In addition, the Cs-137 isotope is widely available, as it is one of the main components of radioactive waste after being stored for several years.

[0027] According to the present disclosure, a PET system configured to perform a method for simultaneously acquiring PET emission and transmission scan data by incorporating a fixed gamma ray source and time-of-flight (TOF)-based coincidence measurement includes a PET detector ring assembly 100 illustrated in Figures 1B and 1C. The PET detector ring assembly 100 includes a plurality of PET detectors arranged in a ring shape around a central opening, where the plurality of PET detector ring assemblies are coaxially arranged along a longitudinal axis L (see Figure 1A) defined through the central opening. A patient tunnel T extends through the central opening of the coaxially arranged PET detector ring assembly 100. The plurality of PET detector assemblies are coaxially arranged along the length of the patient tunnel T.

[0028] Each detector electronics assembly 110 includes a PET detector block 214. Each detector block includes a detector 213 and one or more associated scintillator crystals 216. A number of gamma radiation sources 130 are located behind the PET detector block 214, as illustrated in FIG. 1B. By located behind the PET detector block 214, we mean anywhere on the side of the detector block 214 facing away from the patient tunnel T. The gamma radiation sources 130 are preferably located within the PET gantry 210 for practical reasons of gamma radiation shielding and containment within the PET scanner.

[0029] 1C, in this arrangement of the gamma radiation source 130 and the PET detector block 214, gamma photons 11 enter the first set of detector blocks 214 from the back side. Some of the gamma photons undergo Compton scattering in the scintillator crystals 216 of the first set of PET detector blocks 214 and undergo forward scattering across the FOV of the PET scanner toward the second set of PET detector blocks 214 on the opposite side of the FOV. The scattered gamma photons are then stopped when they encounter the scintillator crystals 216 associated with the second set of PET detector blocks, with each scattered photon producing a flash of light that serves as a stop signal for the coincidence electronics. The scattered photons are indicated by arrows 12. This provides the basis for transmission imaging when the target subject is positioned in the patient tunnel T and is within the FOV.

[0030] In some preferred embodiments of the disclosed systems, the gamma radiation source 130 can include a collimating shield such that gamma photons directed towards the PET detector 214 can shine at full intensity while gamma photons directed in other directions can be absorbed by the shield to reduce potential radiation exposure to people near the scanner. The particular configuration for the collimating shield can be designed to take into account the particular radiation source selected for a given system and optimize the transmission imaging achieved.

[0031] 2 and 3 are conceptual diagrams illustrating an example of one gamma radiation source assembly 130A and associated PET detector ring 100. The gamma radiation source assembly 130A is positioned such that the gamma radiation source 130 is located above a gap between two PET detector blocks 214 in the PET detector assembly 100. The gamma radiation source assembly 130A comprises a gamma radiation source 130 and a radiation shielding housing 136 for the gamma radiation source 130. The gamma radiation source assembly 130A is configured to hold the gamma radiation source 130 and to be switched between an ON configuration and an OFF configuration. In the ON configuration, the remaining gamma radiation photons from the gamma radiation source 130, other than those from the gamma radiation source 130 traveling in a direction toward the PET detector block 214, are substantially absorbed by the housing 136. Radiation absorption is never 100% in reality. Thus, "substantially absorbed" means that the gamma photons are strongly attenuated. The radiation shielding housing 136 may be constructed from a radiation attenuating material such as tungsten.

[0032] In the illustrated example, the gamma radiation source 130 is attached to the end of an articulating rod 134 that can manipulate the position of the gamma radiation source 130 within the housing 136 between an ON configuration and an OFF configuration. The radiation-shielded housing 136 can include a channel 138, and in the OFF configuration, the gamma radiation source 130 can be retracted into the channel 138, as shown in FIG. 3, so that gamma photons emitted from the source are absorbed by the housing 136. In the ON configuration, the articulating rod 134 extends the gamma radiation source 130 out of the channel 138. In the ON configuration, the gamma radiation source 130 can be positioned within an opening 135 that opens towards the detector block 214.

[0033] 4A is a flow chart 300 illustrating a method of using one or more fixed gamma radiation sources 130 in a PET scanner as radiation sources for a transmission scan to generate scan data (list mode data) that can be used to create an attenuation map. The method includes: (a) providing one or more fixed gamma radiation sources 130 in each PET detector ring assembly 100, where the fixed gamma radiation sources are located outside the PET detector ring assembly 100 (Box 310); (b) identifying forward scattered gamma photons 12 that are gamma photons emitted from the fixed gamma radiation sources and that are forward scattered through scintillator crystals in a first set of detector blocks 214 in the PET detector ring assembly, traverse the FOV of the PET scanner, and are detected by scintillator crystals in a second set of detector blocks in the PET detector ring assembly via the coincidence electronics of the PET scanner (Box 320); (c) acquiring list mode data from a blank transmission scan (i.e., no radioactivity in the FOV). (d) acquiring list mode data from a transmission scan with the target object within the FOV (Box 340); (e) simultaneously with step (d), acquiring list mode data from a PET scan with the target object within the FOV (Box 350); (f) generating an attenuation map (mu map) by comparing the list mode data from a blank transmission scan with the list mode data from the transmission scan with the target object within the FOV (Box 360); and (g) applying the attenuation map to the list mode data from the emission scan of step (e) to apply attenuation correction to the list mode data from the emission scan (Box 370).

[0034] In another aspect, a PET scanner system 200 is disclosed configured to perform a method described in flow chart 300. The PET system includes a plurality of detector ring assemblies 100 including a plurality of scintillator crystals; a machine-readable storage medium 250; and a system controller coupled to and in communication with the detector ring assemblies; wherein the machine-readable storage medium is encoded with computer program code such that, when the computer program code is executed by the system controller 290, the system controller performs a method including: The method includes: (a) providing one or more fixed gamma radiation sources in each PET detector ring assembly, where the fixed gamma radiation sources are disposed outside the detector ring assembly; (b) identifying, via coincidence electronics of the PET scanner, gamma photons emitted from the fixed gamma radiation sources that are forward scattered through scintillator crystals of a first set of detector blocks in the PET detector ring assembly, traverse an FOV of the PET scanner, and are detected by scintillator crystals of a second set of detector blocks in the PET detector ring assembly; and (c) acquiring list mode data from a blank transmission scan (i.e., no radioactivity in the FOV). (d) acquiring list-mode data from a transmission scan having a target object in its FOV; (e) simultaneously with step (d), acquiring list-mode data from a PET scan with the target object in the FOV; (f) generating an attenuation map (mu map) by comparing the list-mode data from a blank transmission scan with the list-mode data from a transmission scan in which the target object is within the FOV; and (g) applying the attenuation map to the list-mode data from the emission scan of step (e) to apply attenuation correction to the list-mode data from the emission scan.

[0035] In some embodiments, time-of-flight (TOF) considerations are applied to improve the quality of transmission scan signals from forward-scattered gamma photons. Referring to the flowchart 400 of FIG. 4B, applying TOF considerations to a transmission imaging scan using a fixed gamma source 130 includes (a) calculating the time-of-flight of a scattered photon, which is a photon from a gamma source that has undergone Compton scattering at a first scintillator crystal in a first detector block 214a, to reach a second scintillator crystal in a second detector block 214b based on the distance between the two scintillator crystals (see block 410). This step calculates the TOF required for a gamma photon to travel from the first scintillator crystal across the FOV to the second scintillator crystal on the opposite side in a given PET scanner. In other words, the distance of interest here is from the first detector block 214a (specifically, the scintillator crystal associated with the first detector) where the Compton scattering of the gamma photon occurs, to the second detector block 214b (specifically, the scintillator crystal of the second detector) where the simultaneously scattered gamma photon is detected. Next, (b) a time window (time width) is defined, where the time window has a width centered on the calculated TOF (see block 420). Next, (c) the TOF of the actual scattered gamma photon originating from the first scintillator crystal is measured with the scan target in the FOV of the PET scanner (see block 430). Next, (d) the measured TOF from (c) is compared to the calculated TOF, and the measured TOF that is within the time window is identified (see block 440). Next, (e) the scattered gamma photons corresponding to the measured TOF within the time window are identified as transmission source events resulting from Compton scattering in the first scintillator crystal, thereby distinguishing transmission type data from gamma emission annihilation events and random events within the scanned target object (see block 450). The transmission scan data thus obtained can be used to generate an attenuation map for correcting the primary PET emission scan data.

[0036] In some embodiments, step (c) of flowchart 400 can further include simultaneously acquiring PET emission scan data of a scan target in the FOV of the PET scanner.

[0037] In another aspect, a PET scanner system 200 is disclosed that is configured to perform the method described in flow chart 400. The PET system includes a plurality of detector ring assemblies 100 including a plurality of scintillator crystals; a machine-readable storage medium 250; and a system controller connected to and in communication with the detector ring assemblies. Here, the machine-readable storage medium is coded with computer program code such that, when the computer program code is executed by the system controller 290, the system controller executes a method including: (a) calculating a TOF for a scattered photon, which is a photon from a gamma ray source that has undergone Compton scattering at a first scintillator crystal in a first detector block, to reach a second scintillator crystal in a second detector block based on a distance between the two scintillator crystals. This step calculates the TOF required for a gamma photon to travel from the first scintillator crystal across the FOV to the second scintillator crystal on the opposite side in a given PET scanner. In other words, the distance of interest here is the distance from a first detector block (specifically, a scintillator crystal associated with the first detector) where Compton scattering of gamma photons occurs to a second detector block (specifically, a scintillator crystal in the second detector) where simultaneously scattered gamma photons are detected. Next, (b) a time window is defined, where the time window has a width centered on the calculated TOF. Next, (c) the TOF of the actual scattered gamma photons generated from the first scintillator crystal is measured with the scanned target body within the FOV of the PET scanner. Next, (d) the measured TOF from (c) is compared with the calculated TOF to identify the measured TOFs that are within the time window. Then, (e) among the scattered gamma photons corresponding to those measured TOFs, the scattered gamma photons that are within the time window are identified as transmission source events originating from Compton scattering (forward scattering) in the first scintillator crystal, thereby distinguishing transmission type data from gamma emission annihilation events and accidental events within the scanned target body.

[0038] The use of an improved fixed gamma source to acquire transmission scan data can be applied not only to PET / MR scanners, but also to PET / CT scanners.

[0039] Using Cs-137 as a gamma ray source, the inventors experimentally confirmed that forward scattered Cs-137 gamma rays provide improved transmission imaging in a PET scanner compared to using LSO background radiation by generating transmission imaging data faster to generate attenuation μ-maps without CT hardware. The experimental results are shown below.

[0040] [Experimental data] The disclosed invention was demonstrated using a 6-ring Biograph Vision PET / CT scanner with LSO scintillators and SiPM-based detectors. Each ring of the scanner contains 38 detector blocks of size 64 mm x 32 mm, and each block contains 200 crystals of size 3.2 x 3.2 x 20 mm. The attenuated radiation (511 keV) in the measurement was measured using a sealed 23 MBq detector. 68 A Ge source (germanium) was placed in a thin steel tube and a Cs-137 source sealed in a lightweight housing had a nominal intensity of 115MBq.

[0041] List mode data (LM) were acquired in five scans and can be described with reference to Figures 5(a) and (b). Scan 1 was a 30 minute acquisition with no activity near the scanner except for the LSO background, no objects in the FOV, not even a patient bed. Scan 2 was a 30 minute acquisition with the germanium source at position (A) close to the center of the scanner's axial travel, at or just beyond the end of the axial range. Scan 3 was a 15 minute scan with the cesium source placed at position B, approximately 25 cm beyond the ring of detectors. Scan 5 was the same as scan 4, except that the patient bed was placed in the FOV to provide a simple phantom. The detected radiation corresponded to a variety of physical phenomena, including normal PET coincidence (line DE in FIG. 5(a)); backscattering from 511 keV or 662 keV photons from A (line DF in FIG. 5(a)); forward scattering of photons from a Cs-137 source at location B (line GH in FIG. 5(a)); and background LSO scintillation crystals (line CI in FIG. 5(a)). The LSOs are distributed around the periphery of the PET detector ring 100, with just one exemplary location shown as (C) in FIG. 5(a).

[0042] The measurements were performed with unusual PET scanner settings: the energy window was wide open to accept events with energies between 160 and 730 keV, and the coincidence window was set to 6.64 ns, which corresponds to a maximum chord length of about one meter, both of which are limited by normal PET imaging.

[0043] The LM data contained the following information for each coincidence event: the energy signal of each photon, discretized in bins of width 2.8 keV; the crystal identification of each photon; the TOF; and a bit indicating whether the coincidence was fast or delayed. The data from each scan were separated for analysis in two ways. First, two-dimensional (2D) energy histograms were generated for the fast coincidences, and matching histograms were generated for the delayed coincidences, using 200 × 200 bins, without limiting the TOF. In these histograms, the energy E of the photon at one detector was A is the horizontal axis, and the energy E of the other coincident photon is the horizontal axis. B Second, we created a net-true coincidence sinogram without any energy restriction, using only events whose TOFs corresponded to the distance between the two crystals ±215 ps (picoseconds).

[0044] For transmission imaging, essentially all response lines can be sampled by efficiently using photons emitted from the fixed Cs-137 source 130 and scattering from nearly all crystals in the PET detector ring 100. The Cs-137 source should be close to the PET detector 214 but far enough away to maintain an acceptable singles rate. One way to do this is to distribute the fixed Cs-137 sources 130 within the PET detector ring 100 as shown in FIG. 1C. The Biograph Vision PET scanner uses 19 detector electronic assemblies 110, so 19 angular positions are shown in FIG. 1C. At each angle, two axial positions would be used, bringing the total number of sources to 38.

[0045] Although the 38 source structures were not used in the above experiment, the sinograms obtained can be estimated by the following calculation. The sinograms obtained from scan 1 and scan 4 are denoted as S bg-blank and S cs1-blankAlthough these are 3D sinograms, it appears sufficient for the analysis to sum over all slices and all sinogram segments and treat them as 2D functions of radial and angular sinogram coordinates. It is speculated that new designs could use different source strengths As rather than 115 MBq as in the study described here. In this case, a 38-source sinogram (S cs38-blank ) may be written as follows:

number

[0046] Since different scan times were used for the two input sinograms, S bg-blank was divided by 2 before calculation. In the formula, R is an operator that rotates the sinogram by one of 19 angles, and the factor 2 takes into account the source placement at the two axis positions. Rotation must be done carefully. After converting the sinogram from 180 degrees to 360 degrees, (1) was applied and then converted back to the 180 degree representation.

[0047] Using the cesium technique, it is possible to determine how fast a transmission scan can be performed while still giving approximately the same image quality as a long scan using the LSO background technique. The gains are approximately:

number

number

[0048] [result] Figure 6 shows 2D energy histograms from scans 1, 2, 3, and 4. In these, the energy range for each photon is 160-730 keV as described above. The bottom row shows the rapid coincidences, and the top row shows the net-true coincidences (prompt minus delayed). The middle row shows the net-true histograms with LSO background subtracted, corrected for different scan times. The column representing scan 1 shows the expected spectrum of beta and gamma radiation in one detector, and the expected energy spectrum with a 202 or 307 keV photopeak in the other detector. The column representing scan 2 shows a normal PET coincidence with several features of interest. Near the upper right corner, small circular patterns can be seen representing the 511 keV photopeak in each detector. Below and to the left of that, features of radiation scattered from the source or from either detector can be seen, and at the lower end of the energy range, another prominent feature representing 511 keV backscatter through an angle of approximately 180 degrees. Also, the shaded box in part 2-B shows E A +E B = 511 keV, and E A and E B represents the energy of a pair of coincidence events, referred to as A and B. These represent the backscattering of 511 keV photons, as in line D-F of Figure 5. Most of the backscattering is overlapped with the LSO background. In the column of scan 3, photons are emitted one at a time from the cesium source, showing random coincidences in the lower row, which disappear in the upper row where delayed coincidences have been subtracted. In this example, E for each pair of coincidence events A and B generated by 662 keV photons from the cesium source A +E B Another diagonal trace is shown with E = 662 keV. These represent backscattering of 662 keV photons from a cesium source. Again, this backscattering signal is strongly overlapped with the LSO background. Scan 4 column shows forward scattered radiation on a line like G-H in Figure 5. E A +E B= 662 keV is again shown. Of note is that these photon pair signatures do not overlap as strongly with the LSO background radiation. The more favorable energy range is a potential advantage of the forward scatter approach, in addition to the fact that the source does not need to be placed within the FOV, which is the best spare space for the patient.

[0049] For scan 4, Fig. 7 shows the distribution of the singles rate in the 228 detector blocks of the scanner in a grayscale display. The horizontal direction corresponds to the circumference of the scanner, and the vertical direction corresponds to the axial dimension of the scanner. The maximum singles rate is about 1 × 10 in the region close to the cesium source. 5 counts / block, whereas in areas dominated by singles from the LSO background radiation, the number was approximately 1 × 10 4 counts / block. The shape of the distribution confirmed that the source was located about 25 cm beyond the detector. The actual amount of radiation striking the detector is expected to have been more than this, because the singles rate includes only events with deposited energies above the activation threshold of about 150 keV, and roughly half of the 662 keV Compton scatterers are expected to produce recoil electrons at energies below that. Simple simulations suggested that about 40% of the incident radiation was detected in a single mode. Other relevant effects include absorption and scattering by materials between the detector crystal and the source, such as the electronics and a 3-8 mm thick aluminium channel to hold a coolant to stabilise the detector during use.

[0050] Figure 8 shows the net-true sinograms from scans 1, 4, and 5 (parts A, B, and C of the figure) and the ratio of the sinogram of scan 5 to that of scan 4 (part D). The sinograms have 520 bins in the radial direction and 399 bins in the angular direction. The LSO-background sinogram from scan 1 was largely featureless except for the expected pattern of high and low values ​​at the edges of the detector block and lower counts in the center due to the solid angle variation with position. The cesium spectrum from scan 4 showed a pronounced diagonal orientation characteristic of forward scattered radiation. As expected, the sinogram region corresponding to small-angle scattering was reduced in brightness because less than 150 keV was deposited on the scattering crystal, i.e., the electron recoil energy was too low to be detected. The transmission and ratio sinograms from scan 5 showed the gamma-ray shadow of the patient bed, suggesting that a map of the attenuation coefficient could be estimated from these measurements.

[0051] To illustrate and quantify the sensitivity gain due to cesium, sinograms from scans 1 and 4 were rebinned to 52x19 from 520 radial bins by 399 angular bins. The ratio of the rebinned sinograms in each of the coarse bins was calculated to account for the difference in scan time. The ratio varied between 1.00 and 4.77 (Figure 9). The 38-source sinograms were estimated from (1) and illustrated in Figure 10, which shows that their uniformity is similar to that from scan 1. The estimated 38-source sinogram had 19.6 times more counts in the middle and 23.2 times more counts at both ends of the radial range compared to the sinogram from scan 1. Using (3), the sensitivity gain for the 38-source configuration was calculated as follows:

number

[0052] It is of interest to quantify the efficiency with which radiation from the LSO background and the cesium source was detected in these experiments. Table 1 lists the source activities and sinogram count rates for scans 1-4. To estimate the corresponding efficiencies, 176 Lu or 137 The number of transmitted photons per decay of the Cs nucleus is taken as NΩ / 4π, multiplied by the geometric efficiency for the radiation that provides the starting signal for the coincidence electronics. In the case of the LSO, the starting signal comes from a cascade of beta and gamma radiation, and as Figure 6 shows, either 202 or 307 keV photons can be detected in the coincidence count, so N=2. In this case, 176 Since the Lu atoms were surrounded on all sides by scintillators, Ω was set to 4πsr. For cesium, the initial signal is thought to be due to a scattered 662 keV gamma ray. In this case, only scattered photons can be detected by coincidence counting, so N=1 was used and Ω was calculated from the position of the cesium source as the sum of the solid angles of all detector blocks whose back faces face the source. The table lists NΩ / 4π and a quantity called the coincidence efficiency. The coincidence efficiency is defined as follows:

number

[0053] [Table 1]

[0054] [Consideration of experimental data] The experimental data presented here demonstrated the feasibility of speed-up in transmission imaging based on coincidence counting, such as from an LSO background. This investigation showed that forward-scattered cesium gamma rays can provide such speed-up. The use of the forward-scattered gamma photon approach of the present disclosure offers the following advantages: a hardware fixed gamma source can be directly integrated into the PET scanner; only a few simple moving parts may be required to accommodate the gamma source; potentially no CT scanner is required; and nothing blocks or limits the diameter of the PET FOV.

[0055] In carefully designed implementations, the source is provided with a collimating shield so that gamma rays directed towards the detector are emitted at full intensity, while gamma rays directed in other directions are absorbed by the shield, reducing radiation exposure to people near the scanner.

[0056] However, there are reasons to suspect that the speedup or sensitivity gain predicted by equation (4) underestimates the benefit of the forward scatter approach. First, the scan 1 sinogram used in this analysis is 176 It is a combination of 202 and 307 keV gamma rays emitted by Lu, but experience has shown that it is better to use only 307 keV gamma rays. Secondly, the forward scattered gamma rays from cesium have mostly higher energy than those of lutetium. This can be seen in part 4B of Figure 6. The higher energy has the advantage of lower attenuation as the radiation penetrates the phantom and the patient's body. Also, less scattering is expected to result in fewer complications due to backscattering of 511 keV radiation from the patient.

[0057] The use of forward scatter radiation is influenced by physical effects at the lower end of the energy range. This is shown in part (B) of Figure 8. The angular distribution of Compton scatter at 662 keV favors small angles, but scattering at the smallest angles leaves little energy in the scintillator crystal, which may be below the low energy threshold of the scanner. As a result, much of the scattered radiation cannot be detected. As shown in Figure 1C, the use of multiple cesium sources results in a sinogram without well-defined edges. This is illustrated by Figure 10, where 19 source positions are computationally combined.

[0058] To describe the relative efficiency of the two methods, the LSO background and the cesium forward scattering, we introduce a figure of merit called the coincidence efficiency. Equation (5) and Table 1 show that 176 Lu beta decay is detected in nearly 100% of all decays, and two gamma rays are emitted in each decay, which shows that the LSO background radiation can be used efficiently. In the experiment, the cesium source was placed close to the ring of PET detectors, so that the solid angle was 6.3% of 4π steradians. Taking all factors into account, the forward scattering method is 3.72 × 10 -3 is significantly more efficient than -3It was shown that. Three effects are noted to explain why some coincidences that might have been detected were not actually seen. First, some of the 662 keV gamma rays may have passed through the detector without any interaction at all, or may have been scattered or absorbed by material between the source and the scintillator. These effects are estimated to account for about 40% of the loss for the detector block closest to the cesium source. Second, some of the gamma rays may have entered the detector and scattered at angles less than 40 degrees, leaving an amount of energy that was reduced below the detection threshold. This effect is another factor based on a simple analysis of the Compton scattering formula with weighting by the sine of the scattering angle, which is estimated to reduce the detection rate by about 40%. These two effects alone explain the majority of the missed coincidences. And the third effect is that after scattering, most of the photons occurred on the opposite side of the FOV, not facing the detector. This overall geometric efficiency factor would have been similar for the LSO background radiation, so the ratio of the two coincidence efficiencies in Table 1 is not easily predicted. In summary, this ratio is close to what would be expected based on simple calculations.

[0059] To use forward scatter in a practical way in a PET scanner, there is a trade-off between the real or perceived radiation hazard and the quality of the transmission image. Thus comes the determination of the optimal amount of gamma source material. In the example using Cs-137, 38 sources with As=30MBq per source can be considered, for a total of 1110MBq, which exactly matches the total Cs-137 content used in the ECAT ART PET scanner. Equation (4) suggests that the sensitivity gain in this case is 6.5.

[0060] This experiment verified that gamma forward scattering, using 662 keV gamma rays from Cs-137 as an example, can be used for transmission imaging in a PET scanner. In this application, the energy window and the coincidence time window need to be opened wide. The resulting sinograms can also be added to the sinograms of the LSO background radiation. If such a PET scanner is built with a total of 1110 MBq distributed among 38 cesium sources, the transmission images during a typical 3 min PET acquisition can be of similar quality to a 20 min scan based only on the LSO background.

[0061] The photon energy E detected from each coincidence event A and B from forward scattered gamma photons A and E B is the initial energy E of the gamma photon emitted from the gamma ray source 130 I The principle of addition to E A +E B =E I can be used to make the PET system reject almost all background noise, such as false coincidence events. For any gamma source, the E of the gamma rays from that source can be calculated by I is known. For example, in the case of Cs-137, E I is 662 keV. Therefore, if Cs-137 is used as the gamma ray source 130, requirement E A +E B =E I = 662 keV can be used to screen out true coincidence events arising from forward scattered Cs-137 gamma photons and improve the quality of the identified transmission scan signal.

[0062] For example, in the method summarized in flow chart 300 of FIG. 4A, step (b) satisfies requirement E to identify forward scattered gamma photons. A +E B =E I where E A and EB are Compton scattering in the scintillator crystals in the first set of detector blocks, and event B is Compton scattering in the scintillator crystals in the second set of detector blocks, respectively; E I is the initial energy of the gamma photon emitted by the gamma ray source.

[0063] In another example, in the method summarized in flowchart 400 of FIG. 4B, step (e) may include determining whether a forward scattered gamma photon is to be identified based on requirement E A +E B =E I where E A and E B are the detected photon energies from each pair of coincidence events A and B, respectively, where event A is the Compton scattering in the first scintillator crystal and event B is the Compton scattering in the second scintillator crystal, and E I is the initial energy of the gamma photon emitted by the gamma ray source.

[0064] In some embodiments where the scintillator crystal of the PET detector is an LSO crystal, the LSO crystal background radiation may also be used as an emission scan to generate an attenuation map according to the methods previously disclosed in U.S. patent application Ser. No. 14 / 172,980, filed Feb. 5, 2014, the contents of which are incorporated herein by reference.

[0065] In accordance with the above disclosure, there is provided a positron emission tomography (PET) scanner system including a gantry; a plurality of detector ring assemblies disposed within the gantry, each detector ring assembly comprising a plurality of PET detectors arranged in a ring shape around a central opening, the plurality of PET detector ring assemblies being coaxially arranged along a longitudinal axis defined through the central opening; a patient tunnel extending through the central opening of the coaxially arranged PET detector ring assemblies, the plurality of PET detector assemblies being coaxially arranged along a length of the patient tunnel, each of the plurality of PET detectors including a detector and one or more scintillator crystals associated with the detector; and one or more fixed gamma ray sources disposed in each of the PET detector ring assemblies within the gantry.

[0066] In a PET scanner system according to any of the above-described embodiments, the one or more fixed gamma radiation sources in each of the PET detector ring assemblies may be located rearward of the detector ring assemblies away from the patient tunnel. In a PET scanner system according to any of the above-described embodiments, the fixed gamma radiation sources may be Cs-137, cobalt-60, or sodium-22. In some embodiments of the PET scanner system, the fixed gamma radiation source is Cs-137.

[0067] In any of the above embodiments of the PET scanner system, each of the gamma radiation sources may be provided in an assembly that provides radiation shielding.

[0068] In any of the above embodiments of the PET scanner system, the assembly providing radiation shielding can include an ON configuration and an OFF configuration, and in the ON configuration, the remaining gamma photons from the Cs-137, other than those from the Cs-137 traveling toward the detector assembly, are substantially absorbed by the assembly providing radiation shielding. In any of the above embodiments of the PET scanner system, the assembly providing radiation shielding is at least partially constructed of a material including tungsten.

[0069] In any of the above embodiments of the PET scanner system, the PET scanner may be a PET / MR scanner. In any of the above embodiments of the PET scanner system, the PET scanner may be a PET / CT scanner. The subject matter has been described in terms of exemplary embodiments, but is not limited thereto. Rather, the appended claims should be construed broadly to include other modifications and embodiments that may be made by those skilled in the art. The present disclosure may further include the following configuration: . (Configuration 1) Gantry; a plurality of PET detector ring assemblies disposed within the gantry, each detector ring assembly including a plurality of PET detectors arranged in a ring shape around a central opening, the plurality of PET detector ring assemblies being coaxially arranged along a longitudinal axis defined through the central opening; a patient tunnel extending through the central opening of the coaxially arranged PET detector ring assemblies, where the plurality of PET detector assemblies are coaxially arranged along a length of the patient tunnel, each of the plurality of detectors comprising a detector and one or more scintillator crystals associated with the detector; and one or more fixed gamma radiation sources disposed within the gantry in each of the PET detector ring assemblies; A positron emission tomography (PET) scanner system comprising: . (Configuration 2) 2. The PET scanner system of claim 1, wherein the one or more fixed gamma radiation sources in each of the PET detector ring assemblies are positioned rearwardly of the detector ring assembly and away from the patient tunnel. . (Configuration 2-1) A plurality of PET detectors arranged in a ring shape; one or more fixed gamma ray sources disposed radially outside the ring shape relative to the plurality of PET detectors; A ring assembly comprising: . (Configuration 3) 3. The PET scanner system of claim 2, wherein the fixed gamma ray source is Cs-137, cobalt-60, or sodium-22. . (Configuration 4) 3. The PET scanner system of claim 2, wherein the fixed gamma ray source is Cs-137. . (Configuration 5) 3. The PET scanner system of claim 2, wherein each of the gamma ray sources is mounted within an assembly that provides radiation shielding. . (Configuration 6) 4. The PET scanner system of claim 3, wherein each of the gamma ray sources is mounted within an assembly that provides radiation shielding. 。 (Configuration 7) the assembly providing the radiation shielding has an ON configuration and an OFF configuration; 6. The PET scanner system of claim 5, wherein when in the ON configuration, the remaining gamma photons from the Cs-137, other than gamma photons from the Cs-137 traveling in a direction toward the detector assembly, are substantially absorbed by the radiation shielding assembly. 。 (Configuration 8) the assembly providing the radiation shielding has an ON configuration and an OFF configuration; 7. The PET scanner system of claim 6, wherein when in the ON configuration, the remaining gamma photons from the Cs-137, other than gamma photons from the Cs-137 traveling in a direction toward the detector assembly, are substantially absorbed by the radiation shielding assembly. 。 (Configuration 9) 8. The PET scanner system of claim 7, wherein the assembly providing radiation shielding comprises at least in part tungsten. 。 (Configuration 10) 9. The PET scanner system of claim 8, wherein the assembly providing radiation shielding comprises at least in part tungsten. 。 (Configuration 11) 2. The PET scanner system of claim 1, wherein the PET scanner is a PET / MR scanner. 。 (Configuration 12) 2. The PET scanner system of claim 1, wherein the PET scanner is a PET / CT scanner. 。 (Configuration 13) 1. A method of using one or more fixed gamma radiation sources in a PET scanner as a radiation source for a transmission scan to generate scan data that can be used to generate an attenuation map, the PET scanner comprising a plurality of PET detector ring assemblies, the method comprising the steps of: (a) providing one or more fixed gamma radiation sources for each of the PET detector ring assemblies, wherein the fixed gamma radiation sources are disposed outside the detector ring assemblies; (b) identifying, via coincidence electronics of the PET scanner, forward scattered gamma photons emitted from the fixed gamma ray source that scattered forward through scintillator crystals in a first set of detector blocks in a PET detector ring assembly, traverse a field of view (FOV) of the PET scanner, and are detected by scintillator crystals in a second set of detector blocks in the PET detector ring assembly; (c) acquiring list-mode data from a blank transmission scan with no radioactivity within the FOV; (d) acquiring list-mode data from a transmission scan with the target object within the FOV; (e) generating an attenuation map by comparing the list-mode data from the blank transmission scan with list-mode data from the transmission scan with the target object within the FOV; and (f) applying the attenuation map to the list-mode data from the emission scan of step (e) to apply attenuation correction to the list-mode data from the emission scan; A method comprising: 。 (Configuration 14) 14. The method of claim 13, further comprising, concurrently with step (d), acquiring list-mode data from a PET scan with the target object in the FOV. 。 (Configuration 15) The step (b) satisfies requirement E. A +E B =E I identifying the forward scattered gamma photons using Here, E A and E B are the detected photon energies from each pair of coincidence events A and B, respectively, the event A is a Compton scattering in a scintillator crystal in the first set of detector blocks; the event B is a Compton scattering in a scintillator crystal in the second set of detector blocks; E I is the initial energy of a gamma photon emitted from the gamma ray source; The method according to claim 13 。 (Configuration 16) 1. A method for improving transmission scan signal quality from forward scattered gamma photons by applying time-of-flight (TOF) considerations to a transmission imaging scan using a fixed gamma ray source, comprising: (a) calculating a TOF of a scattered photon, the photon being a photon from the gamma ray source that has undergone Compton scattering in the first scintillator crystal of the first detector block, to reach a second scintillator crystal of the second detector block based on a distance between the two scintillator crystals; (b) defining a time window, where the time window has a width centered around the calculated TOF; (c) measuring the TOF of actual scattered gamma photons originating from the first scintillator crystal while a scanned target object is in a field of view (FOV) of the PET scanner; (d) comparing the measured TOFs from (c) to the calculated TOFs to identify the measured TOFs that are within the time window; (e) identifying the forward scattered gamma photons corresponding to the measured TOF within the time window as transmission source events arising from Compton scattering in the first scintillator crystal, thereby distinguishing transmission type data from gamma emitting annihilation events and chance events within the scanned target; A method comprising: 。 (Configuration 17) 17. The method of claim 16, wherein step (c) further comprises simultaneously acquiring transmission scan data of the scan target object in the FOV of the PET scanner. 。 (Configuration 18) acquiring PET emission scan data of the scan target object in the FOV of the PET scanner simultaneously with step (c); and generating an attenuation map from the transmission scan data for correcting the PET emission scan data; 18. The method of claim 17, further comprising: 。 (Configuration 19) The step (e) satisfies the requirement E A +E B =E I identifying forward scattered gamma photons using Here, E A and E B are the photon beams detected from each pair of coincidence events A and B, respectively, the event A being the Compton scattering in the first scintillator crystal, and the event B being the Compton scattering in the second scintillator crystal; The above E I is the initial energy of the gamma photon emitted from the gamma ray source, The method according to claim 16, 。

Claims

1. Gantry; a plurality of PET detector ring assemblies disposed within the gantry, each PET detector ring assembly including a plurality of PET detectors arranged in a ring shape around a central opening, the plurality of PET detector ring assemblies being coaxially arranged along a longitudinal axis defined to pass through the central opening; a patient tunnel extending through the central opening of the plurality of coaxially arranged PET detector ring assemblies, the plurality of PET detector ring assemblies being coaxially arranged along a length of the patient tunnel, each of the plurality of PET detectors comprising a detector and one or more scintillator crystals associated with the detector; and one or more fixed gamma radiation sources disposed in each of the plurality of PET detector ring assemblies in the gantry, each fixed gamma radiation source configured to emit forward scattered gamma photons through at least a portion of the one or more scintillator crystals included in a corresponding one of the PET detector ring assemblies toward a field of view (FOV) in the patient tunnel; A positron emission tomography (PET) scanner system comprising:

2. 2. The PET scanner system of claim 1, wherein the one or more fixed gamma radiation sources in each of the PET detector ring assemblies are disposed posteriorly away from the patient tunnel relative to the PET detector ring assembly.

3. The PET scanner system of claim 2 , wherein the one or more fixed gamma ray sources are Cs-137, Cobalt-60, or Sodium-22.

4. The PET scanner system of claim 2 , wherein said one or more fixed gamma ray sources are Cs-137.

5. The PET scanner system of claim 2 , wherein each of the one or more fixed gamma radiation sources is mounted within an assembly that provides radiation shielding.

6. The PET scanner system of claim 3 , wherein each of the one or more fixed gamma radiation sources is mounted within an assembly that provides radiation shielding.

7. the assembly providing the radiation shielding has an ON configuration and an OFF configuration; 6. The PET scanner system of claim 5, wherein when in the ON configuration, remaining gamma photons from the fixed gamma ray source other than gamma photons from the fixed gamma ray source traveling in a direction toward the PET detector ring assembly are absorbed by the radiation shielding assembly.

8. the assembly providing the radiation shielding has an ON configuration and an OFF configuration; 7. The PET scanner system of claim 6, wherein when in the ON configuration, remaining gamma photons from the Cs-137 other than gamma photons from the Cs-137 traveling in a direction toward the PET detector ring assembly are absorbed by the radiation shielding assembly.

9. The PET scanner system of claim 7 , wherein the assembly providing radiation shielding comprises at least partially tungsten.

10. The PET scanner system of claim 8 , wherein the assembly providing radiation shielding comprises at least partially tungsten.

11. 10. The PET scanner system of claim 1 which is a PET / MR scanner.

12. 10. The PET scanner system of claim 1 which is a PET / CT scanner.

13. 1. A method of using one or more fixed gamma radiation sources in a PET scanner as a radiation source for a transmission scan to generate scan data that can be used to generate an attenuation map, the PET scanner comprising a plurality of PET detector ring assemblies, the method comprising the steps of: (a) providing the one or more fixed gamma radiation sources in each of the PET detector ring assemblies, wherein the fixed gamma radiation sources are disposed outside the PET detector ring assemblies; (b) identifying, via coincidence electronics of the PET scanner, gamma photons emitted from the fixed gamma ray source that are scattered forward through scintillator crystals in a first set of detector blocks in a PET detector ring assembly, traverse a field of view (FOV) of the PET scanner, and are detected by scintillator crystals in a second set of detector blocks in the PET detector ring assembly; (c) acquiring list-mode data from a blank transmission scan with no radioactivity within the FOV; (d) acquiring list-mode data from a transmission scan with the target object within the FOV; (e) generating an attenuation map by comparing the list-mode data from the blank transmission scan with list-mode data from the transmission scan with the target object within the FOV; and (f) applying the attenuation map to the list-mode data from the emission scan of step (e) to apply attenuation correction to the list-mode data from the emission scan; A method comprising:

14. 14. The method of claim 13, further comprising acquiring list-mode data from a PET scan with the target object in the FOV, simultaneously with step (d).

15. The step (b) satisfies requirement E. A +E B = E I identifying the forward scattered gamma photons using Here, E A and E B are the photon energies detected from each pair of coincidence events A and B, respectively; the event A is a Compton scattering in a scintillator crystal in the first set of detector blocks; the event B is a Compton scattering in a scintillator crystal in the second set of detector blocks; E I is the initial energy of a gamma photon emitted from the fixed gamma ray source; The method of claim 13.

16. 1. A method for improving transmission scan signal quality from forward scattered gamma photons by applying time-of-flight (TOF) considerations to a transmission imaging scan using a fixed gamma source, comprising: (a) calculating a TOF of a scattered photon, the photon being a photon from the fixed gamma ray source that has undergone Compton scattering in a first scintillator crystal of a first detector block, to reach a second scintillator crystal of a second detector block based on a distance between the two scintillator crystals; (b) defining a time window, where the time window has a width centered around the calculated TOF; (c) measuring the TOF of actual scattered gamma photons originating from the first scintillator crystal with the scanned target object in a field of view (FOV) of a PET scanner; (d) comparing the measured TOFs from (c) to the calculated TOFs and identifying the measured TOFs that are within the time window; (e) identifying the forward scattered gamma photons corresponding to the measured TOFs that fall within the time window as transmission source events arising from Compton scattering in the first scintillator crystal, thereby distinguishing transmission type data from gamma emitting annihilation events and chance events within the scanned target; A method comprising:

17. 17. The method of claim 16, wherein step (c) further comprises simultaneously acquiring transmission scan data of the scan target object in the FOV of the PET scanner.

18. acquiring PET emission scan data of the scan target object in the FOV of the PET scanner simultaneously with step (c); and generating an attenuation map from the transmission scan data for correcting the PET emission scan data; 20. The method of claim 17, further comprising:

19. The step (e) satisfies requirement E. A +E B = E I identifying forward scattered gamma photons using Here, E A and E B are the photon beams detected from each pair of coincidence events A and B, respectively, the event A being the Compton scattering in the first scintillator crystal, and the event B being the Compton scattering in the second scintillator crystal; The above E I is the initial energy of a gamma photon emitted from the fixed gamma ray source; 17. The method of claim 16.

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