SCINTILLATION CRYSTAL ASSEMBLY WITH REDUCED DETECTION OF SCATTER RADIATION

Abstract: A scintillation crystal (20) of a scintillation camera (10) is mounted on a first side of a light guide (18). Light detectors (22) are mounted on a second side of the light guide (18). The light guide (18) includes a chemical element having an atomic number greater than or equal to 40. Preferably, the chemical element is a glass material including oxides of at least one of barium, lanthanum, and lead. Gamma rays in the 70 to 600 keV range are photo electrically absorbed in the scintillation crystal resulting in the emission of light. A significant portion of the gamma rays, particularly gamma rays over 300 keV pass through the scintillation crystal and interact with the light guide. When the light guide has a 7 >/= 40, the primary interaction is photo electric absorption. When the light guide has a low-Z, a significant portion of the gamma rays are Compton scattered, some back into the scintillation crystal causing detrimental secondary scintillation events.

Description: SCINTILLATION CRYSTAL ASSEMBLY WITH REDUCED DETECTION OF SCATTER RADIATION Background of the Invention The present invention relates to the art of nuclear medicine and diagnostic imaging. It finds particular application in reducing errant scintillations produced by back scattered rays from reaching an associated scintillation crystal during positron emission tomography ("PET")- It is to be appreciated that the present invention may also be used in conjunction with single photon emission computed tomography ("SPECT"), whole body nuclear scans, transmission imaging, other diagnostic modes and/or other like applications. Heretofore, scintillation crystals are widely used for radiation detection in many industries. In nuclear medicine, large area crystals (e.g., crystal plates about 20 mm thick) of thallium-doped sodium iodide (Nal(Tl)) are most commonly used for gamma camera detectors. The crystals are formed on glass plates about 20 mm thick. The plates are typically a blend of silicon, sodium, boron, calcium, and aluminum oxides selected for their transparency and index of refraction. The other face of the crystal plate was commonly encased with an aluminum sheet to form a hermetic seal. In the camera, the glass plate functions as a light guide between the scintillation crystal and photomultiplier tubes or other photo detectors. The light guide optically spreads the scintillation light pulse to an optimal size. The glass material is chosen by considering its physical properties. The index of refraction is selected to conform with the scintillation crystal. The optical absorption or transparency at the wavelength of the emitted scintillations is minimized. Other properties of the glass are also of interest, such as the strength and thermal expansion properties along with the cost of manufacturing. At energy levels around 80 or 144 keV, substantially all of the incident energy is absorbed by the scintillation crystal and converted into light. Recently, it has become desirable to perform PET scans in addition to traditional SPECT and planar single photon scans. In the PET mode, the camera is subjected to radiation energy levels about 511 keV at very high event rates, often over 1 million counts per second ("Mcps"), to maximize the number of coincidence counts and to optimize image quality. The present inventor has found that at 511 keV, the about 20 mm sodium iodide crystal only absorbs and converts about 12% of the radiation into light. The remaining radiation passes through the crystal into the glass plate with a portion being Compton scattered in the crystal. Much of the radiation that reaches the glass plate is Compton scattered in the plate. Radiation that is back scattered, i.e. scattered back into the scintillation crystal, typically gives up a little over half of its energy to the scattering interaction, depending on the scattering angle. The radiation which is back scattered into the crystal along various angles has an energy of about 170 keV to about 300 keV. In this energy range, particularly at oblique angles, the back scattered radiation is substantially completely converted into light. The scintillations from the back scattered radiation are noise or errors which degrade the resultant image. The present invention provides a new and improved apparatus and method which overcomes the above-referenced problems and others . Summary of the Invention A scintillation camera includes a scintillation crystal mounted on a first side of a light guide. Light detectors are mounted on a second side of the light guide. The light guide includes a chemical element having an atomic number of greater than or equal to 40. A method of detecting gamma radiation includes receiving gamma rays with a scintillation crystal. A portion of the gamma rays are photoelectrically absorbed in the scintillation crystal and emit light. A portion of the gamma rays are passed through the scintillation crystal. The emitted light is passed through a light guide, which includes a chemical element having an atomic number of greater than or equal to 40, to photo detectors. Photoelectric absorption of the portion of the gamma rays that pass through the scintillation crystal in the light guide is maximized. Compton scattering of the gamma rays in the light guide is minimized. A primary advantage of the present invention is that it reduces or eliminates stray radiation events from back scattered radiation. Another advantage of the present invention is that the count rates are increased. Another advantage of the present invention is that if more than one energy window is used, scattered gamma rays contaminating the lower energy window from the higher energy window are reduced. Still further advantages of the present invention will become apparent to those of ordinary skill in the art upon reading and understanding the following detailed description of the preferred embodiments. Brief Description of the Drawings The invention may take form in various components and arrangements of components, and in various steps and arrangements of steps. The drawings are only for purposes of illustrating a preferred embodiment and are not to be construed as limiting the invention. FIGURE 1 illustrates a diagrammatic illustration of a nuclear camera system according to the present invention; FIGURE 2 is a diagrammatic illustration of gamma ray interactions within the scintillation crystal and the light guide of FIGURE 1; FIGURE 3 illustrates a histogram of an energy spectrum of Nal for detected coincidence events achieved in a computer simulation using standard low-Z glass; and FIGURE 4 illustrates a histogram of an energy spectrum of Nal for detected coincidence events achieved in a computer simulation using high-Z glasses. Detailed Description of the Preferred Embodiments With reference to FIGURE 1, a nuclear camera system 10 includes a plurality of detectors heads ("detectors") 12 mounted for movement around a subject 14 in an examination region 16. Each of the detectors 12 includes a light guide 18 and a scintillation crystal 20 that converts a radiation event into a flash of light energy or scintillation. An array of sensors 22, e.g. 59 sensors, is arranged to receive the light flashes from the scintillation crystal 20. In the preferred embodiment, the sensors include photomultiplier tubes. However, other sensors are also contemplated. Each of the sensors 22 generates a respective analog sensor output pulse (e.g., tube output pulse) in response to the received light flash. Furthermore, each of the sensors 22 is electrically connected to analog-to-digital converters 24. The analog-to-digital converters 24 convert the analog sensor output pulses to a series of digital sensor output values. A processor 26 determines coordinates in two dimensions of the location and the energy of the scintillation event that occurred in the crystal. In a PET mode, a coincidence detector 28 determines when scintillations are detected concurrently in both scintillation crystals. When coincident events are detected, the spatial coordinates of the events, which define a connecting ray, are communicated to a reconstruction processor 30. An angular orientation of the detector heads 12 around the subject is determined by an angular position encoder 32. The reconstruction processor 30 reconstructs a volumetric image representation from the rays passed by the coincidence detector. The volumetric image representation is stored in a volume image memory 34. A video processor 36 converts operator selected portions of the volumetric image presentation into an appropriate format for display on a monitor 38, such as a video monitor, CCD monitor, active matrix monitor, high resolution printer, or the like. Typical human readable displays include slice images, volume or surface renderings, projection images, and the like. In the preferred embodiment, the light guide 18 includes a glass material having a chemical element with an atomic number ≥ ~40. Preferably, the light guide 18 includes oxides of barium, lanthanum, and/or lead. The light guide optically spreads light pulses from the scintillation crystal 20 to a predetermined size. Alternatively, the guide 18 includes a transparent plastic material in which atoms with an atomic number ≥ ~40 are bonded into the polymer chain. The chemical composition of a glass light guide having an atomic number ≥ ~40 causes the detection of scattered high energy photons to be minimized. For the case of a gamma camera detecting 511 keV gamma energy rays, the use of high-Z glass is beneficial when considering the properties of sodium iodide crystal and commercially available glasses. With reference to FIGURE 2, the two (2) dominant processes through which photons in the energy range of nuclear medicine (e.g., between ~70 keV and ~600 keV) interact with matter are photoelectric absorption and Compton scattering. In photoelectric absorption, a gamma ray 60 strikes the nucleus 62 in the crystal and is completely absorbed such that its energy transferred to one or more electrons (s) , with the occasional emission of an x-ray. The electrons change quantum levels emitting light 64 of a characteristic wavelength. The light guide passes the light, with an intensity vs. spatial location represented by curve 66. In Compton scattering, a gamma ray 70 strikes a nucleus 72 in the crystal or the light guide and its energy is partially transferred, resulting in an energetic electron and a scattered gamma ray 74 of lower energy, which may itself be photoelectrically absorbed or Compton scattered. With the prior art low-Z glasses, the scattered gamma ray still has sufficient energy, e.g. about 200 keV to undergo photo electric absorption when it strikes a nucleus 76 in crystal causing the emission of light 78. A resultant intensity vs. spatial location curve 80 is spatially shifted relative to the true entry point of the gamma ray 70 into the crystal, causing inaccuracies in the resultant image. The desired scenario for nuclear imaging is complete absorption of gamma rays within the scintillation crystal itself. Such complete absorption triggers a primary scintillation event. At lower energies (e.g., ~70 keV to ~80 keV (Tl-201 gamma rays) and ~140 keV (Tc-99m gamma rays) ) , a majority of the gamma rays are absorbed because the Nal (TI) crystal has a very high stopping power at these energies. However, at higher energies (e.g., ~511 keV (positron annihilation gamma rays in PET imaging) ) , the stopping power of Nal (TI) is much smaller. The dominant process at 511 keV in a Nal(Tl) crystal of typical thickness (between about 10 mm and about 20 mm) is not complete absorption. A gamma ray 90 partially interacts 92 with the crystal, depositing less than ~511 keV in the crystal. A small amount of light 94 is generated as the gamma ray defects and continues on 90' at lower energy. These events, referred to as secondary events, are eliminated from processing by energy discrimination, i.e. its intensity curve 96 is so low it is easily filtered. However, these secondary events can be detrimental to PET imaging on such a system. To assume that only one event is detected at a time, when an event is detected that produces scintillation light, the detection of additional events is blocked for a short time consistent with the duration of a scintillation event. This contributes to the dead time of the detector and helps to limit the count rate capacity of the camera. The majority of 511 keV events, however, pass through the crystal without interacting at all, note gamma ray 70. Aside from a loss of sensitivity, these events would seem to pose no harm to the count rate capacity of the camera, since no interaction had occurred. However, after passing through the crystal, there is a high probability of interacting with other components of the detector (i.e., the glass, the photo detectors, the electronics, the camera "bucket," etc) . Because Compton scattering is by far the dominant mode of interaction with most material at ~511 keV, the issue is complicated even more. Therefore, many gamma rays are scattered and redirected towards the crystal like gamma ray 70. The energy of the scattered gamma rays is lower, typically between ~170 keV and ~300 keV. Even at this reduced energy, there is a high probability that the scattered gamma rays will reach the scintillation crystal and trigger electronics for contributing to the dead time, and limiting the count rate or generating false data. There are other ways for undesired scattered photons to reach the scintillation crystal. X-rays or bremsstrahlung x-rays also may be produced in the components behind the crystal, and these may also be directed towards and detected in the crystal. A large portion of these "backscattered" rays, which produce the secondary events, are prevented from reaching the detector by the present chemical composition of the glass light guide 22 which include materials having high atomic numbers (i.e., high-Z materials). As discussed above, reducing the backscattered rays helps avoid secondary scintillation events, also referred to as pulse pile-up. Although glass light guides having high-Z materials do not completely eliminate the probability of Compton scattering at 511 keV, the probability of photoelectric absorption at the scattered energies is much larger. Typical glasses contain mostly silicon dioxide (Si0 2 ) and other oxides of low-Z elements (e.g., aluminum, sodium, boron, and calcium). Glass of this type is used in gamma cameras. Heavier glasses exist, such as, for example, flint glass (containing lead oxide (PbO) ) , some crown glasses (containing barium oxide (BaO) , and rare earth glasses (containing lanthanum oxide La 2 0 3 ) . Usually, the optical properties of heavier glasses are not as desirable for gamma cameras relative to the lighter glasses. For example, they tend to be less transparent to light from the scintillations. Therefore, such heavier glasses are not typically considered as candidates for light guides in gamma cameras. However, the heavier glasses are much more opaque to the scattered gamma rays. High-Z materials are desirable because photoelectric absorption dominates over Compton scattering at the scattered energies E. The cross-section for photoelectric absorption scales as Z 3 /E 3 . The cross-section for Compton scattering is relatively independent of energy and depends mainly on the mass density of the material. There is a crossover photon energy above which Compton scattering dominates and above which the ability to shield the crystal from scattered rays diminishes. For silicon (Z = 14), zinc (Z = 30), barium (Z = 56), and lead (Z = 82) , the crossover energies are ~30 keV, ~140 keV, -290 keV, and -560 keV, respectively. Therefore, in order to prevent Compton scattering from dominating and to adequately shield the crystal from scattered rays at -511 keV, the glass light guide should contain chemical elements having an atomic number Z > -40. For this reason, glasses containing barium (Z = 56), lanthanum (Z = 57), and lead (Z = 82) oxides are preferred for reducing Compton scattering and adequately shielding the crystal from scattered rays at -511 keV. The number of scattered gamma rays that are detected, therefore, are reduced by using glass containing high-Z elements. The potential benefit is that the maximum count rate is increased, thereby improving performance for coincidence imaging and other studies having very high event rates (e.g, first-pass cardiac imaging) . Furthermore, in multiple energy studies, higher-Z glass light guides reduce the amount of contamination from the scatter behind the scintillation crystal and improve the quality of the image. For example, if dual isotope imaging is performed with 18 F-FDG (511 keV) and 99m Tc-mibi (140 keV) , scatter from the 511 keV gamma rays in the 140 keV energy window is substantial. It is to be understood that using the higher-Z glass is also beneficial for such dual isotope imaging. FIGURE 3 illustrates a histogram of an energy spectrum of Nal for detected coincidence events achieved in a computer simulation using standard low-Z glass. A typical gamma camera detector was modeled as a 10 mm slab of Nal crystal, a 15 mm slab of glass as the light guide, a 10 mm slab of copper to simulate the other elements in the detector (magnetic shielding, photomultiplier tube electrodes, electronics, mounting plates, etc.), and a 12.5 mm slab of lead to simulate the detector bucket. The simulation investigated the case of a dual-head camera system in PET mode, which detects coincident pairs of gamma rays at 511 keV emitted from a scattering medium. Notably, a prominent backscatter peak 100 occurs at -180 keV in addition to a photoelectric peak 102. FIGURE 4 illustrates histograms of an energy spectrum of Nal for detected coincidence events achieved in a computer simulation using standard glass (indicated as the graph with a backscatter peak 110) along with typical compositions of barium crown glass (indicated as the graph with a backscatter peak 112) and light flint glass (indicated as the graph with a backscatter peak 114) . The barium crown glass and light flint glass both contain lead and high-Z glasses. Notably, the intensity of the backscatter peaks 112, 114 are significantly less for the high-Z barium crown glass and light flint glass, respectively, relative to the peak 110 for the standard, low-Z glass. A reduced backscatter peak indicates fewer scattered events are detected, minimizing the contribution of the undesired events to the dead time of the camera. Furthermore, the number of scattered events at -140 keV is much lower, which is helpful for 18 F/" ra Tc dual isotope imaging. In an alternate embodiment, it is contemplated that the photomultiplier tubes be manufactured with high-Z glass instead of, or in addition to, the light guide including the high-Z glass.

Claims: Having thus described the preferred embodiment, the invention is now claimed to be: 1. A scintillation camera (10) including a scintillation crystal (20) mounted on a first side of a light guide (18) and light detectors (22) mounted on a second side of the light guide (18), the light guide (18) comprising: a chemical element having an atomic number of greater than or equal to 40. 2. The scintillation camera (10) as set forth in claim 1, wherein the light guide (18) includes a glass material . 3. The scintillation camera (10) as set forth in claim 2, wherein the glass material includes oxides of at least one of barium, lanthanum, and lead. 4. The scintillation camera (10) as set forth in claim 1, wherein the light guide (18) includes a plastic material . 5. The scintillation camera (10) as set forth in any one of preceding claims 1-4, wherein the light guide (18) optically spreads light pulses from the scintillation crystal (20) to a predetermined size. 6. The scintillation camera (10) as set forth in any one of preceding claims 1-5, wherein: the scintillation crystal (20) absorbs gamma rays (60, 70, 90) and emits light (64, 78, 94) in primary events; the light detectors (22) include photo detectors which convert the emitted light (64, 78, 94) into electrical signals; and the light guide (18) conveys the emitted light (64, 78, 94) from the scintillation crystal (20) to the photo detectors (22) and reduces secondary energy events in the scintillation crystal (20) that interfere with the primary events . 7. The scintillation camera (10) as set forth in any of preceding claims 1-6, wherein the light guide (18) interacts with gamma radiation below about 600 keV primarily by photoelectric absorption. 8. The scintillation camera (10) as set forth in any of preceding claims 1-7, wherein the light guide inhibits Compton scattering of gamma radiation below about 600 keV. 9. The scintillation camera (10) as set forth in any of preceding claims 1-8 further including: a second scintillation crystal disposed across an examination region (16) from the first scintillation crystal (20); photo detectors disposed to receive light from the second scintillation crystal; and a second light guide which conveys light from the second scintillation crystal to the photo detectors; a coincidence detector (28) for detecting coincident scintillations in the first (20) and second scintillation crystals; and a reconstruction processor (26) for reconstructing coincident scintillation information from the photo detectors into an image representation. 10. The scintillation camera (10) as set forth in claim 9, wherein the first and second scintillation crystals are sodium iodide crystals less than about 20 mm thick. 11. A method of detecting gamma radiation, the method comprising: receiving gamma rays (60, 70, 90) with a scintillation crystal (20) ; photoelectrically absorbing a portion of the gamma rays (60, 70, 90) in the scintillation crystal (20) and emitting light (64, 78, 94) and passing a portion of the gamma rays (60, 70, 90) through the scintillation crystal (20); passing the emitted light (64, 78, 94) through a light guide (18), which includes a chemical element having an atomic number of greater than or equal to 40, to photo detectors (22); and maximizing photoelectric absorption of the portion of the gamma rays (60, 70, 90) that pass through the scintillation crystal (20) in the light guide (18) and minimizing Compton scattering of the gamma rays (60, 70, 90) in the light guide (18) . 12. The method as set forth in claim 11, wherein the gamma rays (60, 70, 90) have an energy between about 70 and about 600 keV. 13. The method as set forth in any one of preceding claims 11 and 12, wherein the light guide is a glass material incorporating oxides of at least one of barium, lanthanum, and lead. 14. The method as set forth in any one of preceding claims 11-13, further including: photo electrically absorbing the gamma rays (60, 70, 90) with a second scintillation crystal and emitting light; passing the emitted light with a second ' light guide to photoelectric detectors; generating electrical signals with the photo detectors in response to received light; determining coincident scintillations in the first (20) and second scintillation crystals; and reconstructing the electrical signals from the photo detectors (22) responsive to the coincident radiation into an image representation.

Inventors: DIFILIPPO FRANK P, DIFILIPPO, FRANK, P.
Applicants: MARCONI MEDICAL SYS INC, MARCONI MEDICAL SYSTEMS, INC.
Application year: 2001
Main classification: G01T
IPC codes: G01T1