Abstract:
An improved scintillator detector cell geometry for converting radiation to light with improved light collection is provided. Shaping the exit face of a scintillator to increase the surface area of the exit face results in a decrease in a fraction of angles that undergo total internal reflection within the scintillator. The scintillator has the advantage of preventing total internal reflection parallel, as well as perpendicular, to the detecting surface of a light collection device. Further, providing a specular reflector on a hemispherical dome portion of the radiation detecting surface of the scintillator results in reduced bounce-off the specular reflector before light contacts the scintillator-photodiode interface. Furthermore, implementing a convex shape when coated with the specular reflector increases the fraction of light directed toward the photodiode compared to a plane surface parallel to the photodiode. The present invention further limits the amount of light that is trapped within the scintillator.

Description:
BACKGROUND OF INVENTION 
     The present invention relates generally to radiation detection and, more particularly, to an improved apparatus and method of light collection for use with a radiation emitting medical imaging scanner. 
     Typically, in computed tomography (CT) imaging systems, an x-ray source emits a fan-shaped beam toward an object, such as a patient. The beam, after being attenuated by the object, impinges upon an array of radiation detectors. The intensity of the attenuated beam of radiation received at the detector array is typically dependent upon the attenuation of the x-ray beam by the object. Each detector element of the detector array produces a separate electrical signal indicative of the attenuated beam received by each detector element. The electrical signals are transmitted to a data processing system for analysis which ultimately results in the formation of an image. 
     Generally, the x-ray source and the detector array are rotated about the gantry within an imaging plane and around the object. X-ray sources typically include x-ray tubes, which emit the x-ray beam at a focal point. X-ray detectors typically include a collimator for collimating x-ray beams received at the detector, a scintillator for converting x-rays to light energy adjacent the collimator, and photodiodes for receiving the light energy from the adjacent scintillator. 
     Typically, the scintillator or scintillation detector used with CT systems and other radiation emitting medical imaging scanners are coupled to a light collection device such as a photodiode using a transparent adhesive. The transparent adhesive, however, typically limits the angles at which light can exit the scintillator to those greater than the critical angle of the interface between the scintillator and the adhesive. Moreover, scintillators usually have very high indices of refraction due to the high density required to prevent radiation leakage. As a result, an appreciable index change at the scintillator exit face occurs with resulting total internal reflection of light striking the exit face at relatively shallow angles. 
     Further, these known systems typically implement a rectangular block shaped scintillator to stop the radiation and maximize the area exposed. While this rectangular shape accommodates an easy to fabricate area-filling shape with the appropriate thickness for optimum radiation stopping power, the rectangular shape has several disadvantages associated with light collection by a photodiode. 
     For example, parallel plane walls are typically implemented that are perpendicular to the light detector photodiode thereby causing light that is emitted nearly parallel to the light collector face to bounce repeatedly off the parallel walls either by specular reflection or by total internal reflection. Moreover, there is no means of directing light preferentially toward the light detector in these rectangular shaped scintillators. Diffused reflectors have been implemented to randomize the direction of light rays, but these diffused reflectors result in the need for thicker reflectors to provide sufficient scatter by light refraction. As a result, these thicker reflectors decrease the possible re-fraction of the scintillator. Additionally, these reflectors are also not opaque so an additional opaque component to stop light leakage between cells must be employed. Other scintillators have been designed using a defused opaque coating, but the increased area of such a coating and the opportunity for multiple bounces typically results in a considerable decrease and reflectivity compared to a specular coating. 
     Additional disadvantages often associated with these rectangular shaped scintillators include the reflection of light back into the scintillator when the light hits the exit surface of the scintillator at an angle greater than the critical angle of the scintillator-adhesive interface. This disadvantage, as well as the previously discussed disadvantages, increase the mean number of reflections that light undergoes before striking the detector. Simply, each reflection off the side reflector of the scintillators results in a loss from the absorbance of the reflector. 
     It would therefore be desirable to design an apparatus and method of light collection whereby light throughput of a scintillator is increased and light collection efficiency likewise improves. 
     BRIEF DESCRIPTION OF THE INVENTION 
     The present invention is directed to improved scintillator detector cell geometry overcoming the aforementioned drawbacks. Shaping the exit face of a scintillator to increase the surface area results in a decrease in a fraction of angles that undergo total internal reflection within the scintillator. By convexly shaping the exit face, the scintillator has the advantage of preventing total internal reflection parallel, as well as perpendicular, to the detecting surface of a light collection device. Further, providing a specular reflector on a hemispherical dome portion of the radiation detecting surface of the scintillator results in reduced reflection off the specular reflector before light contacts the scintillator-photodiode interface. Furthermore, implementing a convex shape that is coated with the specular reflector, increases the fraction of light directed toward the photodiode compared to a planar surface parallel to the photodiode. The present invention further limits the amount of light that is trapped within the scintillator. 
     Therefore, in accordance with one aspect of the present invention, a scintillation apparatus for use with a radiation emitting medical imaging scanner is provided. The scintillation apparatus includes an entrance face configured to receive radiation and an exit face having a tetrahedral shape and configured to discharge light. The scintillation apparatus further includes a plurality of plane walls extending from the entrance face to the exit face. 
     In accordance with another aspect of the present invention, a CT system includes a scintillator array having a plurality of scintillation cells. Each scintillation cell of the CT system has at least one of a non-planar radiation reception surface and a non-planar light emitting surface. The non-planar reception surface and the non-planar light emitting surface are symmetrically shaped with respect to one another. The CT system further includes a radiation projection source configured to project radiation toward the scintillator array and a photodiode array having a plurality of photodiodes. The photodiode array is optically coupled to the scintillator array to detect light output therefrom. The CT system further includes a gantry having an opening to receive a subject to be scanned. 
     In accordance with yet another aspect of the present invention, a radiation detector for use with the radiation emitting medical imaging scanner is provided. The radiation detector includes a means for detecting radiation as well as a means for converting the radiation to light energy. The radiation detector further includes a means for emitting light energy toward a light energy detector and a means for reducing light energy bounce off within the radiation detector. 
     In accordance with a further aspect of the present invention, a method of light collection from a scintillation detector of a radiation emitting medical imaging scanner includes directing radiation toward a scan subject and a scintillation detector. The method further includes receiving radiation attenuated by the scan subject and converting the attenuated radiation to light energy. The light energy is then admitted through a non-planar surface of the scintillation detector. The method further includes detecting the admitted light energy from the scintillation detector. 
     Various other features, objects and advantages of the present invention will be made apparent from the following detailed description and the drawings. 
    
    
     BRIEF DESCRIPTION OF DRAWINGS 
     The drawings illustrate one preferred embodiment presently contemplated for carrying out the invention. 
     In the drawings: 
     FIG. 1 is a pictorial view of a CT imaging system. 
     FIG. 2 is a block schematic diagram of the system illustrated in FIG.  1 . 
     FIG. 3 is a perspective view of a scintillation detection cell in accordance with one embodiment of the present invention. 
     FIG. 4 is a cross-sectional view of a top portion of the scintillation cell of FIG. 3 taken along line  4 — 4 . 
     FIG. 5 is a cross-sectional view of the scintillation cell of FIG. 3 taken along line  5 — 5 . 
     FIG. 6 is a vertical cross-sectional view of the scintillation cell of FIG. 3 taken along line  6 — 6 . 
     FIG. 7 is a cross-sectional view similar to that shown in FIG. 6 illustrating a same alternative embodiment or different alternative embodiment of the present invention. 
     FIG. 8 is a perspective view of a scintillation cell in accordance with another alternative embodiment of the present invention. 
     FIG. 9 is a cross-sectional view of the scintillation cell of FIG. 8 taken along Line  9 — 9 . 
     FIG. 10 is a perspective view of a portion of a scintillator array incorporating a plurality of the scintillation cells shown in FIG.  3 . 
    
    
     DETAILED DESCRIPTION 
     Referring to FIGS. 1 and 2, a computed tomography (CT) imaging system  10  is shown as including a gantry  12  representative of a “third generation” CT scanner. Gantry  12  has an x-ray source  14  that projects a beam of x-rays  16  toward a detector array  18  on the opposite side of the gantry  12 . Detector array  18  is formed by a plurality of detectors  20  which together sense the projected x-rays that pass through a medical patient  22 . Each detector  20  produces an electrical signal that represents the intensity of an impinging x-ray beam and hence the attenuated beam as it passes through the patient  22 . During a scan to acquire x-ray projection data, gantry  12  and the components mounted thereon rotate about a center of rotation  24 . 
     Rotation of gantry  12  and the operation of x-ray source  14  are governed by a control mechanism  26  of CT system  10 . Control mechanism  26  includes an x-ray controller  28  that provides power and timing signals to an x-ray source  14  and a gantry motor controller  30  that controls the rotational speed and position of gantry  12 . A data acquisition system (DAS)  32  in control mechanism  26  samples analog data from detectors  20  and converts the data to digital signals for subsequent processing. An image reconstructor  34  receives sampled and digitized x-ray data from DAS  32  and performs high speed reconstruction. The reconstructed image is applied as an input to a computer  36  which stores the image in a mass storage device  38 . 
     Computer  36  also receives commands and scanning parameters from an operator via console  40  that has a keyboard. An associated cathode ray tube display  42  allows the operator to observe the reconstructed image and other data from computer  36 . The operator supplied commands and parameters are used by computer  36  to provide control signals and information to DAS  32 , x-ray controller  28  and gantry motor controller  30 . In addition, computer  36  operates a table motor controller  44  which controls a motorized table  46  to position patient  22  and gantry  12 . Particularly, table  46  moves portions of patient  22  through a gantry opening  48 . 
     Referring now to FIG. 3, a scintillation cell  50  is shown in accordance with one embodiment of the present invention. Cell  50  includes an upper portion or radiation detection region  52 , a body or intermediate region  54 , and a lower portion or light emitting region  56 . In this embodiment, the radiation detection region  52  as well as the light emitting region  56  have a trigonal pyramidal shape. As a result, the body  54  of cell  50  has three cell walls  58 . Cell walls  58  interface with curvilinearly shaped entrance surface walls  60  at one end and interface with emitting surface walls  62  at an opposite end. In this embodiment, the scintillation cell  50  has a symmetrical shape. That is, radiation detection region  52  mirrors light emitting region  56 . 
     FIG. 3 illustrates one preferred embodiment of the present invention having a scintillation cell with only one of a trigonal pyramidal radiation detection region  52  and a trigonal pyramidal light emission region  56 . In this embodiment, the scintillation cell would have a projectionless surface at one end and a trigonal pyramidal shape at an opposite end. Further, the present invention is applicable with detection surface walls and emission surface walls that are not curvilinear in shape but rather linear. Additionally, the present invention contemplates a scintillation cell having more than three cell walls. For example, a scintillation cell with four cell walls as well as four radiation detection surface walls and four light emission surface walls are within the scope of the present invention. Also, the present invention contemplates, in an alternative embodiment, a convex or hemispherical shape for at least one of the radiation detection region and the light emission region of the scintillation cell. In this alternative embodiment, the convex radiation detection region and the convex light emission region would each have a circular cross-section. 
     Referring to FIG. 4, a cross-sectional view taken along line  4 — 4  of FIG. 3 illustrates the triangular cross-section of the preferred embodiment shown in FIG.  3 . FIG. 4 illustrates the triangular orientation of the interface between the radiation detection region  52  and the intermediate or body region  54 . As shown, three cell walls are shown. The cell walls  58  are formed of and enclose scintillation materialwhich, as indicated previously, is designed to convert high frequency electromagnetic energy such as x-rays or gamma rays to light. 
     FIG. 5 is a cross-sectional view of the scintillation cell shown in FIG. 3 taken along line  5 — 5  thereof. As shown and similar to FIG. 4, FIG. 5 illustrates the triangular orientation of the scintillation cell  50 . Three cell walls  58  formed of scintillation material intersect to form a triangular body and enclose scintillation material  64 . FIG. 5 also illustrates the position of a reflective coating  63  about the periphery of cell walls  58 . Interposed between the optically reflective layer  63  and the cell walls  58  is a dielectric layer  65 . The dielectric layer  65  is formed of a dielectric selected to have an optical index that is less than that of the material forming the scintillator. A number of dielectric materials such as clean air may be used if they have an optical index less than that of the scintillator material  64 . The dielectric layer  65  operates to reflect light photons generated by the scintillation material  64  back into the scintillator body thereby increasing light collection efficiency. The dielectric layer  65  operates to increase the number of light photons that eventually exit the light emitting region  56 , FIG.  5  and strike the photodetector  20 , FIG.  2 . 
     Still referring to FIG. 5, reflective coating  63  is disposed around the periphery of the dielectric layer  65 . Light photons that strike the interface between cell walls  58  and the dielectric layer  65  at greater than the critical angle for that interface will pass through the dielectric layer  65  and preferably strike reflective layer  63 . Reflective layer  63  will reflect the light photons back through the dielectric layer  65  and into the scintillation material  64 . In a preferred embodiment, the optically reflective layer  63  has a specular surface and includes a metal having a relatively high reflectance. 
     FIG. 6 is a cross-sectional view of the scintillation cell  50  taken along line  6 — 6  of FIG.  3 . As shown, scintillation cell  50  has a tetrahedral radiation detection region  52  and a tetrahedral light emitting region  56  with the body or intermediate region  54  therebetween. Each region  52 ,  54 ,  56  of cell  50  is formed of a solid scintillation material  64 . X-rays or gamma rays detected by the radiation surface as defined by walls  60  are converted to light by the scintillation material  64  in accordance with known conversion techniques. The light photons are eventually then admitted through the light emission region  56  and detected by a photodetector. Utilizing a tetrahedral or pyramidal shape at the detection region  52  creates a significant directional component within the scintillation cell  50 . Simply, the portion of the light photons generated within the scintillation cell  50  that are directed back toward the radiation detection surface  60  may be, in accordance with the present invention, reflected toward the light emission region  56  such that the number of reflections off the interior surfaces of the scintillation cell are reduced. That is, the present invention reduces the number of parallel plane surfaces within the scintillation cell. As a result, light emitted nearly perpendicular to the plane surfaces are less susceptible to being caught in an optical cavity within the scintillation cell. The pyramidal shape of the detection region  52  improves the light photons progress toward the light emission regionand further implementing a specular reflector  65  and a dielectric layer  63  along a periphery of the side walls  58  of the scintillation cell  50  also improves light direction toward the light emitting surface. 
     By shaping the light emitting region  56  in a tetrahedral or pyramidal shape increases the surface area of the light emitting region  56  but also decreases the fraction of angles that undergo total internal reflection within the scintillation cell  50 . Maximized light throughput is possible with a trigonal pyramidal shape and has the advantage of preventing total internal reflection parallel as well as perpendicular to the light emitting surface. While the light emitting region  56  may have any convex shape or series of shapes, all of which are within the scope of the present invention, the use of multiple protrusions, i.e., ripples, from the light emitting surface can be problematic since the light exiting one protrusion may be refracted back into the scintillator by a neighboring protrusion thereby jeopardizing the light collection efficiency of the scintillator cell and the photodetector. 
     FIG. 7 is a cross-sectional view similar to that shown in FIG. 6 showing one alternate embodiment of the present invention. In this embodiment, a reflective coating extends along a portion of the periphery of the scintillation cell  50 . Whereas in the embodiment of FIG. 6, the reflective coating  63  extended only along an outer periphery of the intermediate region  54 , in the embodiment of FIG. 7, the reflective coating  63  is implemented along the intermediate side walls  58  as well as the detection surface side walls  60  and the light emission surface side walls  62 . In another alternative embodiment, the reflective coating may be extended to align along a portion of the periphery (not shown) of the light emitting region. Coating a portion of the exterior of the scintillator cell  50  with a reflective coating  63  represents only one preferred embodiment of the present invention. That is, a reflective coating may be implemented along an outer periphery of the side walls of the scintillation cell  50  as shown in FIG. 6 or, in another alternative, the reflective coating  63  may be used along an outer periphery of one or both of radiation detection region or the light emission region  56 . 
     Referring to FIG. 8, another alternative embodiment of the present invention is shown, similar to that shown in FIG. 3, which will be described using like numerals and a parenthetical (a) in describing the features thereof. As shown, scintillation cell  50 ( a ) is a three-sided structure having three body side walls  58 ( a ) and a trigonal pyramidal upper region  52 ( a ) and a trigonal pyramidal lower region  56 ( a ). Scintillator  50 ( a ) is different from scintillator  50  shown in FIG. 3 in two respects. First, scintillation cell  50 ( a ) has a flat light emission surface  66 . The light emission surface  66  is flat to ensure a greater connection to a photodetector (not shown). Secondly, the scintillation cell  50 ( a ) includes a reflective coating  63 ( a ) along the entire periphery except for the flat light emission surface  66 . As indicated previously, the reflective coating layer  63 ( a ) improves the light collection within the scintillation cell by reducing light photon bounce-off within the scintillation cell. Further, the reflective coating  63 ( a ), such as a specular reflector, may also prevent radiation that penetrates between adjacent scintillation cells from reaching the photo detector. 
     A cross-sectional view taken along line  9 — 9  of FIG. 8 is shown in FIG.  9 . As may be readily seen, the scintillation cell  50 ( a ) has a conical radiation detection region  52 ( a ) as well as a planar-bottommed trigonal pyramidal light emission region  56 ( a ). As shown, the light emission surface  66  extends horizontally across a tapered region  68  of light emission region  56 ( a ). 
     FIG. 10 illustrates a portion of a scintillator array incorporating a plurality of the scintillation cell  50  shown in FIG.  3 . Scintillator array  70  includes a plurality of scintillation cells wherein each scintillation cell  50  has a trigonal pyramidal radiation detection region  52  and a trigonal pyramidal light emission region  56  and three side walls  58  of an intermediate region  54  therebetween. Radiation is detected by each scintillation cell  50  of scintillator array  70  and converted to light energy which is then detected by a photodiode array (not shown) coupled to the scintillator array  70 . Specifically, each photodiode (not shown) of the photodiode array is coupled to a corresponding scintillation cell  50  of the scintillator array  70 . As indicated previously, the light energy generated by each scintillation cell  50  and detected by each photodiode (not shown) is indicative of the x-rays attenuated by an imaging subject and detected by the scintillator array  70 . Each photodiode detects light and transmits an electrical signal to a data acquisition system, FIG. 2, for subsequent processing and image reconstruction. 
     Accordingly, in accordance with one embodiment of the present invention, a scintillation apparatus for use with a radiation emitting medical imaging scanner is provided. The scintillation apparatus includes an entrance, the entrance face configured to receive radiation, and an exit face having a tetrahedral shape and configured to discharge light. The scintillation apparatus further includes a plurality of plane walls extending from the entrance face to the exit face. 
     In accordance with another embodiment of the present invention, a CT system includes a scintillator array having a plurality of scintillation cells. Each scintillation cell of the CT system has at least one of a non-planar radiation reception surface and a non-planar light emitting surface. The CT system further includes a radiation projection source configured to project radiation toward the scintillator array and a photodiode array having a plurality of photodiodes. The photodiode array is optically coupled to the scintillator array to detect light output therefrom. The CT system further includes a gantry having an opening to receive a subject to be scanned. 
     In accordance with yet another embodiment of the present invention, a radiation detector for use with the radiation emitting medical imaging scanner is provided. The radiation detector includes a means for detecting radiation as well as a means for converting the radiation to light energy. The radiation detector further includes a means for emitting light energy toward a light energy detector and a means for reducing light energy bounce off within the scintillator. 
     The present invention has been described in terms of the preferred embodiment, and it is recognized that equivalents, alternatives, and modifications, aside from those expressly stated, are possible and within the scope of the appending claims.