Abstract:
In nuclear imaging, solid state photo multipliers ( 48 ) are replacing traditional photomultiplier tubes. One current problem with solid state photomultipliers, is that they are difficult to manufacture in the size in which a typical scintillator is manufactured. Resultantly, the photomultipliers have a smaller light receiving face ( 50 ) than a light emitting face ( 46 ) of the scintillators ( 44 ). The present application contemplates inserting a reflective material ( 52 ) between the solid state photomultipliers ( 48 ). Instead of being wasted, light that initially misses the photomultiplier ( 48 ) is reflected back by the reflective material ( 52 ) and eventually back to the radiation receiving face ( 50 ) of the photomultiplier ( 48 ).

Description:
CROSS REFERENCE TO RELATED APPLICATIONS 
     This application claims the benefit of U.S. provisional application Ser. No. 60/892,890 filed Mar. 5, 2007, which is incorporated herein by reference. 
    
    
     DESCRIPTION 
     The present application relates to the diagnostic imaging arts. It finds particular application to nuclear imaging. Although described with reference to positron emission tomography (PET), it is to be appreciated that the present application relates more generally to pixilated detectors in which an array of scintillator material is coupled to an array of photodetectors, such as in single photon emission computed tomography (SPECT). The present application is also applicable to high energy physics experiments and in astronomy and astrophysics. 
     In past nuclear imaging devices, gamma radiation detectors employed scintillators that convert incident gamma radiation into light, which is then detected by photomultiplier tubes (PMTs). As Due to several drawbacks to photomultiplier tubes, there is interest in replacing them with solid state light sensors, such as avalanche photodiodes driven in a Geiger mode, called e.g., silicon photomultipliers (SiPMs). Typical SiPMs have better timing and energy resolution than typical PMTs. Exact timing, down to the nanosecond range is becoming more valuable as time of flight PET (TOF-PET) scanners are becoming more prevalent. But there have been serious impediments to adopting this new technology. 
     Typically, a scintillator, such as a lantium-bromide scintillator, has a light emitting face that is about 4×4 mm Currently, it is cost prohibitive to manufacture SiPMs that large. Typical SiPMs are manufactured more reliably in 3×3 mm or smaller sizes. This is because that the probability of bad sections of the pixel increases over proportionally to the pixel area. For example it is easier to produce four good 2×2 mm SiPMs than it is to produce one good 4×4 mm SiPM. A 2×2 mm SiPM has a smaller radiation receiving face than the radiation emitting face of a 4×4 mm scintillator. If 2×2 mm SiPMs were coupled with an array of 4×4 mm scintillators, there would be dead space where there is no response to light output by the scintillators between the SiPMs. 
     The consequence of this dead space between SiPMs is that some light emitted by the scintillator will not be collected by the SiPM. In the example of a 4×4 mm scintillator coupled to a 2×2 mm SiPM, the collection efficiency is reduced to 25% based relative to a 4×4 mm SiPM, three-fourths of the light output is lost. With a 3×3 mm SiPM coupled to a 4×4 mm scintillator, the efficiency is 56%. This reduced detection efficiency degrades spatial, energy, and time resolution by about a square root of the area fraction, that is, by a factor of 2 for the 2×2 mm SiPM and by a factor of 1.34 for the 3×3 mm SiPM. 
     Another drawback of SiPMs is that they are more sensitive to red-green wavelengths and less sensitive to blue wavelengths, as emitted by most current scintillators. 
     The present application provides a new and improved radiation detection apparatus and methods for its manufacture that overcome the above-referenced problems and others. 
     In accordance with one aspect, a radiation detector is provided. The detector includes a scintillator that emits light in response to being struck by radiation of a characteristic energy level. The scintillator has a light output face of a first area. The detector also includes a solid state photomultiplier that has a light receiving face of a second area optically coupled to the scintillator. The light receiving face is smaller than the light output face of the scintillator. Reflective material is disposed on portions of the scintillator light output face that are not optically coupled to the solid state photomultiplier light receiving face. 
     According to another aspect, a method of diagnostic imaging is provided. A subject is placed in an imaging region of a diagnostic imaging device. The subject is injected with a radiopharmaceutical. Radiation caused by the radiopharmaceutical is detected with a detector array. The detector array includes a plurality of individual detectors. Each detector includes a scintillator that emits light in response to being struck by radiation of a characteristic energy level optically coupled to a solid state photomultiplier. Each photomultiplier has a radiation receiving face smaller than a radiation emitting face of each scintillator. Light created by the scintillator not initially received by the photomultiplier is reflected back to the photomultiplier. The received radiation is reconstructed into an image representation of at least a portion of the subject. 
     In accordance with another aspect, a method of constructing a radiation detector array is provided. Silicon photomultipliers or photodiodes are disposed in an array on a substrate. Radiation receiving faces of the photomultipliers or photodiodes are optically coupled to light emitting faces of scintillators. The light emitting faces each have a greater surface area than the radiation receiving face to which it is coupled. Reflective material is disposed in a space between photomultipliers or photodiodes to increase light received on the radiation receiving faces of the photomultipliers. 
     One advantage lies in more efficient light harnessing. 
     Another advantage resides in superior timing characteristics. 
     Another advantage resides in more reliable solid-state light detection. 
     Still further advantages of the present invention will be appreciated to those of ordinary skill in the art upon reading and understand the following detailed description. 
    
    
     
       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 the preferred embodiments and are not to be construed as limiting the invention. 
         FIG. 1  is a diagrammatic illustration of a diagnostic imaging apparatus, in accordance with the present application; 
         FIG. 2  is a depiction of an array of scintillators and associated photomultipliers in accordance with the present application; 
         FIG. 3A  is a bottom-up view of an array of scintillators and  FIG. 3B  is a top-down view of a mating array of photomultipliers surrounded by reflective material and guard rings; 
         FIG. 4  shows internal reflections of light in a single scintillator; 
         FIG. 5  is a current response graph of an exemplary photomultiplier as a function of time; 
         FIG. 6  shows an embodiment that includes photomultipliers bonded to the underside of a clear flex foil substrate; 
         FIG. 7  shows an embodiment similar to that of  FIG. 6 , except with holes in the flex foil substrate above the photomultipliers. 
     
    
    
     With reference to  FIG. 1 , a diagnostic imaging device  10  includes a housing  12  and a subject support  14 . Enclosed within the housing  12  is a detector array  16 . The detector array  16  includes a plurality of individual detector elements  18 . The array  16  is arranged so that detector elements  18  are distributed evenly about an imaging region  20 . The detector array  16  can be a ring of detectors  18 , multiple rings, or discrete flat panels disposed opposing each other. Whatever the actual placement or arrangement of the detectors  18 , it is preferable to arrange the detectors such that each detector has a plurality of counterpart detectors across the imaging region to facilitate coincidence detection. In positron emission tomography (PET), pairs of gamma rays are produced by a positron annihilation event in the imaging region and travel in opposite directions. These gamma rays are detected as pairs, with a slight delay (on the order of nanoseconds) between detections if one gamma ray travels farther to reach a detector than the other. 
     Before the PET scan commences, a subject is injected with a radiopharmaceutical. The radiopharmaceutical contains a radioactive element coupled to a tag molecule. The tag molecule is associated with the region to be imaged, and tends to gather there through normal body processes. For example, rapidly multiplying cancer cells tend to expend abnormally high amounts of energy duplicating themselves. So, the radiopharmaceutical can be linked to a molecule, such as glucose that a cell typically metabolizes to create energy, gather in such regions and appear as “hot spots” in the image. Other techniques monitor tagged molecules flowing in the circulatory system. 
     For PET imaging the selected radioisotope emits positrons. The positron can only move a very short distance (on the order of nanometers) before it is annihilated in an annihilation reaction that creates two oppositely directed gamma rays. The pair of gamma rays travel in opposite directions at the speed of light striking an opposing pair of detectors. 
     When the leading edge of a gamma ray strikes the detector array  16 , a time signal is generated. A triggering processor  22  monitors each detector  18  for an energy spike, e.g., integrated area under the pulse, characteristic of the energy of each received gamma ray. The triggering processor  22  checks a clock  23  and stamps each detected gamma ray with a time of leading edge receipt stamp. The time stamp is first used by an event verification processor  24  to determine which gamma rays are a pair which defines a line of response (LOR). Because gamma rays travel at the speed of light, if detected gamma rays arrive more than several nanoseconds apart, they probably were not generated by the same annihilation event and are discarded. Timing is especially important in TOF-PET, as the minute difference in substantially simultaneous events can be used to further localize the annihilation event along the LOR. As computer processor clock speeds become faster, the higher the accuracy with which an event can be localized along its LOR. In a SPECT camera, the LOR or trajectory for each detected gamma ray is determined by collimation. 
     LORs are stored in an event storage buffer  34 , and a reconstruction processor  36  reconstructs the LORs into an image representation of the subject using filtered backprojection or other appropriate reconstruction algorithm. The reconstruction can then be displayed for a user on a display device  38 , printed, saved for later use, and the like. 
     With reference to  FIG. 2 , a portion of the detector array  16  is shown. 
     When a gamma ray  40  strikes the detector array  16  it interacts with one of the individual detector elements  18 . First, the gamma ray  40  passes through a light reflective, gamma ray transmissive top coat  42  and strikes a scintillator  44 . The scintillator  44  converts the gamma ray  40  into a burst of light comprising multiple photons. Some of the photons pass through a light emitting face  46  of the scintillator  44  and hit a solid state photomultiplier  48 , such as an SiPM. The light emitting face  46  of the scintillator is larger in surface area than a light receiving face  50  of the photomultiplier  48 , e.g. 4×4 mm vs. 2×2 mm Resultantly, there is dead space  51  between photomultipliers which do not convert incident light into electrical current or potential. This can be seen also in mating  FIGS. 3A and 3B . 
     Only a fraction of light from the scintillation burst strikes the photomultiplier directly. With reference again to  FIG. 2 , a light reflective material  52  is disposed in the dead space  51  between photomultipliers  48 . The reflective material  52  is preferably a white material. The reflective material  52  can be an initially liquid substance that flows between the photomultipliers and later hardens. The reflective material  52  can also be printed on a substrate  54  using offset printing techniques or applied to the light emitting face of the scintillator. With the reflective material  52  in place, photons that try to exit from the radiation emitting face  46  of the scintillator  44  in the dead space are reflected back up into the scintillator  44  and eventually to the photomultiplier  48 . The photomultipliers  48  and the reflective material  52  are preferably applied directly to the substrate  54 , such as a printed circuit board, a ceramic substrate, or a flex-foil. Also mounted to the substrate are reflective guard rings  56  that separate each photomultiplier  48  to prevent cross-talk between photomultipliers. 
     Disposed between the scintillators  44  and the photomultipliers  48  is a layer of index matching material  58 , such as optical coupling grease. When light reaches a boundary between materials, and the materials have different indices of refraction, some of the light will be transmitted, and some reflected. Because reflection is not desired between the scintillator  44  and the SiPM, the index matching layer  58  is interposed to minimize reflection. Everywhere else, however, reflection is desired to channel as much of the scintillated light as possible into the photomultiplier  48 . Thus, the scintillators  44  themselves are encased in a reflective layer  60 . In this manner, the light reflected by the reflective surface  52  is reflected by the reflective top layer  42  and the reflective side layers  60  to the photomultiplier  48 . 
     Preferably, the photomultipliers  48  are solid state silicon photomultipliers, (SiPM) but it is to be understood that photodiodes are also viable, and are certainly contemplated. SiPMs are most sensitive to the green and longer wavelengths of visible light. Typically, exemplary Lantium Bromide scintillators, like many commonly used scintillators, emit light in the shorter, blue wavelengths. Other scintillators, such as cadmium tungstate, (CdW) bismuth germanium oxide, (BGO) gadolinium orthosilicate, (GSO) cerium doped lutetium orthosilicate, (LSO) cerium doped lutetium yttrium orthosilicate, (LYSO) lead sulfate, cerium fluoride, cerium doped lanthanum fluoride, and the like are also contemplated. The SiPMs  48  can still detect the blue wavelengths, but can detect green or longer wavelengths more efficiently. A wavelength shifting material  62  is disposed between the scintillators  44  and the reflective coatings  60  on at least the vertical sides of the scintillators  44  to shift emitted light from the blue portions of the spectrum to the green. When applied to the surface of the scintillator, the wave shift material optionally has an index of refraction or reflectivity that encourages internal reflection. A view of a single scintillator  44  is shown in  FIG. 4 . Part of the blue light is transferred to the light emitting face  46  via total internal reflections, while other portions of the blue light are shifted by the wavelength shifter  62  to green light. An air gap between the scintillator  44  and the reflector  60  ensures that part of the blue light reaches the light emitting face  46  via total internal reflections, without wavelength shifting due to the index of refraction or reflectivity of the wavelength shift layer  62 . Transmission via total internal reflection is efficient and enables good timing. The light that is wavelength shifted reaches the light emitting face  46  via an increased number of reflections. The wavelength shifting material  62  can also be applied to the reflective material  52 ,  42 , and/or  62  and the region between the scintillator and the reflector could be a waveguide material. 
     A known drawback of this type of wavelength shifting is that it slows the response, which blurs the temporal resolution. To account for this drawback, the radiation emitting face  46  is not coated in wavelength shifting material, so that unshifted scintillated light has a direct path to the photomultiplier. With reference to  FIG. 5 , this results in the initial blue light peak portion  64  with a sharp leading edge  66 , i.e. excellent temporal resolution for generating accurate time stamps and weaker energy resolution. The lower amplitude blue peak portion is followed by a larger green peak portion  68  with excellent energy resolution, but weaker temporal resolution. Thus, the temporal resolution is preserved by omitting the shifting material  62  on the radiation emitting face  46  of the scintillator. At the same time, energy resolution is preserved. 
     Other configurations of the solid state photomultipliers  48  are also viable options. With reference to  FIG. 6 , the SiPMs  48  are disposed beneath a clear flex foil substrate  70 . Disposing the SiPMs  48  below the flex foil  70  facilitates bump bonding to the foil  70 . Index matching material  58  is, again, disposed to lessen loss of light due to reflections. Reflective material  52  is again disposed in the dead space between SiPMs  48 . As shown in the embodiment of  FIG. 7 , holes  72  are cut in the flex foil or other substrate  70 . The holes  72  are situated above the solid state photomultipliers  48 . The opposite surface of the substrate  70  is rendered reflective to reflect light back to a coupled scintillator  44 . The holes  72  are filled with optical coupling material  58  to provide a non-reflective coupling between the scintillator  44  and the SiPM  48 . 
     In yet another viable configuration, multiple discrete photomultipliers can be used to receive scintillations from a single scintillator. For example, four 2×2 mm photomultipliers can be used to receive light from a 4×4 mm scintillator. 
     The invention has been described with reference to the preferred embodiments. Modifications and alterations may occur to others upon reading and understanding the preceding detailed description. It is intended that the invention be constructed as including all such modifications and alterations insofar as they come within the scope of the appended claims or the equivalents thereof.