Abstract:
The present invention is a photodetector including improved photosensors configured of an array of small (sub-millimeter) high-density avalanche photodiode cells utilized to readout a single scintillator. Each photosensor comprises a plurality of avalanche photodiodes cells arranged in an (n×n) array of avalanche photodiode cells (where, n&gt;1) that are coupled to a single scintillation crystal. The overall (n×n) array area as the photosensor is the same as the area of a face of the scintillator and each avalanche photodiode cell has a surface area that is not greater than one square millimeter. The photosensor is also configured to facilitate reading the output of each avalanche photodiode cell in the array. By reading out each small avalanche photodiode cell independently, the noise and capacitance are minimized and thereby provide a more accurate determination of energy and timing.

Description:
BACKGROUND OF THE INVENTION 
       [0001]    I. Field of the Invention 
         [0002]    The present invention relates to medical imaging systems; more particularly, the present invention relates to a high density, highly integrated, APD photosensor (Avalanche Photodiode) array for scintillation crystal readout in detectors utilized in imaging systems using positron emission tomography (PET). 
         [0003]    II. Background Information 
         [0004]    Nuclear medicine is a unique medical specialty wherein radiation is used to acquire images, which show the function and anatomy of organs, bones or tissues of the body. Radiopharmaceuticals are introduced into the body, either by injection or ingestion, and are attracted to specific organs, bones or tissues of interest. Such radiopharmaceuticals produce gamma photon emissions, which emanate from the body and are captured by a scintillation crystal, with which the photons interact to produce flashes of light or “events.” Events are detected by an array of photo detectors, such as photomultiplier tubes, and their spatial locations or positions are calculated and stored. In this way, an image of the organ or tissue under study is created from detection of the distribution of the radioisotopes in the body. 
         [0005]    One particular nuclear medicine imaging technique is known as Positron Emission Tomography, or PET. PET is used to produce images for diagnosing the biochemistry or physiology of a specific organ, tumor or other metabolically active site. Measurement of the tissue concentration of a positron emitting radionuclide is based on coincidence detection of the two gamma photons arising from positron annihilation. When a positron is annihilated by an electron, two 511 keV gamma photons are simultaneously produced and travel in approximately opposite directions. Gamma photons produced by an annihilation event can be detected by a pair of oppositely disposed radiation detectors capable of producing a signal in response to the interaction of the gamma photons with a scintillation crystal. Annihilation events are typically identified by a time coincidence between the detection of the two 511 keV gamma photons in the two oppositely disposed detectors, i.e., the gamma photon emissions are detected virtually simultaneously by each detector. When two oppositely disposed gamma photons each strike an oppositely disposed detector to produce a time coincidence event, they also identify a line of response, or LOR, along which the annihilation event has occurred. 
         [0006]    An example of a PET method and apparatus is described in U.S. Pat. No. 6,858,847, which patent is incorporated herein by reference in its entirety. After being sorted into parallel projections, the LORs defined by the coincidence events are used to reconstruct a three-dimensional distribution of the positron-emitting radionuclide within the patient. PET is particularly useful in obtaining images that reveal bioprocesses, e.g. the functioning of bodily organs such as the heart, brain, lungs, etc. and bodily tissues and structures such as the circulatory system. 
         [0007]    In order to minimize patient exposure to radiation, detectors utilized in PET imaging systems must be able to detect low levels of incident optical photons or ionizing particles. In such imaging devices it is often advantageous to employ radiation detection devices having internal gain; avalanche photodiodes (APDs) are commonly used in such devices to provide the desired detection sensitivity. An APD is a semiconductor device that is biased near the breakdown region such that charge generated as a result of the absorption of an incident photon is amplified in the APD itself as a result of a cascading effect as charge is accelerated by the high bias potential applied across the p-n junction of the device. In such imaging devices, it is desirable that the APD exhibit low noise and high gain. Certain devices, such as medical imagers (e.g., using gamma radiation), also require relatively large arrays (e.g., about 5 cm.sup.2 or larger) of high quality, low noise APDs. 
         [0008]    One detector configuration in PET imaging systems utilizing APD arrays is configured in a one-to-one coupling configuration, where one APD is coupled to one scintillation crystal. To collect the maximum amount of light from the scintillator, the APD has to have the same surface area as the scintillator crystal to which it is coupled. This can result in large surface area APDs, which increases the APD noise and capacitance. The noise and capacitance of an APD is directly proportional to its surface area. As the noise and capacitance of the APD increases, its capacity to accurately determine the proper energy and timing of an event decreases. This results in poor PET detector performance and is a main reason why APDs are not commonly used as photosensors in detectors in PET imaging systems. There is a need for a photosensor array configuration that can take advantage of the high gain resulting from the use of APDs in the array while simultaneously reducing the noise and capacitance resulting from increasing the size of APDs when used in photosensors. 
         [0009]    One detector which utilizes the concept of a high density photosensor readout coupled to a single scintillator is the SiPM (Silicon Photomultiplier). A SiPM uses a very dense array of Geiger-mode APDs which are typically resistively connected in parallel to provide a single readout channel. A significant drawback of this type of this type of detector is that each of the SiPM APD cells is nonlinear since they operate in Geiger mode. Detectors utilizing SiPM Geiger mode APDs in each cell operate in binary mode. Accordingly, the respective output of SiPM APD cells can only be zero or one. This is a fundamental problem, because each SiPM APD cell is only capable of counting one photon and cannot indicate that more than one photon has been received. For example, if two photons reach the same SiPM APD cell at the same approximate time, there is no way of knowing that two photons have been received by the SiPM APD cell. To increase the linearity of a SiPM, the SiPM cell density must be increased. However, increasing the cell density of a SiPM causes a decrease in the fill factor of the device. For most SiPM devices, there is a trade-off between linearity and fill factor. There is a need for a configuration of APD arrays that overcomes this limitation wherein an array of APD cells operate in proportional mode and are linear. 
         [0010]    Some developers in the industry have investigated operating SiPMs below the breakdown voltage in proportional mode for CT imaging to overcome the inherent nonlinearity of the device. However, the proposed detector still uses a common readout of the summed SiPM cells. There is a need for a system that facilitates individual cell readout of the cells comprising an APD array. 
       SUMMARY OF THE INVENTION 
       [0011]    Consistent with embodiments of the present invention, the present invention comprises an improved photodetector including improved photosensors configured as an array of small (sub-millimeter) high-density avalanche photodiode cells utilized to readout a single scintillator. Each photosensor comprises a plurality of avalanche photodiode cells arranged in an (n×n) array (where, n&gt;1) that are coupled to a single scintillation crystal. The overall (n×n) array area of the photosensor is substantially the same as the area of a face of the scintillator and each avalanche photodiode cell has a surface area that is not greater than one square millimeter. The photosensor array is also configured with circuitry to facilitate reading the output of each avalanche photodiode cell in the array separately. 
         [0012]    It is to be understood that both the foregoing summary of the invention and the following detailed description are exemplary and explanatory only, and should not be considered restrictive of the scope of the invention, as described and claimed. Further, features and/or variations may be provided in addition to those set forth herein 
     
    
     
       BRIEF DESCRIPTION OF THE DRAWINGS 
         [0013]    The accompanying drawings, which are incorporated in and constitute a part of this disclosure, illustrate various embodiments and aspects of the present invention. In the drawings: 
           [0014]      FIG. 1   a  is a side view of a sub-millimeter 2×2 APD array coupled to a single scintillator crystal in one embodiment of the present invention; 
           [0015]      FIG. 1   b  is a top view of a sub-millimeter 2×2 APD array coupled to a single scintillator crystal in one embodiment of the present invention; 
           [0016]      FIG. 2   a  is a side view of a sub-millimeter 10×10 APD array configured with an independent bias circuit, in one embodiment of the present invention; 
           [0017]      FIG. 2   b  is a side view of a sub-millimeter 10×10 APD array configured with a common bias circuit, of the present invention; 
           [0018]      FIG. 3  is a schematic example of an embodiment of the biasing circuit and the signal output connection of the APD array illustrated in  FIG. 2   b  in one embodiment of the present invention; 
           [0019]      FIG. 4   a  is a side view of a package configuration of a sub-millimeter APD array connected to a frond end APD-ASIC in one embodiment of the present invention; 
           [0020]      FIG. 4   b  is a side view of a package configuration of a sub-millimeter APD array connected to a frond end APD-ASIC in a second embodiment of the present invention illustrating a smaller ASIC die; 
           [0021]      FIG. 5  is an embodiment of a schematic block diagram of a single APD readout circuit from the bias circuit to the energy and timing outputs of the APD ASIC; and 
           [0022]      FIG. 6  is an illustration of an embodiment of back-end circuitry within the APD ASIC or system processing electronics of the present invention that may be utilized to determine energy, position and timing of each PET event. 
       
    
    
     GENERAL DESCRIPTION 
       [0023]    Consistent with embodiments, the present invention is an improved photodetector configured for use in PET imaging systems, wherein the improved photodetector shall include at least one scintillator crystal and an APD photosensor array coupled to the at least one scintillator crystal. The APD photosensor array is sized so that its surface area is substantially equivalent to a surface area of a face of the scintillator crystal. The APD photosensor array comprises a plurality of avalanche photodiodes arranged in an n×n array of cells that include support circuitry. Each APD cell has a separate output for the avalanche photodiode positioned thereon so that each avalanche photodiode can be read independently. Each APD cell is defined as sub-millimeter because the surface area of each APD cell is &lt;1.0 mm×&lt;1.0 mm. In one embodiment, the surface area of each APD cell shall be 0.05 mm 1.0 mm×0.05 mm-1.0 mm. By reading the output from each small sub-millimeter APD cell independently, the noise and capacitance are minimized and thereby provide a more accurate determination of energy and timing. 
       DETAILED DESCRIPTION 
       [0024]    The following detailed description refers to the accompanying drawings. Wherever possible, the same reference numbers are used in the drawings and the following description to refer to the same or similar parts. While several embodiments and features of the invention are described herein, modifications, adaptations and other implementations are possible, without departing from the spirit and scope of the invention. Rather these embodiments are provided so that this disclosure will be complete and will fully convey the invention to those skilled in the art. For example, substitutions, additions or modifications may be made to the components illustrated in the drawings, and the methods described herein may be modified by substituting, reordering or adding steps to the disclosed methods. Accordingly, the following detailed description does not limit the invention. Instead, the proper scope of the invention is defined by the appended claims. 
         [0025]    The present invention comprises the use of a high density, highly integrated APD (Avalanche Photodiode) array to readout one or more scintillation crystals within a photo detector used in PET imaging applications. One embodiment of the invention, shown in  FIGS. 1A and 1B , illustrate a block detector  100 , utilizing a high density, highly integrated APD array  120  to read each scintillation crystal  112 ,  114 ,  116 ,  118  within an array of scintillation crystals  102 . 
         [0026]    Unlike a typical detector in PET imaging systems that utilize a one-to-one crystal to photosensor configuration, or a block detector where the number of photosensors is less than the number of crystals in the configuration, the detector embodied in the present invention utilizes a configuration in which the number of photosensors, APDs, is always greater than number of crystals. In the example embodiment illustrated in  FIGS. 1A and 1B , there are twenty-five photo sensors to each scintillation crystal. In the embodiment illustrated, a two by two array of scintillation crystals  102  is coupled to a ten by ten array of photo sensor cells  120 , wherein each cell includes an APD. As illustrated, each scintillation crystal  112 ,  114 ,  116 ,  118  is coupled to a five by five array of APD photo sensor cells positioned to read out each scintillation crystal  112 ,  114 ,  116 ,  118  separately. As illustrated in  FIG. 1B , crystal one  112  is coupled to a first array of APD photo sensors  122   a - 122   y ; crystal two  114  is coupled a second array of APD photo sensors  124   a - 124   y ; crystal three  116  is coupled to a third array of APD photo sensors  126   a - 126   y ; and crystal four  118  is coupled to a fourth array of APD photo sensors  128   a - 128   y . As illustrated in  FIG. 1B , the combination of scintillation crystals one  112 , two  114 , three  116  and four  118  into a array shows that that photo detectors  100  of the type generated using the present invention may be a subset of a larger array detector. 
         [0027]    Each APD in the APD array  120  reading light from the block of scintillation crystals  112 ,  114 ,  116  and  118  are read out independently through ball grid array (BGA) connections to a highly integrated ASIC  136  to provide accurate energy and timing information.  FIG. 1A  illustrates the BGA  132   a - 132   e  connecting APD cells  122   a - 122   e  to ASIC  136  and the BGA  134   a - 134   e  connecting APD cells  124   a - 124   e  to ASIC  136 . (Similar BGA connecting the remaining APD cells to ASIC  136  exists but are not shown in  FIG. 1A ). 
         [0028]    The APD photo sensor array  120  incorporated into block detector  100  comprising the present invention is a monolithic or assembled array of sub-millimeter APDs. A top view of embodiments of the APD photo sensor array  120  illustrated in  FIGS. 1A and 1B  is illustrated in  FIGS. 2A and 2B . In the embodiments illustrated, a ten by ten array of APD photo sensors or a five by five array of APD photo sensors are shown. It is to be understood that the specific array of APD photo sensors illustrated in  FIGS. 1B ,  2 A and  2 B are for illustrative purposes only and are not intended to limit the scope of the present invention. The dimensions of an array of APD photo sensors may be any n×n array configuration assembled in accordance with the present invention. Generally, the dimensions of the APD photo sensor array may be determined by the size of the scintillator crystal and the overall desired size of the detector. In one embodiment, it is contemplated that the each APD photo sensor cell in the array of APD photo sensors  320 ,  420  shall be approximately 1.0 millimeters (mm) per side or less (each cell has an area of &lt;1.0 mm×&lt;1.0 mm) and each APD photo sensor cell operates independently. In some embodiments, each APD photo sensor cell in the array of APD photo sensors  320 ,  420  may have an area of &lt;0.05-1.0 mm×&lt;0.05-1.0 mm. As illustrated in the embodiments, shown in  FIGS. 2A and 2B , each APD photo sensor cell in the array of APD photo sensors  302 ,  402  is connected to a common bias circuit to facilitate high voltage bias and signal readout. The common bias provides for easier integration and a higher fill factor. Although it is not shown in the Figures, it is contemplated that in another embodiment, each APD photo sensor cell in the array of APD photo sensors  320  may have an independent bias circuit. 
         [0029]    Of the two embodiments illustrated in  FIGS. 2A and 2B , the embodiment illustrated in  FIG. 2B  is preferred, as the embodiment illustrated in  FIG. 2A  is more difficult and costly to manufacture. A problem with the design illustrated in  FIG. 2A  is the high and low voltage lines  312  and  314  are too close. Electrical lines, such as lines  312  and  314 , configured on standard printed circuit board material that are in such close proximity cannot function properly. Since there will be a substantial difference between the high and low voltage lines  312  and  316 , arcing will occur between the high and low voltage lines, unless the base material, of which the array of APD photo sensors is comprised, is a material such as ceramic or Teflon. Accordingly, using an inexpensive base material such as FR4 to create the array of APD photo sensors is not feasible because the physical properties of the base material cannot support voltage lines of varying voltage being in close proximity. However, the problem resulting from having two voltage lines in close proximity on a chip may be overcome when a design, such as that illustrated in  FIG. 2A , is manufactured using a base material such as Teflon or ceramic. While it is contemplated that an embodiment of the invention can be manufactured in accordance with design set forth in  FIG. 2A , manufacturing such a design adds substantial expense to the process when base materials such as ceramic and Teflon are used. 
         [0030]    An alternative embodiment, illustrated in  FIG. 2B  allows for the use of inexpensive underlying base material, such as FR4 by eliminating the arcing problem through increasing the distance between the high voltage lines  442 ,  446 ,  450 ,  454 ,  458 ,  462  and the low voltage lines  444 ,  448 ,  452 ,  456 ,  460 . As illustrated in  FIG. 2B , this is accomplished by having the anode and cathodes of two adjacent APDs in APD photo sensor cells within an array share high and low voltage lines. As shown, APD cell  424   e  is connected to high voltage line  442  through connection  404 . APD cells  424   e  and  424   j  are both connected to low voltage line  444  through connections  406  and  408 , and APD cells  424   j  and  424   o  are both connected to high voltage line  446  through connections  410  and  412 . It is to be understood that whether the anode or cathode of an APD is an APD cell, in the present invention, connected to high or low voltage is irrelevant. The important point is to have a voltage difference between the anode and cathode of the APD. In one embodiment, the anode may be connected to ground and the cathode connected to a positive high voltage. In an alternative embodiment, the cathode may be connected to ground and the anode connected to a negative high voltage. 
         [0031]    Referring to  FIG. 3 , this figure presents an illustration of a schematic drawing of an example of the common biasing circuit and the signal output connection of an array of APD photo sensors  140  comprised of ten APD photo sensor cells  142   a - 142   j . This illustration is a more detailed schematic of a portion of the array of APD photo sensor array  302  illustrated in  FIG. 2B . As illustrated, the array of APD photo sensors  140  illustrated comprises a five by two array of APD photo sensor cells  140 . As illustrated, each of the ten APD photo sensor cells  142   a - 142   j  are comprised of the same components and circuitry and are connected to ground and a common high voltage line in the same manner. Accordingly, of the ten APD photo sensor cells  142   a - 142   j , cell  142   a  will be explained to illustrate operation of each APD photo sensor cell in array  140 . The input of APD photo sensor cell  142   a  is connected to high voltage line  144 , wherein the cathode of the APD  150   a  is connected to the high voltage line  144  through resistor  146   a . The cathode of APD  150   a  is also connected to a capacitor  148   a  and ground  158   a . The anode of APD  150   a  is connected to ground  158   a  through resistor  152   a  and the output of the APD cell  142   a  to the ASIC (not shown) through a bump-bond connection  156   a  through capacitor  154   a . The high voltage capacitor  148   a  is used to decouple noise from the high voltage line; the regular low voltage capacitor  154   a  is AC-coupling the APD signal to subsequent signal chain. Capacitor  154   a  can be eliminated if the DC-coupling signal path is desired. Node  156   a  D 11 , subsequence CFA, capacitor  156   a , and ground plane  158   a  also form a low-impedance current loop for signal high frequency components, facilitating event timing detection. 
         [0032]      FIGS. 4A and 4B  are side views of two embodiments of package configurations of the sub-millimeter APD photo sensor arrays  160  and  170  and the respective front-end APD-ASIC  166  and  176  connected through bump-bond connections  168  and  178 . As illustrated in  FIG. 4B , the embodiment shown also includes a printed circuit board sandwiched between the APD photo sensor array  170  and the front end APD-ASIC  176 , allowing the size of the APD-ASIC  176  to be substantially reduced. 
         [0033]    When light enters a photo sensor, it is desirable to know two pieces of information, when the photon is received by the photo sensor (timing) and how large the signal received is (amount of energy or the number of photons received at a given instance in time). This configuration facilitates the ability to determine the timing at which photons are received and the number of photons received by facilitating the ability to determine the amount of energy received by each APD at an instance in time, thereby making the embodiment linear.  FIG. 5  illustrates the circuitry within the APD cell  184  and the APD ASIC  180  utilized to determine the timing and energy information necessary in PET imaging applications. As illustrated, a sub-millimeter APD cell  184  having its output  196  connected to a charge sensitive pre-amplifier (CSP)  202 , is positioned on the APD ASIC  180 . The CSP  202  facilitates the collection of the charges from the APD  190 . Next the CSP  202  output signal is split into to channels, a slow channel  204  and fast channel  206 . The signal output by the CSP  202  has a high frequency and a low frequency component. The fast channel  206 , which receives the signal output by the CSP  202 , is performing filtering or shaping of the signal received to emphasize the high frequency components. The output of the fast channel  206  is received by a trigger  208 , and a constant fraction discriminator (CFD) or a leading edge (LE) trigger  210 . The slow channel  204 , which receives the signal output by the CSP  202 , is performing filtering or shaping of the signal received to emphasize the low frequency components. The output from slow channel  204  is representative of the energy received by the APD cell  184  and is proportional to the number of photons received by the APD cell  184 . 
         [0034]      FIG. 6  is an illustration of the circuitry within the APD ASIC, including the circuitry from  FIG. 5  at  180   a - 180   d , where stages of sub-millimeter APD cells, each are connected to CSPs positioned on the APD ASIC and connected to a slow channel and a fast channel that provide energy outputs  212   a - 212   d  and timing outputs  210   a - 210   d . Each energy output  212   a - 212   d  is received by an energy multiplexer  222  and a cell location  224  unit (a look up table). Each timing output is received by a timing multiplexer  226 . The output of the timing multiplexer  226  is received by a timing digital convertor  236 . The output of the energy multiplexer  222  is received by an analog to digital convertor  232 . The output of the cell location is received by the look up table  234  to perform event fine positioning. All signals exiting the analog to digital convertor  232 , the look up table  234  and the timing digital convertor have been digitized and are then processed by the processor  238  which is a commercially available chip on the market. It is contemplated, as illustrated in  FIG. 6 , that the circuitry shown, other than the APD Cell and the processor, is on the APD ASIC. However, it is to be understood that it is contemplated that the circuitry illustrated in  FIG. 6  is not required to be implemented through the APD ASIC and may be implemented in various embodiments on other printed circuit boards. 
         [0035]    The above specification, examples and data provide a description of the manufacture and use of the invention. Since many embodiments of the invention can be made without departing from the spirit and scope of the invention, the invention resides in the claims hereinafter appended.