Abstract:
A diagnostic imaging system utilizing a reduced crystal design pattern is utilized to image a subject and collect event data. The reduced crystal design pattern includes filled crystal locations and empty crystal locations. A processor accounts for empty crystal locations by selecting windows that include nearest neighbor filled crystal locations. The nearest neighbor filled crystal locations include event data which is averaged by the processor and assigned to the empty crystal location. A weighted average based on distance or event strength is incorporated.

Description:
BACKGROUND 
       [0001]    The following relates generally to radioemission imaging. It finds particular application in conjunction with positron emission tomography (PET), and will be described with particular reference thereto. A digital PET scanner has several benefits in terms of image quality, but it is costly to design a digital PET scanner. For example, in one design there are over 30,000 crystals to build the illustrative PET scanner. 
         [0002]    The crystal cost is a substantial portion of the overall PET system cost, and reducing the crystal cost would make PET scanners more affordable for less affluent medical facilities, such as those in developing countries. However, since each crystal corresponds to a detector pixel, reducing the number of crystals results in a corresponding reduction in PET detector resolution and hence a corresponding reduction in image resolution. 
       SUMMARY 
       [0003]    The following provides a new and improved system and method which overcome these problems and others. 
         [0004]    In accordance with one aspect, a nuclear imaging system comprises a radiation detector having a regular array of detector pixel locations and including: radiation detector elements occupying some of the detector pixel locations of the regular array, and unoccupied detector pixel locations of the regular array that are not occupied by radiation detector elements. The system may further include one or more processors configured to process radiation event data acquired of a subject using the radiation detector to generate a reconstructed image of the subject by operations including: estimating radiation event data for the unoccupied detector pixel locations based on radiation event data acquired of the subject by radiation detector elements occupying detector pixel locations of the regular array that neighbor the unoccupied detector pixel locations; and reconstructing a data set including both the radiation event data acquired of the subject using the radiation detector and the estimated radiation event data for the unoccupied detector pixel locations to generate the reconstructed image of the subject. 
         [0005]    In accordance with another aspect, an imaging method comprises: acquiring radiation event data of a subject in an imaging region using at least one crystal module arranged around the imaging region, the module having scintillator crystals defining a regular array of detector pixels with some missing detector pixels; estimating radiation event data for the missing detector pixels; and reconstructing the combination of the acquired radiation event data and the estimated radiation event data to generate a reconstructed image of the subject. 
         [0006]    In accordance with another aspect, an imaging system comprises: a radioemission imaging scanner including a scintillator-based radiation detector with some missing scintillator crystals; and a processor programmed to reconstruct acquired radioemission data acquired by the radioemission imaging scanner by operations including: estimating radioemission data for the missing scintillator crystals based on the acquired radioemission data, and reconstructing the combination of the acquired radioemission data and the estimated radioemission data for the missing scintillator crystals to generate a reconstructed image. 
         [0007]    One advantage resides in reduced cost of an imaging system. 
         [0008]    Another advantage resides in little loss in image quality. 
         [0009]    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. 
     
    
     
       BRIEF DESCRIPTION OF DRAWINGS 
         [0010]    The present application 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. 
           [0011]      FIG. 1  illustrates a nuclear imaging system with an iterative-based reconstruction system using an in-reconstruction filter for smoothing. 
           [0012]      FIG. 2  illustrates a typical detector of the nuclear imaging system of  FIG. 2 . 
           [0013]      FIG. 3  illustrates reduced crystal designs. 
       
    
    
     DETAILED DESCRIPTION 
       [0014]    Disclosed herein is a PET scanner with a reduced number of scintillator crystals to reduce production costs of scanners without a corresponding loss of image fidelity. 
         [0015]    With reference to  FIG. 1 , a nuclear imaging system  10  employing a nuclear imaging modality to image a subject is provided. The nuclear imaging modality detects radiation, such as gamma photons, received from a target volume of the subject for imaging. Examples of such nuclear (also called radioemission) imaging modalities include positron emission tomography (PET) and single-photon emission computed tomography (SPECT). As illustrated, the system  10  is a PET imaging system. 
         [0016]    The system  10  includes a nuclear scanner  12 , illustrated as a PET scanner. The nuclear scanner  12  generates raw scan data and includes a stationary gantry  14  housing a plurality of gamma detectors  16  built up from individual detector array units  17  (variously referred to in the art as “modules”, “tiles”, or so forth) arranged around a bore  18  of the scanner  12 .  FIG. 1  shows a plan view of one such detector module  17  as a representative inset. The bore  18  defines an examination volume  20  for receiving a target volume of a subject to be imaged, such as a brain, torso, or the like. The detectors  16  are typically arranged in one or more stationery rings which extend the length of the examination volume  20 . However, rotatable heads are also contemplated. The detectors  16  detect gamma photons from the examination volume  20  and generate the raw scan data. 
         [0017]    With reference to  FIG. 2 , each of the detector modules  17  includes one or more scintillators  22  typically arranged in a regular grid pattern. The pattern of scintillators  22  is further described in detail below. The scintillators  22  scintillate and generate visible light pulses in response to energy depositions by gamma photons. As illustrated, a gamma photon  24  deposits energy in a scintillator  26 , thereby resulting in a visible light pulse  28 . The magnitude of a visible light pulse is proportional to the magnitude of the corresponding energy deposition. Examples of scintillators  22  include sodium iodide doped with thallium (NaI(Tl)), cerium-doped lutetium yttrium orthosilicate (LYSO) and cerium doped lutetium oxyorthosilicate (LSO). 
         [0018]    In addition to the scintillators  22 , the detector module  17  includes a sensor  30  detecting the visible light pulses in the scintillators  22 . The sensor  30  includes a plurality of light sensitive elements  32 . The light sensitive elements  32  are arranged in a grid of like size as the grid of scintillators  22  and optically coupled to corresponding scintillators  22 . In the illustrated embodiment, the light sensitive elements  32  are silicon photomultipliers (SiPMs), but photomultiplier tubes (PMTs) are also contemplated as well as digital silicon photomultipliers (dSiPMs). In illustrative  FIG. 2  there is a one-to-one correspondence between the scintillator crystals  22  and the SiPMs  32 , but this is not required—for example, in another embodiment there is a four-to-one scintillator-to-SiPM ratio in which each SiPM is covered by a 2×2 array of scintillator crystals. 
         [0019]    Each of the SiPMs  32  includes a photodiode array (e.g., Geiger-mode avalanche photodiode arrays), each photodiode corresponding to a cell of the photodiode array. Suitably, the SiPMs  32  are configured to operate in a Geiger mode to produce a series of unit pulses to operate in a digital mode. Alternatively, the SiPMs can be configured to operate in an analog mode. Where the light sensitive elements  32  are PMTs, there is often a many-to-one correspondence between the scintillators  22 . Regardless of the scintillator-to-detector element ratio, in a typical configuration each scintillator crystal serves as a detector “pixel”, that is, the detected scintillation  28  is localized to a single identified scintillator crystal  22 . To this end, the scintillator crystal  22  may be coated with a light-reflective material to contain the scintillation light in the crystal. Additionally or alternatively, signal processing techniques such as Anger logic may be employed to locate which scintillator crystal detected the scintillation event  28 . It is also contemplated to further localize the scintillation event within the scintillator crystal  22  in which it occurs—for example, a depth of interaction (DOI) analysis may be performed based on the spread of light observed at the light detector  30 . 
         [0020]    The scintillator crystals  22  are arranged in a regular (e.g. Cartesian) grid, for example rows and columns of scintillator crystals  22 . More generally, a regular array or regular grid of detector pixel locations comprises detector pixel locations arranged in a repeating pattern across the face of the detector, e.g. in Cartesian rows and columns in the illustrative examples, or in a regular hexagonal layout with a hexagonal repeating primitive unit repeated over the detector face, or so forth). 
         [0021]    However, as seen in  FIG. 2  some scintillator crystals are omitted from the regular grid. In illustrative  FIG. 2  these “missing” scintillator crystals are replaced by filler elements  23 , for example glass elements cut to the same shape as the scintillator crystals  22 . Substituting filler elements  23  for missing scintillator crystals can have advantages in terms of improved structural integrity of the scintillator array, since in some designs neighboring crystals may be in contact and such contact contributes to structural strength of the overall layout. However, it is also contemplated to omit the filler elements  23  and instead have unfilled openings in the (otherwise) regular grid of scintillator crystals  22 . The filler elements  23 , if used, should be inexpensive (at least compared with the scintillator crystals  22 )—for example, the cost of a single LYSO crystal is on the order of $10 USD, whereas an equivalently sized glass element purchased in bulk is much less expensive than this. Thus, the omitted scintillator crystals represent a substantial cost savings. The filler elements  23  are made of a material (e.g. glass) that does not produce an appreciable scintillation in response to absorption of a gamma ray, or the filler elements  22  do not absorb gamma rays at all (i.e. are transparent to the gamma rays). 
         [0022]    More generally, the radiation detector has a regular array of detector pixel locations and includes: (i) radiation detector elements  22  (e.g. scintillator crystals  22 ) occupying some of the detector pixel locations of the regular array, and (ii) unoccupied detector pixel locations of the regular array that are not occupied by radiation detector elements. In the scintillator-based embodiment of  FIG. 2 , the scintillator crystal  22  is the radiation detector element, since the scintillator crystal  22  absorbs the radiation particle  24  thus effectuating its detection. In illustrative embodiments, the unoccupied detector pixel locations (that is, the missing detector pixels) are arranged in a regular sub-array, e.g. as shown in the illustrative examples of  FIG. 3 . 
         [0023]    The illustrative detector module  17  includes an 8×8 array of scintillator crystals  22 , but other sizes are contemplated. Moreover, various module/submodule combinations are contemplated depending upon the sizes of constituent elements, with various nomenclatures for the various modules and sub-modules, e.g. “tiles”, “podules”, “modules”, or so forth. It is appreciated that that the shown geometries are merely illustrative examples, and other geometries with other fill factors can be used and are contemplated. 
         [0024]    Referring back to  FIG. 1 , during a scan of a subject using the scanner  12 , a target volume of the subject is injected with a radiopharmaceutical or radionuclide. The radiopharmaceutical or radionuclide emits gamma photons, or causes gamma photons to be emitted, from the target volume. The target volume is then positioned in the examination volume  20  using a subject support  34  corresponding to the scanner  12 . Once the target volume is positioned within the examination volume  20 , the scanner  12  is controlled to perform a scan of the target volume and event data is acquired. The acquired event data describes the time, location and energy of each scintillation event detected by the detectors  16  and is suitably stored in a data buffer  36 , illustrated as a PET data buffer. 
         [0025]    Subsequent to acquisition, or concurrently therewith, an event verification processor  38  filters the buffered event data. The filtering includes comparing energy (cell counts in the digital mode) of each scintillation event to an energy window, which defines the acceptable energy range for scintillation events. Those scintillation events falling outside the energy window are filtered out. Typically, the energy window is centered on the known energy of the gamma photons to be received from the examination volume  20  (e.g., 511 kiloelectron volt (keV)) and determined using the full width half max (FWHM) of an energy spectrum generated from a calibration phantom. 
         [0026]    For PET imaging, the event verification processor  38  further generates lines of response (LORs) from the filtered event data. A LOR is defined by a pair of gamma photons striking the detectors  16  within a specified time difference of each other (i.e., a coincident event). The specified time difference is small enough to ensure the gammas are from the same annihilation event. For SPECT imaging, the event verification processor  38  further generates a projection line or small-angle cone (generally referred to as a “projection”). A projection is defined by a gamma photon striking the detectors  16 . The LORs or the projections are stored in a list of a list mode memory  40 . Each list item corresponds to a LOR or a projection. 
         [0027]    When using detectors  16  comprising detector modules  17  with missing scintillator crystals (e.g. replaced by the filler elements  23 ), no gamma rays are detected by these missing scintillator crystals. If the number of missing scintillator crystals were sufficiently small, then this would not be a significant problem as the effect would be statistically insignificant. However, in the embodiments disclosed herein, the missing crystals amount to a substantial fraction of the total number of scintillator crystals that would be present in a completely filled regular array. For example, the module  17  would, if completely filled, have 8×8=64 scintillator crystals  22 . As seen  FIG. 2  (or in the inset of  FIG. 1 ), there are 16 missing scintillator crystals in the detector module  17 —that is, 25% of the crystals of the “filled” 8×8 array are missing. In other embodiments disclosed herein (see  FIG. 3 ) the fraction of crystals missing may be 50% or even 75% (other fractions of missing crystals are also contemplated). Accordingly, the “lost” gamma ray acquisitions due to this large fraction of missing scintillator crystals is substantial. 
         [0028]    With continuing reference to  FIG. 1 , to compensate for these lost acquisitions, a missing detector pixel compensator  41  generates estimated, i.e. approximate, data for each missing scintillator crystal based on the acquisition statistics of the neighboring scintillator crystals  22  that are present, as disclosed herein. The approximate data are suitably formatted in list mode format with the event detector location corresponding to the missing scintillator crystal. Thus, the output of the missing detector pixel compensator  41  is a “full” list mode data set can thereafter be processed as if it had been acquired by detectors  16  having no missing detector pixels. 
         [0029]    A reconstruction processor  42 , illustrated as a PET reconstruction processor, reconstructs the list mode data (i.e., a list of projections or LORs, depending upon the imaging modality, with the estimated data for missing scintillator crystals filled in by the missing detector pixel compensator  41 ) into a reconstructed image of the target volume that is stored, e.g. in a PET image memory  44 , and/or is displayed on a display device  48  or so forth. The various data processing components  38 ,  41 ,  42  are suitably implemented as a computer  46  or other electronic data processing device programmed by suitable software or firmware to perform the various functionality as disclosed herein. The computer  46  or other electronic data processing device may be further programmed to serve as a controller or user interface via which a radiology technician or other medical personnel operate the imaging scanner  12  using suitable user interface device(s) such as an illustrative keyboard  50 , mouse, trackball, touch-sensitive display, or so forth. It will also be appreciated that the various data processing components  38 ,  41 ,  42  can be implemented as separate components (as shown) or can be variously integrated together—for example, the missing detector pixel compensator function can be integrated with the event verification component to generate the list mode data. 
         [0030]    With reference to  FIG. 3 , some illustrative reduced crystal design patterns for the detector module are depicted. The full pattern  300  is the conventional crystal design pattern for a crystal module where every crystal location is occupied by a scintillator crystal  22 , i.e. 100% crystal capacity. In one embodiment, a reduced pattern that eliminates 25%  302  of the crystals is used. (This corresponds to the illustrative detector module  17  of  FIG. 1 , inset and  FIG. 2 .) In another embodiment, a reduced pattern that eliminates 50%  304  of the crystals is used. In another embodiment, a reduced pattern that eliminates 75%  306  of the crystals is used. The patterns will be discussed in further detail below. It is appreciated that the patterns shown in  FIG. 3  are diagrammatic tabular representations with numerical indicators depicting the absence/presence of crystals, where each square, i.e. table entry, represents a potential crystal location. In this case, “1” indicates the presence of a crystal  308  at a crystal location and “0” indicates the absence of a crystal  310  at a crystal location. In some embodiments, the missing detector pixels are themselves arranged in a regular grid or array, as this can simplify computing estimated data for the missing detector pixels, as described herein. 
         [0031]    As stated above, list mode data is acquired of a subject using a reduced crystal pattern. Due to the reduced crystal pattern design, the missing detector pixel compensator  41  accounts for missing event data at the pixels where crystals are absent  310  by adding estimated data for those missing pixels based on the actual acquisition statistics of neighboring detector pixels (including directly neighboring, i.e. adjacent, detector pixels and optionally also including further-distant, e.g. next-nearest neighboring, detector pixels). 
         [0032]    In one embodiment, the missing detector pixel compensator  41  accounts for the missing crystals by using an average of event data, i.e. singles counts, from adjacent crystals. For example, in the 25% reduced crystal design pattern  302  (upper right example of  FIG. 3 , also corresponds to the illustrative detector module  17  of  FIG. 1 , inset, and  FIG. 2 ), a 3×3 window  312  can be selected to include 8 filled crystal locations surrounding an absent crystal location. The missing detector pixel compensator  41  averages the event data of the 8 nearest neighbor crystal locations and stores the average as the event data corresponding to the absent crystal location. For absent crystal locations on the edge of the module, and 3×1 window  314  or a 1×3 window  316  can be used to average event data of the filled crystal locations and store the average as the event data for the absent crystal location. It is contemplated that other sized windows can be used. 
         [0033]    In another example, in the 50% reduced crystal design pattern  304 , a 2×2 window  318  3×3 window  319  can be used to average event data of the filled nearest neighbor crystal locations and store the average as the event data for the absent crystal location. For absent crystal locations on the edge of the module, and 3×1 window  314  or a 1×3 window  316  can be used to average event data of the filled crystal locations and store the average as the event data for the absent crystal location. It is contemplated that other sized windows can be used. 
         [0034]    In yet another example, in the 75% reduced crystal design pattern  306 , a pair of nearest neighbor diagonal adjacent filled crystals  320  or two pairs of nearest neighbor diagonal adjacent filled crystals  322  can be used to average event data of the filled crystal locations and store the average as the event data for the absent crystal location. For absent crystal locations  310  that do not have diagonal adjacent filled crystals, a 3×1 window  324  or 1×3 window  326  can be used to average event data of the filled crystal locations and store the average as the event data for the absent crystal location. For absent crystal locations on the edge of the module, and 3×1 window  314  or a 1×3 window  316  can be used to average event data of the filled crystal locations and store the average as the event data for the absent crystal location. It is contemplated that other sized windows can be used. 
         [0035]    In these examples, only immediately neighboring, i.e. adjacent, detector elements are used in estimating the data for the missing scintillator crystal. In another approach, further-distant crystals may also be factored in. For example, next-nearest neighbor elements may be added in. In such a case, since the next-nearest neighbor crystal data are not likely to be as representative as the nearest neighbor pixels, the data can be weighted accordingly, i.e. a lower weight assigned to data from next-nearest neighbor crystals as compared with nearest neighbor crystals. 
         [0036]    To implement the missing detector pixel compensator  41 , a table can be provided, with an entry for each missing detector element (i.e. each missing scintillator crystal) that identifies the extant detector elements that are to be combined to estimate data for the missing detector element along with the weight to assign to the data from each extant scintillator crystal. The estimation can then be rapidly performed as it merely requires retrieving the extant crystal data, weighting, and summing. 
         [0037]    More particularly, in one embodiment, the missing detector pixel compensator  41  adopts a weighted average approach to construct event data values for absent crystal locations  310 . The event verification processor  38  uses a linear or non-linear weight guided function designed to determine the events at the absent crystal location. The weights of the function can be chosen using various weighting schemes. In one embodiment, the weights are chosen using a distance guided approach. The distance guided approach considers the distance of absent crystal locations  310  to filled crystal locations  308  as a factor to determine the weight to be given to the events at a particular crystal (i.e., data from directly adjacent crystals is weighted more heavily than next-nearest neighbor crystals, et cetera). In another embodiment, the event verification processor  38  considers the strength, i.e. number of singles counts, of events at a filled crystal location  308  to determine weighting for that location. 
         [0038]    The foregoing can be directly applied in the case of single photon emission computed tomography (SPECT) data, because in SPECT each detector pixel acquires lines-of-response (LORs) independently of the other detector pixels. By contrast, in PET each LOR is defined by two (nearly) simultaneous 511 keV gamma particle detection events. To account for this, the described averaging can be performed for each detector pixel pair. For example, consider the LOR(i,j) defined between detector pixels indexed i and j, where detector pixel i is a missing pixel and detector pixel j is an existing detector pixel. The LOR count between each existing detector pixel neighboring the missing pixel i and detector pixel j can then be averaged. In the case of estimating a LOR(i,j) where both detector elements i and j are missing, various approaches can be taken. In one approach, these data are omitted from the list mode data set that is reconstructed. Although this may impact the reconstructed image quality to some extent, the impact is relatively low, especially for designs in which the fraction of detector pixels missing is low, e.g. 25% as in the embodiment of  FIG. 1 , inset, and  FIG. 2 . Another approach is to: (1) estimate for LOR(i,j+1) where “j+1” denotes an existing immediately neighboring pixel to missing detector j (so that the situation is converted to the just-described one-missing-pixel case); (2) estimate for LOR(i+1,j) where “i+1” denotes an existing immediately neighboring pixel to missing detector i; and (3) averaging the results of steps (1) and (2). Optionally, this may be repeated for some or all of the surrounding pixels of the missing pixels i and j, depending upon the desired balance between computational complexity and estimation quality. 
         [0039]    Depending upon the nature of the image reconstruction performed by the reconstruction processor  42 , the estimated LORs for missing detector pixels may need to be assigned estimated time stamps. (The time stamp values are not of significance if, for example, the image reconstruction performs forward/backward projection all LORs of the data set without regard to time information). For a static imaging subject which does not move or otherwise change over time, the time stamps for missing detector pixels can be assigned random or pseudorandom values within the time interval of the data acquisition, so as to uniformly fill the time interval. If it is desired to estimate variation in time of the time stamp statistics (as might be appropriate if, for example, in a dynamic imaging task in which inflow/washout of the radiopharmaceutical is measured) then the time stamp distributions over the acquisition time interval for the neighboring detector pixels can be chosen as a “template” and the time stamps for the estimated LORs of the missing detector pixel chosen to conform with that statistical distribution. 
         [0040]    In one embodiment, reduced scan times are achieved due to the reduced number of crystals in the standard field of view. In another embodiment, reduced scan times are achieved by increasing the spacing between full detector blocks and increasing the field of view. 
         [0041]    Referring back to  FIG. 1 , a control system, such as the illustrative computer  46 , suitably provides a graphical user interface (GUI) to allow users to control the scanner  12  to image a subject. For example, the user can coordinate a PET image of a target volume of the subject. Further, by way of the GUI, the control system can be employed to view and, optionally, manipulate images stored in the image memory  44 . For example, an image of the image memory can be displayed on the display device  48 . 
         [0042]    In some instances, one or more of the memories  36 ,  40 ,  44  and/or the processing components  38 ,  41 ,  4  are integrated with the control system, e.g. as a unitary computer system  46 . For example, the reconstruction processor  42 , the missing detector pixel compensator  41 , and the event verification processor  38  can share a common processor. 
         [0043]    As used herein, a memory includes any device or system storing data, such as a random access memory (RAM) or a read-only memory (ROM), a hard disk drive, optical disk, or so forth. Further, as used herein, a processor includes any device or system processing input device to produce output data, such as a microprocessor, a microcontroller (typically with ancillary components such as working RAM memory), a graphic processing unit (GPU), an application-specific integrated circuit (ASIC), an field-programmable gate array (FPGA), and the like; a controller includes any device or system controlling another device or system, and typically includes at least one processor; a user input device includes any device, such as a mouse or keyboard, allowing a user of the user input device to provide input to another device or system; and a display device includes any device for displaying data, such as a liquid crystal display (LCD) or a light emitting diode (LED) display. 
         [0044]    The illustrative embodiments employ a radiation detector design in which each detector pixel corresponds to a single scintillator crystal  22 . Functionally, a detector pixel is the smallest element of the radiation detector to which a radiation detection event can be localized (although some further ancillary localization, such as DOI, may be possible). The detector pixel size thus determines the spatial resolution of the radiation detector. In another contemplated embodiment, direct-detection solid state detector elements are employed, in which radiation is directly detected by absorption of a radiation particle by a solid state detector without an intervening scintillator/scintillation event generating light. In a solid state detector, the detector pixel corresponds to a single solid state detector element that generates a current pulse (or other signal) in response to absorbing a radiation particle. Analogously to scintillator-based embodiments disclosed herein, such a solid state detector can be advantageously constructed in accord with the principles disclosed herein by omitting certain solid state detector elements from the regular array of solid state detector elements making up the radiation detector, and estimating LORs for the missing solid state detector elements based on LOR data acquired by neighboring solid state detector elements. 
         [0045]    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 construed as including all such modifications and alterations insofar as they come within the scope of the appended claims or the equivalents thereof.