Abstract:
Methods of a high resolution, stationary imaging detector for use in systems for positron emission tomography or single photon emission tomography that uses shadowing effects from intensity attenuation to provide three dimensional positioning information for a source of activity within a field of view of the detector.

Description:
CROSS-REFERENCE TO RELATED APPLICATIONS 
       [0001]    This application claims priority to U.S. Provisional Application No. 61/932319, filed Jan. 28, 2014, which is incorporated herein by reference. 
     
    
     FIELD OF THE INVENTION 
       [0002]    The following relates to positron emission tomography (PET) imaging, and more specifically to a single detector for PET imaging. It also relates to single photon emission computed tomography (SPECT) imaging. 
       BACKGROUND OF THE INVENTION 
       [0003]    Within medical imaging, there are several different methods used to develop images for medical diagnosis of a patient. These methods include ultrasound (US), magnetic resonance imaging (MRI), computed tomography (CT), single photon emission computed tomography (SPECT), and positron emission tomography (PET). 
         [0004]    For PET and SPECT imaging, a patient is injected with a radiopharmaceutical. The radiopharmaceutical associated with PET imaging has a radionuclide that produces gamma particle photon pairs with opposing trajectories from positron annihilation. In SPECT imaging, single photons are produced from a radionuclide with a trajectory from the source of activity in the radionuclide. 
         [0005]    Current PET Technology traditionally uses two distinct, separate, and opposing detectors to determine a Line of Response (LOR) of a positron emission event. This could be in the form of a ring (as in Whole Body PET) or as paddles in a high resolution PET system. This is needed in order to detect two distinct gamma particles moving in opposite directions. Gamma particles are created from an event where a positron interacts with an electron and annihilation occurs. 
         [0006]    When photons impact scintillation crystals, some gamma particles have energy transferred to visible light. This light is detected by a photomultiplier tube (PMT) or a silicon photo multiplier (SiPM). Electrical signals from the PMT or SiPM are used for event and position detection. These signals are typically in a pulse format that are sent to electrical circuits for amplification and pulse height detection. 
         [0007]    With PET, when two gamma particles come into contact of each opposing detector, this is known as a true coincidence event. The timing window for this contact being detected between each detector typically has a range between 0 to 8 nanoseconds. A random event occurs when only one of the gamma particles comes into contact with one of the detectors. The random event cannot provide a LOR since two points were not detected to determine a line. A random coincidence event occurs when two gamma particles from two different annihilation events within the coincidence timing window. This can generate the LOR for imaging, but is incorrect since the LOR was created from two independent events and not a common single coincidence event. Random coincidence events can have a negative impact on the performance of a PET imaging machine. 
         [0008]    The annihilation event occurs within the Field of View (FOV) in order for the true coincidence event to occur and determine the line of response for PET imaging. The time between the two gamma particles impacting the detectors for scintillation can be used to discriminate for random events. Random coincident events may be discriminated by other methods since they have the same time occurrences as the coincident event. 
         [0009]    When gamma particles generated has a trajectory through material, there are three types of interactions that can occur. They are photoelectric process, Compton scattering process, and pair production process. The combined effects from these three processes are known as attenuation. The gamma photons will either pass through the material, be absorbed by the material or change its trajectory and “scatter”. Based on a beam of photons entering into the material of initial intensity (I o ), the intensity attenuation of the gamma photons (I t ) can be determined: 
         [0000]      I t =I o e −ux    
         [0000]    Where x is the thickness of the material and −u is the attenuation coefficient. The attenuation coefficient is dependent on the density of the material, and the photon energy of the gamma particle. For positron-electron annihilation and single photon emissions, the photon energy is typically 511 keV. 
         [0010]    With SPECT technology, a single photon is emitted from events of radionuclide activity that is injected into a patient. The photons are detected through the use of a gamma camera where a 2D image is captured. The gamma camera uses collimators for line of sight detection of the emitted gamma photon. The camera is moved with different position and angles so that a 3D image can be generated. 
         [0011]    As discussed above, conventional PET systems use two separate and opposing detectors for determining true coincidence of annihilation events. With the drawbacks of conventional systems discussed above, it would be desirable to have a single detector that can be used for three dimensional imaging in medical diagnostics. A multi-detector configuration, such as a ring configuration, is not needed with the use of a single detector or detector arrays. This single detector embodiment provides high resolution stationary scans with the detector in close proximity with the patient&#39;s body. 
       SUMMARY OF THE INVENTION 
       [0012]    The present invention generally provides improved devices, systems, and methods for three dimensional imaging in medical diagnostics using a single detector. Some of the advantages of a single detector are lower cost and higher mobility than current PET devices. 
         [0013]    Provided is a plurality of detector configurations that uses gamma particle intensity attenuation materials that are positioned next to scintillation crystals. The attenuation materials can provide either a shadow or collimation effect to the scintillation crystals based on the location gamma particle source of activity above the detector configurations. This shadow effect provides angle information about location of source activity without the use of an LOR from coincidence events or a two detector PET system. 
         [0014]    The plurality of single detector configurations of this embodiment is independent from coincidence events for PET imaging. These detector configurations are then inherently immune to random and random coincidence events. Therefore, method for discrimination of these types of events is not needed by the system of this embodiment. 
         [0015]    A single detector gamma camera is used for SPECT imaging systems and is dependent on scanning methods where the detector is moved to different positions and angles in order to reconstruct a 3D image. The plurality of single detector configurations of this embodiment provides 3D imaging from single photon emissions from a stationary position. Unlike SPECT imaging systems, these detector configurations include collimated as well as non-collimated photons for imaging. 
     
    
     
       BRIEF DESCRIPTION OF THE DRAWINGS 
         [0016]      FIG. 1  provides a system view of the single detector imaging machine from photon event occurring to image reconstruction. This figure can be applied to both PET and SPECT technologies. 
           [0017]      FIGS. 2   a  and  2   b  show two different views of one embodiment of a detector construction.  FIG. 2   a  shows a perspective view of a three dimensional representation of a detector block configuration.  FIG. 2   b  shows a sectional side view of a portion of the detector array. It continues in each horizontal direction since the construction materials are alternating repetitively. 
           [0018]      FIG. 3  shows a sectional side view of a portion of the detector illustrating the shadow technique created from photon intensity attenuation due to interaction with material. The image shows two beams of gamma particles with a trajectory that is passing through attenuation material and one beam that is not due to line of sight. The image indicates a shadowing effect being detected by the scintillation crystals based on the trajectories of these two photon beams. 
           [0019]      FIG. 4  shows a sectional side view of the detector with different angles of shadow projections being created based on the location of the source of activity. It also shows the effects from a collimator when activity is in line of sight of all scintillation crystals of a crystal element. 
           [0020]      FIGS. 5   a  and  5   b  show top views of two embodiments of different possible geometries of the detector array. The pixelated crystals and attenuation shielding material can provide the shadow effects using various shapes.  FIG. 5   a  shows a hexagonal geometry for both the crystal pixels and the shielding attenuation material.  FIG. 5   b  shows a square geometry for both the crystal pixels and shielding attenuation material. 
           [0021]      FIG. 6  is a flow chart that shows a method for determining the height of the source of activity needed for 3D reconstruction. The flow chart is based on the number of crystal elements and axis of rotation that the geometry the detector provides. 
       
    
    
     DETAILED DESCRIPTION 
       [0022]    The present invention provides a single detector used to detect particles, for example gamma particles, from a positron emission can provide a low cost, extremely portable solution. A single detector allows for a hand held scanner module similar to an ultra sound scanner. 
         [0023]    The present invention also provides a method for a Field of View (FOV) deep into the chest wall, which is a limitation from the current PET/PEM technology and digital mammography. It also addresses the issue where digital mammography can provide false positives with dense breast tissue. 
         [0024]    A single ended detector PET scanner adds flexibility in that it can easily adapt to scanning different parts of the body such as the thyroid, and other soft tissue. The single detector is not dependent on coincident events and is immune to random and random coincident events that occur within the FOV. 
         [0025]    Current PET scanners require a radiopharmaceutical injection for the whole body. For high resolution localized scanning, the injection must not be in close proximity to the FOV for the scan since random coincident events affect the image performance. With a single ended detector, a localized injection could be done since random coincident event have no effect on detection. A localized injection minimizes prep time, exposure and makes it more convenient than current PET scans. 
         [0026]    Some embodiments of the single detector places pixelated crystals next to shielded material with a greater height, alternating the pixelated crystal slices with shielding material slices. A point source within the FOV space will emit gamma particles toward the single detector. Depending upon the location of the point source with the FOV, the shielding will reduce the radiation intensity on the opposite side of the shield creating a “shadow” effect that ends based on the angle from the top of the shield to the end of the shadow. This angle can then be extrapolated to the position of the point source for mapping. Since there is more than one shield slice, the shadow angle increases for each slice positioned closer to the source location. Eventually, the slices closest to the point source will be fully exposed to the source and no shadow will exist. 
         [0027]    The single detector imaging system is used to detect photon events from a source of activity and reconstruct an image to indicate the location of the source of activity.  FIG. 1  shows the basic assemblies for this embodiment in imaging. Events from a source of activity  101  are created from positron emission or single photon emission within a radionuclide. When these events occur within the FOV of the detector assembly  103 , some events can be detected that have a trajectory  102  that impacts the detector assembly  103 . The detector assembly  103  is able to convert the energy of the photon into electrical pulses  104 . Characteristics of electrical pulses  104  provides data representation of X and Y positioning information for the location of where scintillation occurred within the detector from impact of the photon particle. The electrical pulses are detected by front end electronic assembly  105  and convert the X and Y data from the pulses  104  into binary data  106 . The binary data  106  provides a format for the X and Y positioning that is understood by a computer assembly  107 . The computer captures the binary data  106 , generates a two dimensional histogram for all binary data  106  captured, and reconstructs an image that is then projected onto a display within the computer assembly  107 . 
         [0028]      FIG. 2   a  shows one embodiment of a basic construction of a single detector block  200  that is used to detect gamma photons emitted from a radiopharmaceutical injected into a patient. The radiopharmaceutical can provide single emission photons as in SPECT imaging or paired positron emission photons for PET imaging. In some embodiments, multiple detector blocks can be connected together to form an array and increase detection surface area. Increased incidental surface area improves system detection performance for higher sensitivity and increased field of view (FOV). The shape of the scintillation crystals  202  in this embodiment are square, but can be in different shapes and sizes. In some embodiments, the material of crystals  202  may be bismuth germinate (BGO), sodium iodide doped in thallium NaI(TI), lutetium yttrium oxyorthosilicate doped with cerium (LYSO:Ce), or other crystals used for a scintillation process. The crystals  202  are pixelated in that they are typically cut 1 to 4 mm in width and length but not limited to these dimensions. They are connected together with a thin optical isolation material or film  203  between them for separation. This film  203  is reflective on both sides so the light energy is not lost or contaminated into neighboring crystals during the scintillation process. The reflective film  203  is also applied to the top of the crystals  202  to ensure maximum light transfer to the PMT  205 . 
         [0029]    The pixelated crystals  202  are connected and arranged to form a crystal element that is surrounded by attenuation material  201 . Material  201  is typically used as shield from radiation exposure or as a collimator in gamma cameras in SPECT systems. In the construction of this detector block, the walls of material  201  are thick enough to significantly attenuate photons. The materials used for a detector block for attenuation can be tungsten, lead, or other high density materials that attenuate gamma particles. For tungsten, the half value layer (HVL) is between 3 and 4 mm which is the thickness needed to attenuate half the photons that incidentally enter the attenuating material  201 . This value is based on the intensity attenuation equation. Like the shape of the pixelated crystal  202 , the geometry of the attenuation material  201  around each crystal element can be in the form of various shapes and different from the crystal  202  shape itself. Multiple or single crystal elements can be formed on a single detector block. 
         [0030]    The depth or thickness of the pixelated crystals  202  is shown  FIG. 2   b . The thickness of these crystals is dependent on the density or stopping power of the material. This is usually between 10 mm and 15 mm, but the invention is not limited to this range. The number of pixels for the crystal  202  should be greater than six in order to develop a histogram count pattern with certain resolution. More pixels will give a better resolution of the count pattern and system. On each side of the crystals  202  at the end of each crystal element, the attenuation shield  201  is connected. This material is relatively taller than the crystal surface and extends down to and even with the bottom of the crystals  202  or can continue and make contact with the PMT  205  directly. 
         [0031]    Both the shield  201  and the crystals  202  are connected to one side of an optic coupler  204 . The opposite side of the coupler is connected to a PMT or SiPM  205 . The coupler  204  provides a method to appropriately transfer the light energy created from the scintillation process of the crystals  202  to the PMT  205 . It also provides adhesion of the crystals  202  and shield  201  to the PMT  205 . Materials can vary for the coupler  204  and can include light guides, translucent adhesive, resin or glass, but is not limited to these materials. 
         [0032]      FIG. 2   b  also indicates the detector block being extended horizontally into a detector array. This extension can continue until the length is sufficient for the FOV identified by the system geometry such as ring, partial ring, or a paddle. 
         [0033]    For  FIG. 3 , the height of the attenuation shield  301  from the crystal surface should be between half of the width to the full width of the crystal element  302 . The height can be outside this range, but if the shield  301  is too high; it will collimate events for each crystal element within the FOV. If it is too short, the shield will not serve its purpose of attenuating events since most crystal elements will have a line of sight of the activity source  303 . The height of the shield  301  directly impacts the dimensions of the FOV. This cross sectional view of a crystal element  302  and associated raised shields  301  in  FIG. 3  are used to show how a shadow is created from narrow beams  305  with a line of sight to the crystal element  302 . Narrow beams  304  have a trajectory through the attenuation shield  301  and are not in the line of sight of the crystal element  202 . As photons interact with the shield material  301 , some are attenuated and the intensity of the photon beam  304  is less that than the line of sight photon beam  305 . 
         [0034]    A photon intensity difference between beams  304  and  305  casts a gamma particle shadow on the crystal element  302 . The detector is able to provide X and Y scintillation position information to the imaging system. A count of scintillations at each pixel of the crystal element  302  is accumulated. A histogram  306  is produced with each histogram bin representing a pixel of the crystal element  302 . The histogram pattern is able to indicate the bin where beam  304  shadow ends and the line of sight beam  305  begins. This location on the crystal element  302  provides an angle to the source of activity  303 . As the height or location of the activity source  303  changes, the shadow responds with a histogram pattern that provides an appropriate angle to the repositioned activity source  303 . 
         [0035]    The detector may have multiple crystal elements  402  within its array as shown in  FIG. 4 . The location of each crystal element  402  varies relative to the source of activity  403 . If a crystal element  407  and  408  has pixels that partially have a line of sight to the source  403 , a shadow line  404  is created. The angles formed for all shadow lines  404  in each crystal element with partial line of sight can indicate the height and the source of activity  403 . With all pixels of a crystal element  406  having line of sight for the source of activity  403 , the crystal element  406  acts as a collimator for the incident photons. This collimator  406  can provide X and Y positioning, the height, as well as the size of the activity source  403 . 
         [0036]    Identifying the location of activity source  403  in  FIG. 4  can be done if it is located with the FOV  405 . The FOV  405  dimensions are dependent on the height of the attenuation shield  401 , the width of the crystal element  402 , and the size and geometry of the detector array. 
         [0037]      FIG. 3  and  FIG. 4  show a two dimensional representation of the shadow effect created with a two dimensional histogram shown. The detector is able to provide a three dimensional histogram, where the histogram pattern is reviewed along multiple axes that the detector geometry can support.  FIG. 5   a  and  FIG. 5   b  show embodiments of a top view of two detector array geometries that could be constructed.  FIG. 5   a  shows hexagonal geometry, such that the histogram pattern can be reviewed in three axes  503 . The square geometry shown in  FIG. 5   b  can support a histogram pattern review on two axes  506 . 
         [0038]    The geometry of the detector is not limited to a hexagon or a square and can be in a variety of shapes such as octagon, rectangle, or circle. The shape of the attenuation shield  501  and  504  does not have to match the shape of the pixelated crystals  502  and  505 . 
         [0039]    With two dimensional histogram data captured, the process for determining the location of a source of activity for the X and Y position can be done with from collimated crystal elements. This process is similar to that of a gamma camera. The process for determining height location of a source of activity is shown in  FIG. 6 . With the two dimensional histogram data, a pattern is reviewed based on an axis and row associated with the geometry of the detector array. The source of activity location will generate peak and collimated bin counts for crystal elements below the activity. The dimensions of the source of activity will determine the quantity and location of collimated crystal elements detected. Non-collimated crystal elements next to the collimated elements identified can provide shadow line information that represents the angle to the source of activity. From a two dimensional histogram perspective, shadow line angles can provide a three dimensional representation for the height of the source of activity above the detector surface. The process in flow chart  FIG. 6  determines the peak bin count from the histogram. It then determines if the histogram bins for the element are collimated or non-collimated. If the peak element has collimated bin counts, an axis and row is determined for histogram pattern review. The bin count pattern is reviewed along the axis and row for each element and determining if the element is collimated or non-collimated. If it is non-collimated and a shadow line exists, the angle is calculated for the line toward the source of activity. The review along a row stops if no shadow line exists and the process begins in the next row along the predetermined axis. When all rows have been reviewed within the predetermined axis, the review is shifted to the next axis and the process is conducted again along the rows until all axes have been reviewed. The review can produce enough angles and collimated crystal elements to reconstruct the X, Y and Z positioning of the source of activity. To detect for multiple sources, the histogram is reviewed for additional collimated peak elements. The review can produce enough angles and collimated crystal elements to reconstruct the X, Y and Z positioning of the multiple sources of activity. 
         [0040]    Although this invention has been described with respect to specific embodiments, it is not intended to be limited thereto and various modifications which will become apparent to the person of ordinary skill in the art are intended to fall within the spirit and scope of the invention as described herein taken in conjunction with the accompanying drawings and the appended claim.