Patent Abstract:
An apparatus and method for obtaining a high-resolution digital image of an object or objects irradiated with radiation having a wavelength in the x-ray or gamma ray spectrum generated by a radiation source, or of an object or objects emitting radiation within the x-ray or gamma ray spectrum. The apparatus comprises a detector matrix and a radiation mask. The detector matrix comprises a plurality of detector pixels, each comprising a detection surface having a respective surface area which generates a signal in response to an energy stimulus. The radiation mask has an opaque portion, and a plurality of apertures. The aperture size and position relative to the detector array determines the image resolution not the size of the detector pixels. The mask is positioned between the detector matrix and the radiation source, such that the opaque portion prevents portions of the radiation from passing through the mask, and each of the apertures permits a portion of the radiation which has passed through or has been emitted from a respective portion of the object to propagate onto an area of the detection surface, less than the surface area, of a respective one of the detector pixels. The signal from a large detector pixel or from a group of small detector pixels represent an image of the respective portion of the object. The detector matrix and radiation mask are moved in synchronism in relation to the object to enable the areas of the detection surfaces of the detector pixels to receive portions of the radiation propagating through or emitted from other portions of the object, and to output signals representative of those other portions. These steps of moving the detector pixels and mask and irradiating the object are repeated until digital images of all portions of the object have been obtained. Alternatively, the x-ray source can be moved to image all portions of the object. The images are then arranged into an image representative of the entire object.

Full Description:
CROSS-REFERENCE TO RELATED APPLICATIONS 
     Related subject matter is disclosed in a U.S. patent application of Cha-Mei Tang entitled “A Method and Apparatus for Making Large Area Two-Dimensional Grids”, Ser. No. 08/879,258, filed on Jun. 19, 1997, issued as U.S. Pat. No. 5,949,850 on Sep. 7, 1999, the entire contents of which is expressly incorporated herein by reference. 
    
    
     BACKGROUND OF THE INVENTION 
     1. Field of the Invention 
     The present invention relates to a method and apparatus employing detector pixels for obtaining an image having a resolution which is not directly related to the sizes of the detector pixels. More particularly, the present invention relates to a method and apparatus which obtains a series of spatially filtered high-resolution digital x-ray or gamma ray images of portions of an object or objects while minimizing image degradation due to conversion blurring and radiation scattering, and which arranges the spatially modulated images into a larger complete image of the object or objects. 
     2. Description of the Related Art 
     Various techniques currently exist and many are under development for obtaining digital x-ray and gamma ray images of an object for purposes such as x-ray diagnostics, medical radiology, non-destructive testing, and so on. Known devices include line digital detectors, which obtain images along essentially one direction, and therefore must be scanned across an object to obtain sectional images of the object which can be arranged into an image of the entire object. Also known are two-dimensional digital detectors which can obtain an image of the entire object at one time, and thus can operate faster than an apparatus which includes a line detector. 
     A digital x-ray imager creates a digital image by converting received x-rays, which are used to form the image, into electrical charges, and displaying the charge as a function of position. Digital x-ray detectors typically have the potential of high sensitivity and large dynamic range. Therefore, when used in medical applications, a digital x-ray detector will generally be capable of obtaining a suitable image of the patient without requiring the patient to receive a large dose of x-ray radiation. 
     Digital image data is also much easier to store, retrieve and transmit over communication networks, and is better suited for computer-aided diagnostics, than conventional film x-rays. Digital x-ray images can also be displayed more easily than conventional film x-rays, and provide greater image enhancement capabilities, a faster data acquisition rate, and simplified data archival over conventional film x-rays. These advantages make digital x-ray imaging apparatus more desirable than film x-ray apparatus for use in many diagnostic radiology applications, such as mammography. 
     The general construction and operation of digital x-ray detectors will now be described. As discussed briefly above, digital x-ray detectors collect electrical charges produced by x-rays as a function of position, where the amount of charge is directly proportional to the x-ray intensity. Two general approaches for x-ray conversion are currently under investigation for flat-panel digital x-ray detectors. These approaches are generally referred to as the indirect method and the direct method. 
     In the indirect method, x-rays are converted to low-energy photons by a scintillator, and the low-energy photons are then converted to electrical charges by solid-state detectors. This method is described in a publication by L. E. Antonuk et al., “Signal, Noise, and Readout Considerations in the Development of Amorphous Silicon Photodiode Arrays for Radiotheraphy and Diagnostic Imaging”  Proc. SPIE  1443:108 (1991), the entire contents of which is incorporated by reference herein. 
     In the direct method, x-rays are converted to electron-hole pairs by photoconductors. An electric field applied to the photoconductor separates the electrons from the holes. This method is described in a publication by J. A. Rowlands et al. entitled “Flat Panel Detector for Digital Radiology Using Active Matrix Readout of Amorphous Selenium,”  Physics of Medical Imaging SPIE  3032: 97-108(1997), and in an article by R. Street, K. Shah, S. Ready, R. Apte, P. Bennett, M. Klugerman and Y. Dmitriyev, entitled “Large Area X-Ray Image Sensing Using a PbI hd 2  Photoconductor,”  Proc. SPIE  3336: 24-32 (1998). The entire contents of both of these papers are incorporated by reference herein. Many types of photoconductors are under development by medical imaging community. 
     A type of flat-panel, two-dimensional, digital x-ray, imager comprises a plurality of charge-coupled devices (CCDs) on a silicon substrate. The CCDs can be easily made on the silicon substrate to have a pixel pitch smaller than 10 μm×10 μm. However, because the maximum size of silicon substrates is limited, to achieve the dimensions needed for a large-area flat-panel x-ray detector, multiple wafers have to be patched together. Some of the CCD x-ray detectors are described in the following publications: F. Takasashi, et al., “Development of a High Definition Real-Time Digital Radiography System Using a 4 Million Pixels CCD Camera”,  Physics of Medical Imaging SPIE  3032: 364-375 (1997); J. M. Henry, Martin J. Yaffe and T. O. Tumer, “Noise in Hybrid Photodiode Array—CCD X-ray Image Detectors for Digital Mammography,”  Proc. SPIE  2708: 106 (1996); and M. P. Andre, B. A. Spivey, J. Tran, P. J. Martin and C. M. Kimme-Smith, “Small-Field Image-Stitching Approach to Full-View Digital Mammography,” Radiology 193, Suppl. Nov.-Dec., 253-253 (1994), the entire contents of each being incorporated by reference herein. 
     Alternatively, a flat-panel imager can include active matrix arrays of thin film transistors (TFTs) on a glass substrate. Because glass substrates can be large, the digital x-ray imager can, in principle, be made of a single substrate. However, it is very difficult to make a digital detector with a pixel pitch much smaller than 100 μm using substrates other than silicon wafers, as described in the following publications: L. E. Antonuk et al., “Development of Thin-Film, Flat-Panel Arrays for Diagnostic and Radiotherapy Imaging”,  Proc. SPIE  1651: 94 (1992); L. E. Antonuk et al., “Large Area, Flat-Panel, Amorphous Silicon Imagers”,  Proc SPIE  2432: 216 (1995); and L. E. Antonuk et al., “A Large-Area, 97 μm Pitch, Indirect-Detection, Active Matrix Flat-Panel Imager (AMFPI)”,  SPIE Medical Imaging  1998  Technical Abstracts,  San Diego, 83 (1998), the entire contents of each being incorporated by reference herein. 
     As discussed above, digital x-ray imaging techniques represent a vast improvement over conventional film x-ray apparatus. However, digital x-ray imaging systems experience certain drawbacks with regard to image resolution. 
     It has been a common belief that the resolution of the digital image can be no better than the pixel pitch (pixel periodicity) of the imaging apparatus, and is rather often much worse due to various types of blurring phenomena which occur during image acquisition. However, as can be appreciated from the description of the operation of digital x-ray detectors set forth below, pixel pitch is only one of the many factors that influence the resolution of a digital image obtainable by a digital imaging apparatus. 
     Detectors for digital radiography are composed of discrete pixels which generally have a uniform size, shape and spacing. The “fill factor” is defined as the active portion of each detector pixel that is used for charge collection relative to pixel pitch or, in other words, the fraction of the pixel area occupied by the sensor for x-ray detection. A flat-panel imager having thin-film transistors (TFTs), for example, has a fill factor which decreases dramatically as the pixel pitch decreases. The TFTs are large compared to transistors on silicon substrates, and the various electrode lines occupy much surface area of the glass substrate. Hence, the fill factor decreases greatly as the pixel pitch decreases. 
     For example, the fill factor is 57% for a 127 μm pixel pitch array, and is 45% for a 97 μm pixel pitch array which performs indirect x-ray conversion and has been aggressively designed, as described in the article entitled “A Large-Area, 97 μm Pitch, Indirect-Detection, Active Matrix Flat-Panel Imager (AMFPI)” cited above. 
     The fill factor approaches zero as the pixel pitch decreases toward 50 μm in a detector employing indirect converters. When the fill factor is small, the sensitivity of the detector suffers greatly. Fortunately, however, the fill factor can be improved using direct x-ray converters and a vertical stacking architecture. However, such device becomes increasingly difficult to fabricate as pixel pitch decreases. Thus, development costs for such a device are very high, and it is unclear what the smallest achievable pixel pitch could be with this technique. 
     In addition, connecting the data and control lines from the detectors to the gate driver chips and readout amplifiers of the pixel array presents severe packaging problems. Currently, bonding of large array of leads from substrate to cable is limited to a device having no less than about an 80-100 μm pixel pitch. By increasing the pixel resolution, multiplexed contacts or new bonding techniques must be developed to create input and output terminals for the device. 
     The modulation transfer function (MTF), which is a function of spatial frequency f versus location on the detector, is useful for analyzing spatial resolution. Larger MTF values mean better resolution. For existing flat-panel detectors, MTFs are important in analyzing two steps of the image acquisition sequence: the detector pixel pitch, and the blurring produced during the conversion of x-rays to charges. (See, e.g., an article by J. M. Henry, Martin J. Yaffe and T. O. Tumer, “Noise in Hybrid Photodiode Array—CCD X-ray Image Detectors for Digital Mammography,”  Proc. SPIE  2708:106(1996), the entire contents of which is incorporated by reference herein). 
     The charges generated by x-ray conversion can become blurred spatially. The source of blurring for indirect conversion using phosphor is different from that for direct conversion. For most detectors, the measured MTF is dominated primarily by the blurring of the converter when the pixel pitch is 100 μm or smaller. 
     In addition, settled phosphor scatters light generated by the x-rays. The lateral spreading of the light is approximately equal to the thickness of the layer. For settled phosphor, spatial resolution becomes finer, but the quantum efficiency decreases as the thickness of the phosphor decreases. Optimized thin photoconductors are expected to produce smaller spread. Although the light spread may be less of a problem for thick collimated CsI phosphor, the boundaries of the CsI grains are not perfect. 
     Furthermore, spatial resolution can be degraded due to x-rays striking the detector at an oblique angle. This problem exists for both direct and indirect x-ray converters. The extent of the charge spread collected by the detector is a function of the incidence angle. Since the x-ray incidence angle is a function of location on the detector relative to the x-ray point source, the modulation transfer function (MTF) of conversion blurring and oblique x-ray incidence blurring MTF conversion  is also a function of the location on the detector. The MTF conversion  of for Lanex Regular is much worse than Lanex Thin. The MTF of for Lanex Thin is 0.2, at 5 cycles/mm, as described in the article entitled “A Large-Area, 97 μm Pitch, Indirect-Detection, Active Matrix Flat-Panel Imager (AMFPI)”, cited above. 
     The final system MTF is the product of the MTF associated with various components of the system, including the detector array MTF introduced by the detector pixel pitch and the MTF of conversion blurring. For these reasons, the reduction of pixel pitch alone is not as good the combination of reduction of pixel pitch and reduction of conversion blurring. The resolution of the detector is also effected by a variety of other factors that will not be discussed in detail here such as signal statistical noise, charge conversion noise and electronic noise. 
     Gamma rays are radiation generated by nuclear process. The energy of gamma rays are typically higher than that of the x-rays, but low energy range of the gamma rays can overlap the high energy end of the x-rays. These detector concepts can also be applied to the detection of gamma rays and megavolt radiation. A thick scintillator or a metal plate/phosphor screen combination is used. This is described in a publication by L. E. Antonuk, et al., “Demonstration of Megavoltage and Diagnostic X-ray Imaging with Hydrogenated Amorphous Silicon Arrays,”  Med. Phys.  19: 1455 (1992), the entire contents of which is incorporated by reference herein. 
     In summary, the major problems expected with small pixel detector development are complicated circuit architecture, increased number of leads to be bonded, the small pitch of the leads necessary for bonding, and resolution being increasingly dominated by scintillator blurring and the oblique x-ray incidence effect. These drawbacks result in decreased manufacturing yield, high risk and expensive development. 
     Accordingly, a continuing need exists for an apparatus capable of obtaining high-resolution digital x-ray or gamma ray images without the drawbacks discussed above. 
     SUMMARY OF THE INVENTION 
     An object of the present invention is to provide a method and apparatus for obtaining high-resolution digital x-ray or gamma ray images of an object or objects emitting x-rays or gamma rays, or of an object or objects irradiated with radiation having a wavelength within the x-ray or gamma ray spectrum. 
     Another object of the present invention is to obtain digital x-ray or gamma ray images at a resolution better than the pixel pitch of the detectors used to obtain the digital images. 
     Another object of the present invention is to reduce scattered x-rays or gamma rays detected by the digital detector while also improving image resolution. 
     A further object of the present invention is to minimize blurring of the digital x-ray or gamma ray images which can occur when the x-rays or gamma rays are directly converted into electron-hole pairs in a photoconductor and collected by the active area of the digital detectors. 
     A still further object of the present invention is to minimize blurring of the digital x-ray or gamma ray image which occurs when the x-rays or gamma rays are indirectly converted into electric charges first by converting x-rays or gamma rays to a longer wavelength radiation, for example, optical radiation, and then collecting and converting these radiation and converting them to electrical charge. 
     These and other objects of the present invention are substantially achieved by providing an apparatus and method for obtaining a digital image of an object or objects generating x-rays or gamma rays, or of object or objects irradiated with radiation having a wavelength in the x-ray or gamma ray spectrum generated by a radiation source. The apparatus comprises a detector matrix and a radiation mask. The detector matrix comprises a plurality of two-dimensional array of detector pixels, each of which comprises a detection surface having a respective active surface area and being adapted to generate an electrical signal in response to a radiation stimulus applied thereto. The radiation mask has an opaque portion and a plurality of apertures therein. The mask is positioned between the detector matrix and the radiation source. The radiation can pass through the mask to the detector only through the apertures of the mask. The image resolution is related to the aperture size and system configuration. Many modes of operation of this detector system are described below. 
     In the first mode of operation, the detector images object or objects that give radiation. The mask is placed between the object and the active detector pixels. The mask allows radiation from selected portions of the objects to be imaged by the detector for a single imaging frame. 
     In the second mode of operation, the object or objects are placed between a radiation source and the mask. Again, the mask allows a selected portion of the object or objects to be imaged by the detector for a single image frame. 
     In the third mode of operation, the object or objects are placed between the mask and the detector array, such that the opaque portion of the mask prevents portions of the radiation from passing therethrough, and each of the apertures permits a portion of the radiation which has passed through a respective portion of the object or objects to pass therethrough and propagate onto an active area of the detection surface of a respective one of the detector pixels. The detector pixels therefore each output a respective signal of the respective portion of the object. 
     The imaging apparatus further includes a conveying device which moves the detector matrix and radiation mask in unison in relation to the object to enable the areas of the detection surfaces of the detector pixels to receive portions of the radiation propagating through other portions of the object, and to output signals representative of those other portions. In particular, the detector matrix and radiation mask are moved along a pattern of movement in increments which are a fraction of the pixel pitch of the detector pixels. After each exposure of the detector to the radiation source, the charges collected by the detector array are read out to a computer and the detector array is reinitialized and the detector and mask are moved to the next appropriate position. This process is repeated so those portions of the object or objects which would not normally be imaged by this detector in the stationary mode can be imaged. These steps of moving the detector pixels and mask, and irradiating the object, are repeated until digital images of all portions of the object or objects have been obtained. The digital data are then arranged into an image representative of the entire object or objects. 
    
    
     BRIEF DESCRIPTION OF THE DRAWINGS 
     The various objects, advantages and novel features of the present invention will be more readily appreciated from the following detailed description when read in conjunction with the accompanying drawings, in which: 
     FIG. 1 is a schematic side view illustration of a high-resolution x-ray or gamma ray imaging apparatus according to an embodiment of the present invention; 
     FIG. 2 is a schematic illustration of the high-resolution imaging apparatus shown in FIG. 1 in relation to an object being imaged and a point x-ray or gamma ray source; 
     FIGS. 3 a  and  3   b  are schematic illustrations showing the scattering of light generated in phosphor screens by incident x-ray energy in relation to the thickness of the phosphor screens which can be employed to perform x-ray conversion in the imaging apparatus shown in FIGS. 1 and 2; 
     FIG. 4 is a schematic illustration showing charge smear generated in a photoconductor, which can be employed to perform x-ray conversion in the imaging apparatus shown in FIGS. 1 and 2, in relation to various angles of incidence of x-ray energy striking the photoconductor; 
     FIG. 5 is a schematic illustration of a top plan view of a mask which can be employed in the imaging apparatus shown in FIGS. 1 and 2; 
     FIG. 6 is a schematic top plan view of an example of a detector pixel array which can be employed in the imaging apparatus shown in FIGS. 1 and 2; 
     FIG. 7 a  is a schematic illustration showing the pattern of electromagnetic radiation which passes through the mask shown in FIG.  5  and strikes the scintillator adjacent the active area of the detector pixels of the detector pixel array shown in FIG. 6; 
     FIG. 7 b  is a diagram illustrating an exemplary sequence of movements of the detector pixel array shown in FIG.  6  and the mask shown in FIG. 5 of the imaging apparatus shown in FIGS. 1 and 2 with respect to the object being imaged according to an embodiment of the present invention; 
     FIG. 8 is a schematic top plan view of another example of a mask which can be employed in the imaging system shown in FIGS. 1 and 2; 
     FIG. 9 a  is a schematic illustration showing the pattern of electromagnetic radiation which passes through the mask shown in FIG.  8  and strikes the scintillator adjacent the active area of the detector pixels of the detector pixel array shown in FIG. 6; 
     FIG. 9 b  is a diagram illustrating an exemplary sequence of movements of the detector pixel array shown in FIG.  6  and the mask shown in FIG. 8 of the imaging apparatus shown in FIGS. 1 and 2 with respect to the object being imaged according to an embodiment of the present invention; 
     FIG. 10 is another diagram illustrating an exemplary sequence of movements of the detector pixel array shown in FIG.  6  and the mask shown in FIG. 5 of the imaging apparatus shown in FIGS. 1 and 2 with respect to the object being imaged according to an embodiment of the present invention; 
     FIG. 11 is a schematic top plan view illustration of another example of a mask which can be employed in the imaging system shown in FIGS. 1 and 2; 
     FIG. 12 a  is a schematic showing the pattern of electromagnetic radiation which passes through the mask shown in FIG.  11  and strikes the scintillator adjacent the active area of the detector pixels of the detector pixel array shown in FIG. 6; 
     FIG. 12 b  is a diagram illustrating an exemplary sequence of movements of the detector pixel array shown in FIG.  6  and the mask shown in FIG. 11 of the imaging apparatus shown in FIGS. 1 and 2 with respect to the object being imaged according to an embodiment of the present invention; 
     FIG. 13 is a schematic top plan view of another example of a detector pixel array which can be employed in the imaging apparatus shown in FIGS. 1 and 2; 
     FIG. 14 is a schematic top plan view of another example of a mask which can be employed in the imaging system shown in FIGS. 1 and 2; 
     FIG. 15 a  is a schematic illustration showing the pattern of electromagnetic radiation which passes through the mask shown in FIG.  14  and strikes the scintillator adjacent the active area of the detector pixels of the detector pixel array shown in FIG. 13; 
     FIG. 15 b  is a diagram illustrating an exemplary sequence of movements of the detector pixel array shown in FIG.  13  and the mask shown in FIG. 14 of the imaging apparatus shown in FIGS. 1 and 2 with respect to the object being imaged according to an embodiment of the present invention; 
     FIG. 16 is a schematic top plan view of another example of a detector pixel array which can be employed in the imaging apparatus shown in FIGS. 1 and 2; 
     FIG. 17 is a schematic top plan view of another example of a mask which can be employed in the imaging system shown in FIGS. 1 and 2; 
     FIG. 18 a  is a schematic illustration showing the pattern of electromagnetic radiation which passes through the mask shown in FIG.  17  and strikes the scintillator adjacent the active area of the detector pixels of the detector pixel array shown in FIG. 16; 
     FIG. 18 b  is a diagram illustrating an exemplary sequence of movements of the detector pixel array shown in FIG.  16  and the mask shown in FIG. 17 of the imaging apparatus shown in FIGS. 1 and 2 with respect to the object being imaged according to an embodiment of the present invention; 
     FIG. 19 is a schematic top plan view of another example of a detector pixel array which can be employed in the imaging apparatus shown in FIGS. 1 and 2; 
     FIG. 20 is a schematic top plan view of another example of a mask which can be employed in the imaging system shown in FIGS. 1 and 2; 
     FIG. 21 a  is a schematic illustration showing the pattern of electromagnetic radiation which passes through the mask shown in FIG.  20  and strikes the scintillator adjacent the active area of the detector pixels of the detector pixel array shown in FIG. 19; 
     FIG. 21 b  is a diagram illustrating an exemplary sequence of movements of the detector pixel array shown in FIG.  19  and the mask shown in FIG. 20 of the imaging apparatus shown in FIGS. 1 and 2 with respect to the object being imaged according to an embodiment of the present invention; 
     FIG. 22 is a schematic illustration of a high-resolution x-ray or gamma ray imaging apparatus according to another embodiment of the present invention in relation to an object being imaged and a point x-ray or gamma ray source; 
     FIG. 23 a  is a schematic top plan view of an example of a mask which can be employed in the imaging system shown in FIG. 22; 
     FIG. 23 b  is a diagram illustrating an exemplary sequence of movements of the mask shown in FIG. 23 a  of the imaging apparatus shown in FIG. 22 with respect to the object being imaged according to an embodiment of the present invention; 
     FIG. 24 a  is a schematic illustration of a high-resolution x-ray or gamma ray imaging apparatus according to another embodiment of the present invention in relation to an object being imaged and a point x-ray or gamma ray source; 
     FIG. 24 b  is a diagram illustrating an exemplary pattern of movement of the x-ray source of the apparatus shown in FIG. 24 a  with respect to the object being imaged according to an embodiment of the present invention; 
     FIGS. 25 a,    25   b  and  25   c  are schematic cross-sectional views of examples of masks which can be employed in an imaging apparatus as shown in FIGS. 1,  2 ,  22 ,  24   a,    26 ,  27  or  28 , when the imaging apparatus is used with a point x-ray source; 
     FIGS. 25 d  and  25   e  are schematic cross-sectional views of examples of masks which can be employed in an imaging apparatus as shown in FIGS. 1,  22 ,  27  or  28 , when the imaging apparatus is used with a parallel beam x-ray source; 
     FIG. 26 is a schematic illustration of a high-resolution x-ray or gamma ray imaging apparatus according to another embodiment of the present invention in relation to an object being imaged and a point x-ray or gamma ray source; 
     FIG. 27 is a schematic illustration of a high-resolution x-ray or gamma ray imaging apparatus according to a further embodiment of the present invention in relation to an object being imaged and a point x-ray or gamma ray source; 
     FIG. 28 is a schematic illustration of a high-resolution x-ray or gamma ray imaging apparatus according to still another embodiment of the present invention in relation to an object being imaged and a point x-ray or gamma ray source; 
     FIG. 29 is a schematic illustration of a detector array such as a charge coupled device (CCD); and 
     FIG. 30 is a schematic illustration showing the pattern of electromagnetic radiation which passes through the mask and strikes the scintillator adjacent the active area of the detector pixels of the detector pixel array shown in FIG.  29 . 
    
    
     DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS 
     An embodiment of a high-resolution x-ray or gamma ray imaging apparatus  100  is exemplified in FIGS. 1-7 b.  In particular, FIG. 1 is a schematic diagram illustrating a view of a side of the imaging apparatus  100  lying in the x-z plane. The imaging apparatus  100  includes a substrate  102 , which can be a silicon or glass substrate or any other appropriate material as described in the Background section above, a detector pixel array  103  with detector pixels  104  which are disposed on the substrate  102 , and a scintillator  106 . The active area of the detector pixels  104  can be any type of pixel as described in the Background section above. 
     In this embodiment, the scintillator  106  converts x-rays or gamma rays to electron-hole pairs or visible photons. The electron hole pairs or visible photons are converted to electrical charge, current or voltage collected on the active radiation detector area of the pixel  104 . In the typical digital x-ray or gamma ray detectors and visible imagers, the active area of the detector pixels  104  each measure the amount of charge collected per pixel. In general, the active area of the detector pixel  104  measures the change of electrical properties, material properties, physical properties, and so on, produced by the variation of the electromagnetic radiation intensity on the active area of the detector pixel  104 . 
     A mask or mask/antiscatter grid  108  (hereinafter “mask  108 ”) having aperture openings  110  therein is disposed on the upper surface of the scintillator  106 . Each aperture opening  110  is aligned with a corresponding active area of the detector pixel  104  as shown. For many applications, the mask  108  can be rigidly attached to the scintillator  106 , or can be directly attached to the active area of the detector pixels  104 . The mask  108  must be opaque enough to substantially block the penetration of the electromagnetic radiation except through the aperture openings  110 . 
     The active area of each detector pixel  104  is larger than its respective aperture opening  110 , and detects the electromagnetic radiation (x-rays or gamma rays) passing through its respective aperture opening  110 . As discussed below, the size of the aperture openings  110  and the number of images taken, not the detector pixel pitch, determines the image resolution. 
     The detector shown in FIG. 1 can be used to image objects that radiate x-rays or gamma rays. For example, the detector can be used for x-ray astronomy. 
     FIG. 2 is a schematic drawing illustrating a side view of the embodiment of the imaging apparatus  100  shown in FIG. 1 being used in an x-ray radiography application to image the interior of an object  112 , which can be, for example, a human body (or a portion thereof) or any other object. An x-ray source  114  is also illustrated schematically. Also, the source  114  could be a gamma ray source, or any energy source. 
     As shown, the object  112  to be imaged is positioned between the x-ray source  114  and the x-ray mask  108  of the imaging apparatus  100 . After the x-ray source  114  emits a pulse of x-rays and the x-rays penetrate the object  112 , the x-rays reach the mask  108 . The mask  108  blocks all the x-rays from hitting the scintillator  106  except at the mask openings  110 . 
     The scintillator  106  can be a phosphor screen, which converts the x-rays to optical radiation, and the photodiodes on each detector  104  covert the optical radiation to electrical charge. Alternatively, the scintillator  106  can be of the type that converts the x-rays directly to charge, such as a photoconductor, photocathode, or the like. The geometry and dimensions of the active area of the detector pixels  104  and x-ray mask openings  110  are such that the x-rays passing through a single mask or mask/antiscatter grid opening  110  will strike preferably only a single detector pixel  104 . Preferably, the active detector area of one pixel  104  captures the charges created by one x-ray beamlet. The charge collected per pixel is then output via data lines (see FIG.  6 ), and processed in a manner known in the art. 
     The arrangement of the imaging apparatus  100  will improve the detector system MTF and increase the Nyquist frequency of even the existing best known detector pixels arrays to obtain a resolution much higher than that obtained by the same detector without a mask and without motion. The detector system MTF is the product of MTF associated with various component of the detector. Two MTF will be discussed: MTF associated with detector geometry and MTF associated with x-ray conversion. 
     As will now be explained, the operation of the imaging apparatus  100  will improve MTF associated with the detector system geometry for detectors which perform either direct or indirect conversion of the x-rays or gamma rays as discussed above. 
     FIGS. 3 a  and  3   b  are schematic diagrams illustrating the manner in which phosphor screens scatter the light generated by the x-rays during indirect x-ray conversion. As shown, the light scatter is proportional to the thickness of the phosphor screen. A thicker phosphor screen will provide a greater light scatter. 
     FIG. 4 is a schematic diagrams illustrating that for direct conversion of x-rays, charge smear is minimal when the x-ray incidence angle is zero degrees, and increases as the x-ray incidence angle increases. For both of these situations, an active pixel detector area much larger than the x-ray mask aperture will reduce conversion blurring and improve conversion MTF. 
     The active area of the detector pixels  104  and mask  108  can have a wide range of pattern or layout. For example, FIG. 5 is a schematic diagram of mask  108  of the imaging apparatus, with apertures  110  viewed in the x-y plane in FIG.  1 . The apertures  110  are square or essentially square, and each have a length and width equal to d 1 . The area of each aperture is d 1 ×d 1 , and the pitch of the aperture is equal to the pixel pitch D 1  in both directions. The arrangement of the apertures  110  forms a uniform grid of openings in the mask  108 . As discussed above, the electromagnetic radiation to be detected has to be completely blocked by the mask  108  except at apertures  110  in the mask  108 . The apertures  110  are used to control the area and position at which the electromagnetic radiation hits the detector pixels. 
     In this embodiment, the pixel pitch D 1  is an integer multiple of d 1 . To enable the object to be  112  imaged without missing any areas and without double-exposing any areas, the imaging apparatus  100  is configured and operated so that the beamlets will each “fit” into a respective active area of the detector pixel  104  an exact number of times. In other words, D 1 =nd 1 , and n is an integer equal to or greater than 2. FIG. 5 shown an aperture arrangement where D 1 =2d 1 . 
     FIG. 6 is a generalized schematic illustration of a top view of a possible layout of the detector pixel array  103  and the active area of the detector pixels  104  for the imaging apparatus  100  as shown in FIGS. 1 and 2. The active radiation detector areas of the pixels  104  are shown shaded with hatched lines. It is noted that the dimensions of the active area of the detector pixels  104  vary greatly from one manufacturer to another, and that the shapes of the active radiation detector areas of the pixels  104  can vary widely and are represented as squares only for illustration purposes. Row control (selection) lines  116 , which are disposed on the substrate  102  (see FIGS.  1  and  2 ), are spaced uniformly from each other at the distance D 1  as shown. Column data lines  118 , which are also disposed on substrate  102 , are also spaced uniformly from each other at the distance D 1 . Typically, data is read out one row at a time (but could be more than one row at a time) through the column data lines  118  to a processing device, such as a computer  119  or the like, as controlled by the row control lines  116 . 
     FIG. 7 a  is a schematic representation of the radiation beamlets  120  that pass through the apertures  110  of the mask  108  which has been superimposed over the active area of the detector pixels  104 . Specifically, the electromagnetic radiation beamlets  120  are illustrated as white squares on the pixels  104 , with each white square having a dimension d 1 ×d 1 , which is equal to or essentially equal to the dimension of the aperture  110  through which the beamlet  120  has passed. In summary, as shown in FIG. 7 a,  the radiation beamlets  120  hit the scintillator above the active area of the detector pixels  104  with dimension d 1 ×d 1 . The distance between the centers of adjacent apertures  110  is equal to D 1 , which is the pitch of the active area of the detector pixels  104 . The relationship between the dimensions of each active area of the detector pixel and the dimensions of the radiation beamlets when they hit the detector pixel is D 1 =nd 1 , where n=2 in this example. Also, the x-rays are only allowed to impact the detector during the x-ray exposure time, but not during the data read out time or while the mask or detector is being moved. 
     To assure that the entire object  112  (FIG. 2) is imaged, a conveying device  124  (see FIG.  1 ), such as a stepper motor, servo motor, motorized table, or any other suitable device, is configured to move the imaging apparatus  100  in a controlled manner. The imaging apparatus  100  is moved with respect to the object  112  in increments equal to d 1  along the pattern shown in FIG. 7 b.  That is, after one exposure of the object  112  to the x-rays, a x-ray image of a respective portion of the object  112  is obtained by each pixel  104 . The data produced by the pixels  104  is output through the column data lines  118 . The imaging apparatus  100  is then moved in the x-y plane by a distance d 1  along an arrow in FIG. 7 b.    
     This process is repeated n 2  times with the imaging apparatus  100  (i.e., the detector pixels grid  103 , scintillator  106  and mask  108 ) moved systematically in the x-y plane, for example, in the directions along arrows  126 ,  128 ,  130  and  132  for each exposure and reading, so that every part of the object  112  is imaged. After all four x-ray image patterns (n 2 =4 in this example) have been obtained and stored, they are reconstructed by a processing device, such as the computer  119  or the like into a complete image representative of the entire object  112 . The reconstructed image has higher resolution than any single x-ray image pattern obtained with or without the mask  106 . 
     The principle of improvement of image resolution is explained first assuming no x-ray conversion blurring and then expanded to include x-ray conversion blurring. 
     For the fill factor of the active area of the detector is 100%, 
     
       
           MTF   geometry =sin(π fD )/(π fD ), 
       
     
      Where MTF geometry  is the MTF associated with the geometry of the detector system in one direction, D is the dimension of the pixel pitch, and f is the spatial frequency. The Nyquist frequency is 1/2D. 
     When the linear dimension of the active area of the detector pixel is reduced to d 1 , for D=2d 1 , 
     
       
           MTF   geometry =sin(π f ( d   1 ))/(π f ( d   1 )), 
       
     
      and the Nyquist frequency is still 1/2D. 
     When the linear dimension of the active area of the detector is d 1  and D=2(d 1 ), and the detector is moved as shown in FIG. 7 b  and D=2(d 1 ), then 
     
       
           MTF   geometry =sin(π f ( d   1 ))/ (π f ( d   1 )), 
       
     
      and the Nyquist frequency is increased to 1/4D. This technique is used to reduce aliasing and improve image resolution for infrared cameras. The technique is called microscanning, dithering and microdithering, as described in the following publications: J. C. Gillette, T. M. Stadtmiller and R. C. Hardie, “Aliasing reduction in staring infrared imagers utilizing subpixel techniques,” Optical Engineering 34, 3130-3137 (1995); R. C. Hardie, K. J. Barnard, J. G. Bognar, E. E. Armstrong and E. A. Watson, “High-resolution image reconstruction from a sequence of rotated and translated frames and its application to an infrared imaging system,” Optical Engineering 37, 247-260 (1998), the entire contents of each being incorporated by reference herein. 
     For x-ray and gamma ray imaging, there is conversion blurring. Conversion blurring can eliminate the benefits of microscan without mask and significantly reduce the signal. For example, for a TFT digital x-ray detector having an active area of the pixel with a dimension d 1 ×d 1 , if If N number of x-rays impinges on this active area of the pixel and M number of electrons are created per x-ray, then the total number of electrons created per pixel would be MN. When there is no conversion blurring, the total number of charge collected by this pixel would be MN. Due to conversion blurring, the percentage of charge collected by this pixel decreases as the pixel dimension decreases, and the remaining charges are spread to the neighboring pixels. 
     In the detector system of the present invention as shown, for example, in FIGS. 1-2, the aperture size of the mask determines the Nyquist frequency and the MTF associated with the pixel, while the active area of the pixel is kept large to increase the percentage of charge collected as the aperture of the mask decreases. 
     The small aperature of the mask and large detector pixel size also improves the MTF associated with the conversion blurring, MTF conversion . The detector system MTF, MTF system , is the product of the MTF associated with the various aspects of the system, 
     
       
           MTF   system   =MTF   geometry   *MTF   conversion   *MTF   others , 
       
     
     Where MTF others  is the MTF associated with other component of the detector system. 
     The detector system described in FIGS. 1-2 and  5 - 7   b  with a mask and motion has a higher Nyquist frequency, larger values for the MTF within the Nyquist frequency and improve signal as compared to imaging without the mask and motion. As explained above, the detector pixel array  103  and mask  108  arrangement can have a wide variation of patterns and dimensions. For example, FIG. 8 is a schematic of a top view of a mask  134  which can be used in the imaging apparatus  100  shown in FIGS. 1 and 2 instead of mask  108 . Mask  134  includes apertures  136  which are square or essentially square and have a dimension d 2 ×d 2 , such that the pixel pitch D 1  of detector pixels  104  is equal to 3(d 2 ), (D 1 =3(d 2 )) in both directions. 
     FIG. 9 a  is a schematic view showing the electromagnetic radiation that has passed through the mask  134  and has impacted on the scintillator above detector pixels  104 . That is, the x-ray beamlets  138  pass through respective apertures  136  in the mask  134  and strike the center of the active radiation detection area of the respective pixel  104 . 
     To obtain an entire x-ray image of the object  112  with an imaging apparatus  100  including mask  134 , the imaging apparatus  100  is moved along a pattern as shown, for example, in FIG. 9 b.  That is, as discussed above with regard to FIGS. 7 a  and  7   b,  after each exposure of the object  112  to x-rays and, generation of an x-ray image sub-pattern by the pixels  104 , and read-out of the pixel data through column data lines  118 , the imaging apparatus  100  is moved to a new location. The imaging apparatus  100  is moved sequentially each time an x-ray image is taken, and is moved in a possible pattern shown in FIG. 9 b  with each arrow representing one successive movement(d 2 =D/3). This process is repeated n 2 =9 times with the detector  103 , scintillator  106  and mask  134  moved in unison so that every part of the object  112  will be imaged. After all of the x-ray image sub-patterns have been obtained and stored, they are combined by a processor such as a computer or the like to provide an x-ray image representative of the entire object  112 . 
     In addition, aliasing can be further minimized and MTF improved by oversampling and applying appropriate mathematical algorithms. That is, returning to the example discussed with regard to FIGS. 7 a  and  7   b,  instead of moving the imaging apparatus  100  including detector  108  by a distance d 1  between successive x-ray or gamma ray exposures, the imaging apparatus  100  is moved by a distance of (d 1 )/2=(D 1 )/2n, so the total number of sub-frames required is (2n) 2 . The value (d 1 )/2=(D 1 )/2n. The arrows shown in the diagram of FIG. 10 suggest a possible sequence of movements for imaging apparatus  100  including detector pixels array  103 , scintillator  106  and mask  108  for a detector motion of (d 1 )/2 between exposures, with the distance d 1  being equal to one-half the pixel pitch D 1  (i.e., D 1 /d 1 =2). 
     An example of sampling variation by increasing the size of the apertures in the mask without changing the detector size or the distance between exposures is exemplified in FIGS. 11,  12   a  and  12   b.  FIG. 11 shows a mask  140  with apertures  142  each having a dimension d 3 ×d 3 , where D 1 /(n−1)&gt;d 3 &gt;D 1 /n. In this example, n=2. FIG. 12 a  shows the spot size of the radiation beamlets  144  formed by mask  140  on the scintillator above the detector pixels  104 . After each x-ray exposure and data readout operation is performed in the manner discussed above, the detector is moved a distance D 1 /n along the arrows shown in FIG. 12 b.  This process is repeated n 2  times with the detector  103 , scintillator  106  and mask  140  moving in unison so that every part of the object  112  is imaged. The suggested motion is similar to that of the example showing FIG. 10 to reduce aliasing. The aliasing reduction is dependent on the amount of overlapping image. 
     It is noted that the periodicity of the detector pixel pitch need not be square. For example, as shown in FIG. 13 shows a detector pixel array  146  having the active area of the detector pixels  148  within the D 1 ×0.75(D 1 ) pixel pitch. For some applications, a rectangular area of the detector pixel layout is more effective than a layout of square detector pixels. 
     FIG. 14 is a schematic illustration of a mask  150  having apertures  152  appropriate for the detector pixels  148  shown in FIG.  13 . In this example, n=D 1 /d 4 =3. 
     FIG. 15 a  is a schematic diagram illustrating the location of the radiation beamlets  154  passing through the apertures  152  of the mask  150  onto the scintillator above the detector pixels  148 . Preferably, the x-ray beams that pass through each aperture  152  in the mask  150  are centered on the active radiation detection area of a respective pixel  148 . After each x-ray exposure to the object  112  and data readout is performed in the manner discussed above, the imaging apparatus  100  including detector pixel array  146  and mask  150  is moved a distance (D 1 )/4 along the arrows shown in FIG. 15 b.  This process is repeated 6 times with the detector grid  146  and mask  150  moved systematically so that every part of the object  112  will be imaged. 
     FIG. 16 is a schematic of a top view of a variation in the layout of the detector pixels for the imaging apparatus  100  shown in FIGS. 1 and 2. In the pixel array  156 , the active areas for radiation detection of the pixels  158  are shown shaded with hatched lines. The shape of each pixel  158  is shown as a square for schematic purpose only. In general, the pixel shape can vary from one product to another and from one manufacturer to another. 
     As shown, the detector pixels  158  are staggered in formation. The periodicity of the pixel is 2D 1  in the horizontal direction and D 1  in the vertical direction. The arrangement further includes column data lines  160 , which are similar to the column data lines  118  discussed above and are spaced uniformly a distance D 1  apart. Each data line will be connected to all the pixels  158  in a respective column of pixels. Control lines  162  run in a staggered zigzag pattern from left to right in this embodiment, and are spaced uniformly a distance D 1  apart. 
     FIG. 17 is a schematic illustration of the aperture layout of the mask  164  employed in the imaging apparatus  100  shown in FIGS. 1 and 2 having a detector pixel layout as shown in FIG.  16 . The apertures  166  are arranged in a staggered fashion as shown, and D 1 /(d 5 )=2. 
     FIG. 18 a  is a schematic illustration showing the locations at which the radiation beamlets  168  pass thought the apertures  166  overlaying the detector pixel array  156 . FIG. 18 b  is a diagram showing an example of movement of the mask  164  and detector pixel array  156  for four x-ray exposures and returning to its original position and image readings which occur in the manner discussed above. As shown, the mask  164  and detector pixel array  156  move along the arrows by a distance d 5  between each exposure and image reading. The minimum number of exposures is n 2 , and n=2 in this example. 
     In general, there are many variations in direction and distance in which the detector pixel array  156  and mask  164  can be moved. For instance, D 1 /(d 5 ) can be any number greater than or equal to 2, and various image data sampling algorithms can be implemented. Also, the pixel pitch does not have to be square. 
     For example, FIG. 19 is another schematic illustration of a top view of a detector pixel array  170  which can be employed in imaging apparatus  100  shown in FIGS. 1 and 2 in place of detector pixel array  103 . This figure is similar to FIG. 16, except the periodicity of the pixel detectors  172  is 3(D 1 ) in the x direction. 
     FIG. 20 is a schematic illustration of a mask  174  which can be employed in an imaging apparatus  100  which includes detector pixel array  170  shown in FIG.  19 . The apertures  176  of the mask  174  are arranged in a staggered fashion along the x direction, and D 1 =3(d 6 ). 
     FIG. 21 a  is a schematic illustration of the positions at which the radiation beamlets  178  which pass through the aperture of the x-ray mask  174  overlaying the detector pixel array  170  strike the detector pixels  172  of the grid  170 . FIG. 21 b  is a diagram of an example of the manner in which the detector pixel array  170  and mask  174  are moved for nine exposures by a distance d 6  between exposures and returning to its original position. As can be appreciated from FIG. 21 b,  the staggered formation of the detector pixels grid  170  and mask  174  enable the entire object to be imaged by moving the grid  170  and mask  174  in one direction (i.e., the x direction), as opposed to in the x and y directions as from a non-staggered grid discussed above. 
     Another mask variation is that the apertures are not squares. For some applications, other x-ray aperature shapes might be more appropriate. 
     Although only several examples of masks and detector pixel array arrangements are described above, various types of mask having various apertures patterns can be used in the imaging system  100  to provide a wide variety of possible image system configurations. Also, as discussed below, the masks need not be attached to the scintillator, but rather, could be positioned at any appropriate location between the x-ray or gamma ray source and the detector pixel array. 
     For example, FIG. 22 is a schematic illustrating an embodiment of an imaging apparatus  180  which includes a substrate  182 , a detector pixel array  184  including detector pixels  186 , a scintillator  188 , and a mask  190  having apertures  191  therein similar to those described above. The imaging apparatus  180  can also include an antiscatter grid  192  which is disposed over the scintillator as shown. An example of an antiscatter grid is disclosed in related copending U.S. patent application Ser. No. 08/879,258, cited above. An x-ray source  194  and object  196  being imaged are also illustrated in relation to the apparatus  180 . 
     Unlike imaging apparatus  100 , in this embodiment the object to be imaged  196  is positioned between x-ray mask  190  and the detector pixel array  184 . As shown, the x-ray energy propagates out of a point x-ray source in a cone shape. 
     FIG. 23 a  shows the mask  190  as viewed in the x-y plane. The apertures  191  are shown as having a square shape, but could have any suitable shape as discussed above for the other masks configurations. Primarily, the size and arrangement of the apertures  191  on the mask  190  should be such that they permit uniform sized and equally spaced beamlets to form on the detector pixels  186 . 
     The periodicity of the square digital detector pixels is defined to be D 1 ×D 1 . The dimension of each x-ray beamlet as it hits the detector pixel (the “x-ray spot size”) is equal to d 7 ×d 7 , where d 7 &lt;D. Using Euclidean geometry, if the x-ray source  194  is considered to be a vertex of a triangle, the x-ray beamlet on the detector pixels  186  is the base of the triangle, and the distance between the x-ray source  194  and the detector pixel is L (distance measured orthogonally), then if the x-ray mask  190  is placed a distance αL from the x-ray source where α is a fraction less than 1, the dimensions of the apertures  191  in the x-ray mask  190  must be equal to α(d 7 )×α(d 7 ). Also, as with the variations discussed above, the apertures of the mask and the detector pixels can vary in size and shape depending on the need. 
     The operation of the imaging apparatus  180  will now be described. When the x-ray source  194  emits a pulse of x-ray energy which strikes the x-ray mask  190 , the mask blocks all of the x-rays from striking the object except at the mask apertures  191 . The x-ray beamlets which pass through the apertures of the mask penetrate the object  196  and propagate toward the antiscatter grid  192 . The antiscatter grid  192  eliminates the scattered radiation, so that only the primary radiation impacts the scintillator  188 . As in the imaging apparatus  100  shown in FIGS. 1 and 2, the scintillator  188  can be a phosphor screen, which converts the x-rays to optical radiation. A photodiode on each detector pixel coverts the optical radiation to electrical charge. Alternatively, the scintillator  188  can be of the type that converts the x-rays directly to electrical charge, such as photoconductor, photocathodes, and so on. 
     The geometry and dimensions of the detector pixels  186  and x-ray mask openings  191  are such that each x-ray beamlet passing through a respective aperture in the mask and a respective aperture in the antiscatter grid  192  will strike within a single detector pixel  186 . Preferably, the active detector area of one pixel  186  captures the charges created by the impacting x-ray beamlet. After each exposure, the x-ray source is turned off or x-ray shutter is closed. The charges collected by the pixels  186  are then output via data lines in a manner similar to that described above for imaging apparatus  100 . 
     For this example, n=D 1 /d 7 =2. After one exposure and data read out, the detector grid  184  (and hence the substrate  182 , scintillator  188  and antiscatter grid  192 ) is moved a distance D 1 /2 for n=2 in a sequence as shown in FIG. 12 b  and the x-ray mask  190  is moved by a distance αd 7  in the same sequence as shown in FIG. 23 b  while the object  196  (patient) remains stationary, to expose a different portion of the object  196 . This process is repeated n 2 times with the detector and mask moved in unison so that every part of the object will be imaged. After all the necessary sub-images have been output and stored, the data is processed to produce one image in a manner similar to that described above. Even though n 2  exposures are taken, the tissue is exposed to the same dose of x-ray as in one exposure without the mask, because each exposure is 1/n 2  the area of an exposure without the mask. The data is then reconstructed digitally to produce the high-resolution image. 
     Variations of the embodiments for the mask and the detector grid layout are the same as those exemplified in FIGS. 8 through 21, except that each aperture of the mask is reduced in size by the factor a and the motion of the mask is reduced by the same factor. 
     FIG. 24 a  is a schematic diagram illustrating that the image filtering concept can be obtained by moving the location of the x-ray source  194  without moving the mask  190 . For the detector shown in FIG.  6  and D 1 /d 7 =n=2, the detector motion is shown in FIG. 12 b,  the corresponding x-ray source displacement is shown in FIG. 24 b,  where the distance between displacement is d 8  and d 8 ≈(D 1 /n)(α/(1−α)). The direction of motion for the source, shown in FIG. 24 b,  is opposite to the direction of motion for the detector, shown in FIG. 12 b.  The range for α is between 0 and 1, and the optimal values for a are near  0 . 5 . The positions for the x-ray source  194  are such that every part of the object will be imaged. Variations of the embodiments for the mask and the detector grid layout are the same as those exemplified in FIGS. 8 through 21, except that each aperture is reduced in size by the factor α. 
     Another variation of FIG. 24 a  is to move the location of the x-ray source  194  and the x-ray mask  190 , but not move the detector  184 , the scintillator  188  or the antiscatter grid  192 . 
     As discussed above, the x-ray mask  190  should be made of high atomic number materials  191  on x-ray transparent substrate  192 , so that the x-rays can be substantially completely blocked with even a thin mask. The desirable thickness will dependent on the allowable transmitted x-rays and the x-ray energy. Gold is most commonly used as x-ray lithography masks. The attenuation factor of gold over the density, μ/ρ, varies with x-ray energy. For example, at x-ray energy of 22.16 keV, μ/ρ=59.7 cm 2 /g and at x-ray energy of 30 keV, μ/ρ=25.55 cm 2 /g, where ρ=19.3 g/cm3 is the density of gold. The amount of x-ray that penetrates the mask is equal to exp(−μL), where L is the thickness of the mask. Typically gold masks of can produce apertures with dimensions of 75 μm to 100 μm and vertical walls are routinely used to block x-rays in the 5-20 keV range. The mask needs to be thicker as the x-ray energy increases. The aperture walls of the mask should ideally be slanted along the direction at which the x-rays are received. If the x-ray source is from a point, then the mask should have the configuration shown schematically in FIGS. 25 a,    25   b  or  25   c,  in which the slant angles increase with distance from the center of the mask. The top layer of the mask in FIG. 25 c  does not have to have the same thickness as the bottom layer. 
     On the other hand, if the x-ray source is a parallel beam, the mask should have a configuration like that shown schematically in FIG. 25 d,  in which the aperture walls are all substantially vertical. The photoresist used in making the x-ray mask  193  does not have to be removed if it is x-ray transparent material, as shown in FIG. 25 e.  This is also true for a mask focused to a point x-ray source. 
     In an imaging apparatus  100  as shown in FIGS. 1 and 2, x-ray scatter can be reduced if the mask is thick and configured as an antiscatter grid. However, in the imaging apparatus  180  as shown in FIG. 22, x-ray scatter can be reduced even without the use of an antiscatter grid. 
     That is, when the x-ray sensitive area ε of the detector pixels is small compared to the area associated with the detector pitch E, the scatter is reduced by approximately the ratio ε/E. Alternatively, a thin mask  200  with aperture d 9 ×d 9  can be used in the imaging apparatus  180  in place of the antiscatter grid  192 , as shown schematically in FIG. 26, to reduce x-ray scatter by the ratio of (d 9 /D 1 ) 2 . 
     FIG. 27 is a schematic illustration of another embodiment of an imaging apparatus according to the present invention. Imaging apparatus  202  includes a substrate  204 , a digital detector pixel array  206  comprising detector pixels  208 , a scintillator  210 , and an x-ray mask  212  having apertures d 10 ×d 10 . However, in this embodiment, the mask is placed a distance λ 1  above the scintillator, and the object (not shown) to be imaged is placed above the x-ray mask  212 . The mask wall thickness and the distance x can act as an antiscatter grid. Alternatively, a properly aligned double mask  214 , having apertures d 11 ×d 11  and individual mask portions separated by an appropriate distance λ 2 , can be used to reduce scatter as shown schematically in FIG.  28 . 
     The invention as described with regard to FIGS. 1-28 employs a detector having a detector pixel pitch that is larger than the x-ray mask opening. The following embodiment of the invention employs detectors that have small pixels to obtain high-resolution images. A schematic of a CCD is shown in FIG.  29 . The pixel sizes of the CCD can have dimensions d 12 ×d 12 , with d 12  being less than 10 μm. However the resolution of the conventional x-ray image is degraded by the phosphor so that the small pixels of the CCD still cannot produce high-resolution images. 
     The concept described above is also applicable to the CCD detector. A group of the CCD can be configured together to collect data for one x-ray image pixel, where d 12  is the pixel pitch of the CCD. The CCD arrays can be used in configurations shown in FIGS. 1,  2 ,  22 ,  24 ,  26 ,  27  and  28 . 
     FIG. 30 is a schematic illustration showing the pattern of x-rays which passes through the mask overlaying the active area of the detector pixels of the detector pixel array shown in FIG.  29 . The example shown in FIG. 30 utilizes 3×3 CCD pixels to collect the information relating to x-ray intensity for one x-ray image pixel, i.e., 3(d 12 )=D 2 , and d 13  is the x-ray spot size overlapping the CCD. 
     The signal collected by each group of CCD pixels with dimension D 2 ×D 2  under an x-ray beamlet will be grouped together to form the signal for the x-ray beamlet. Each D 2 ×D 2  group of pixels is effectively a macro pixel analogous to a single pixel of D 1 ×D 1  as shown, for example, in FIG.  6 . For illustration purposes, nine CCD pixels form a macro pixel in FIG.  30 . 
     If the CCD pixels are much smaller than D 2 , then slight misalignment of the CCD array with respect to the mask can be tolerated by redistributing the signal of the CCD pixel to different macro pixels using software algorithms. The amount of misalignment may be on the order d 11  over a distance of tens of D 2 . 
     When CCD detectors are used and d 13 /d 12  is greater than or equal to one, only the mask, and not the detector, needs to move for configurations shown in FIGS. 1,  2 ,  22 ,  27  and  28 . Neither the mask nor the detector are required to move for the configuration shown in FIG. 24 a.    
     The high-resolution x-ray imaging apparatus discussed above according to the present invention has many applications. In addition to medical applications (e.g., mammography), such imaging apparatus can be used in scientific research, defense and security environments, biotechnology, x-ray microscopy, x-ray astronomy, three-dimensional x-ray tomography and various industrial applications such as those in which non-destructive testing is required. 
     For example, radiographic testing is used in industry in process control to detect manufacturing flaws and is increasingly integrated as a crucial component on the manufacturing floor. The trend of non-destructive testing is moving toward the use of real-time, non-film radioscopic systems over traditional film-based systems. Digital non-destructive evaluation offers all the traditional benefits of detecting microscopic flaws and providing permanent inspection records. It enables new capabilities such as computer-based inspection methods and cost reduction. The electronics and automotive industries have moved fastest to adopt radioscopy; many other industries are following this trend. 
     The spatial filtering which is performed by the present invention to obtain high-resolution digital x-ray or gamma ray images provides several advantages. The imaging apparatus can use either direct or indirect x-ray or gamma ray conversion to generate signals representative of the image. The invention provides an improvement of the MTF beyond the limitation of the pixel pitch of the detector pixel array. Image degradation by conversion blurring caused by phosphor screens can be minimized, and image degradation by oblique x-ray incidence can be minimized, thus providing improved image resolution as well as more spatially uniform image resolution. In medical applications, the method and apparatus of the present invention also allow for x-ray detection efficiency beyond the limitation of the fill factor of the imager, without the need for increasing the x-ray or gamma ray dosage to a patient. 
     In addition, a wide range of image resolutions can be achieved using the present invention, with digital x-ray or gamma ray images having a resolution as small as 1 μm. This concept of using mask to select the resolution is independent of the dimensions. Typically, the pixel size of gamma cameras are large while the pixel size of the CCDs are typically small. The pixel size depends on the energy of the radiation to be detected, the application and availability of detectors. Similarly, the mask thickness and the aperature size depends on the application&#39;s needs, the x-ray energy and the ability to fabricate the aperture size with the appropriate mask thickness. 
     Although only a limited number of exemplary embodiments of the invention have been described in detail above, those skilled in the art will readily appreciate that many modifications are possible in the exemplary embodiments without materially departing from the novel teachings and advantages of this invention. Accordingly, all such modifications are intended to be included within the scope of the invention as defined in the following claims.

Technology Classification (CPC): 6