US 20020196899 A1
A system for spectroscopic imaging of bodily tissue in which a scintillation screen and a charged coupled device (CCD) are used to accurately image selected tissue. An x-ray source generates x-rays which pass through a region of a subject's body, forming an x-ray image which reaches the scintillation screen. The scintillation screen reradiates a apatial intensity pattern corresponding to the image, the pattern being detected by a CCD sensor. The image is digitized by the sensor and processed by a controller before being stored as an electronic image. Each image is directed onto an associated respective CCD or amorphous silicon detector to generate individual electronic representations of the separate images.
1. An x-ray imaging device comprising:
a support surface that supports a region of interest of a patient to be imaged between an x-ray source and an x-ray conversion device, the conversion device having a first two-dimensional area and emitting a signal correlated with the received radiation, said signal extending across the first two-dimensional area; and
a solid state device having a two dimensional array of pixels and generating an electronic representation of the region of interest, the signal received at the array of pixels extending across a second two-dimensional area.
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9. An x-ray imaging device comprising:
a support surface that supports a region of interest of a patient to be imaged between an x-ray source and an x-ray detector, the detector including a solid state device having a two dimensional array of pixels and generating an electronic representation of the region of interest, the signal received at the array of pixels; and
a controller connected to the detector to control readout of image data to an electronic memory.
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 This application is a continuation of U.S. patent application Ser. No. 09/389,760 filed Sep. 2, 1999 which is a continuation of U.S. patent application Ser. No. 08/469,895 filed Jun. 6, 1995, now U.S. Pat. No. 6,031,892 which is a continuation-in-part of U.S. patent application Ser. No. 08/438,800 filed May 11, 1995 which is a continuation-in-part of U.S. patent application Ser. No. 07/853,775 filed Jun. 2, 1992, now U.S. Pat. No. 5,465,284, which is the U.S. National Phase of International Application No. PCT/US90/07178, filed Dec. 5, 1990 and which is a continuation-in-part of U.S. patent application Ser. No. 07/446,472, filed Dec. 5, 1989, now U.S. Pat. No. 5,150,394. The entire contents of the above applications are incorporated herein by reference in entirety.
 In recent years the use of radiological examining equipment to make measurements of bone density in patients has continually increased. In particular, the use of such equipment in diagnosing and analyzing osteoporosis has become prevalent in the medical community. Osteoporosis is characterized by the gradual loss of bone mineral content or atrophy of skeletal tissue, resulting in a corresponding overall decrease in average bone density. Such a condition is common in elderly women and greatly increases the risk of fracture or similar bone related injury.
 The presently available techniques for the radiological measurement of bone density utilize a rectilinear scanning approach. In such an approach, a radiation source, such as a radionuclide source or an x-ray tube, and a point detector are scanned over a patient in a raster fashion. This scan results in an image which has been derived from the point-by-point transmission of the radiation beam through the bone and soft tissue of a patient. The calculation of the bone-mineral concentration (the “bone density”) is usually performed by a dual energy approach.
 The current rectilinear scanning approach is generally limited by its long scanning time and its lack of good spatial resolution. The poor spatial resolution results in an inability to provide an image displaying high anatomical detail and which will permit accurate determination of the area in the scan occupied by bone. Moreover, the output of the x-ray source and the response of the detector must be closely monitored in order to assure high accuracy and precision.
 In accordance with the present invention, a stationary bone densitometry apparatus is provided for examining a subject's body. A dual energy x-ray source directs a beam of x-ray radiation toward the subject's body. The radiation is applied to the entire region of the body being examined. A scintillation screen receives the x-ray radiation passing through the body of the subject, and emits radiation in the visible spectrum with a spatial intensity pattern proportional to the spatial intensity pattern of the received x-ray radiation.
 A charge coupled device (CCD) then receives radiation from the scintillation screen. This CCD sensor generates a discrete electronic representation of the spatial intensity pattern of the radiation emitted from the scintillation screen. A focusing element between the screen and the CCD sensor focuses the scintillation screen radiation onto the CCD sensor. To prevent ambient radiation from reaching the CCD sensor, the present embodiment employs a shade or hood surrounding a region between the scintillation screen and the CCD sensor. A CCD controller then processes the electronic representation generated by the CCD sensor, and outputs corresponding image data.
 A dual photon x-ray source is used to allow the examination to be performed with x-rays at two different energy levels. This source can be an x-ray tube, or a radionuclide source with a filter element to remove one of the energy levels when desired. Correlation of the image data retrieved using each of the two x-ray energy levels provides quantitative bone density information.
 A focusing element between the scintillation screen and the CCD sensor can take the form of a lens or a fiber optic reducer. An image intensifier can be used in conjunction with the CCD sensor. The image intensifier can be a “proximity type” image diode or a microchannel based device. It can also be directly attached to the CCD. An image store used with the CCD controller allows manipulation of the CCD sensor output signals by a data processor. This includes the correlation of measurements utilizing x-ray beams of two different energy levels. The system can also be adapted to operate at higher shutter speeds enabling the counting of x-ray transmissions. This provides energy measurements of x-ray transmissions that are useful in certain applications.
 In an alternative embodiment, a detector made of amorphous silicon is used to receive and detect the radiation from the scintillation screen to generate the electronic representation of the spatial intensity pattern of the x-ray pattern. The amorphous silicon detector can replace the CCD detector or it can be used to receive the x-rays directly.
 In another preferred embodiment, the apparatus of the invention includes two scintillation screens, each of which is associated with its own respective CCD detector or amorphous silicon detector. One of the scintillators is reactive to high-energy x-rays and generates an optical image of the spatial intensity pattern of the high-energy x-ray pattern. Its associated detector detects the image and generates an electronic representation of the high-energy x-ray pattern. The other scintillator is reactive to low-energy x-rays simultaneously generate an optical image of the low-energy pattern. Its associated detector generates an electronic representation of the low-energy x-ray pattern. The data processor performs the correlations of the measurements for the x-rays at two different energy levels.
 An additional preferred embodiment is directed to systems and methods of imaging spectroscopy where charge coupled device (CCD) is optically coupled to a scintillator and measures or counts the spatial intensity distribution of a radionuclide that has been introduced into bodily tissue, either in vivo or in vitro. CCD's of sufficient thickness can be used to measure gamma ray events without the use of a scintillator in certain applications. The CCD has sufficient resolution and sensitivity to measure such distributions accurately, usually in less than two minutes. Radiation sources that emit radiation having an energy in a range between 10 and 2,000 keV, and preferably in the range between 20 and 600 keV, are delivered to the cancerous tissue or any other suitable pathologic abnormality.
 The CCD acquires “frames” of information by counting the number of gamma-ray events over a selected period of time. Each frame, or a sequence of frames that have been added or summed to provide an image, can be filtered using pulse height analysis techniques to substantially reduce or eliminate scattered radiation. Pulse height analysis can also be utilized to discriminate between signals having different energy levels that contain diagnostically significant information. The system's discrimination and energy measuring capabilities render it suitable for diverse applications.
FIG. 1 is a perspective view of the imaging system of the present invention.
FIG. 2 illustrates in schematic view a bone density measuring apparatus using a lens to focus image data from a scintillation onto a CCD sensor.
FIG. 3 illustrates in schematic view a bone density measuring apparatus using a fiber optic reducer to deliver an image from a scintillation screen to a CCD sensor.
FIG. 4 illustrates another preferred embodiment for the scintillation screen employing a fiber optic plate.
FIG. 5 is an illustration of the pixel array of a binnable CCD sensor.
FIG. 6 is an alternative preferred embodiment to the bone density measuring apparatus of FIG. 2.
FIG. 7 is another alternative preferred embodiment to the bone density measuring apparatus of FIG. 2.
FIG. 8 is a perspective view of a scanning system of the present invention.
FIG. 9 is a schematic sectional view illustrating the sensor control system.
FIG. 10 is a schematic sectional view illustrating a frame transfer CCD used for both emission and transmission studies.
FIG. 11 is a schematic sectional view of a CCD imaging system for both emission and transmission studies.
FIG. 12 illustrating a process flow sequence that is used in performing the imaging methods of the present invention.
FIG. 13 is an alternate embodiment of a CCD imaging system that can be employed for both emission and transmission studies.
FIGS. 14A and 14B illustrate a process flow sequence for conducting emission and transmission studies of tissue
FIG. 15 is a schematic diagram of an alternative preferred embodiment to the bone densitometry measuring apparatus of FIG. 2 using dual scintillation screens and dual detectors.
FIG. 16 is a schematic diagram of an alternative preferred embodiment to the bone densitometry measuring apparatus of FIG. 15.
FIG. 17 is a schematic diagram of a variation of the bone densitometry measuring apparatus of FIG. 16.
FIG. 18 is a schematic diagram of another alternative preferred embodiment to the bone densitometry measuring apparatus of FIG. 15 having dual amorphous silicon image sensors.
FIG. 19 is a schematic diagram of an alternative detection structure including dual amorphous silicon image sensors which can be used with the various embodiments of the bone densitometry measuring apparatus of the invention.
FIG. 20 is another preferred embodiment of a bone densitometer for static, scanning or stepped imaging procedures.
 FIGS. 21A-21C illustrate alternate embodiments for the detector assembly of FIG. 20.
 FIGS. 22A-22B illustrate scanning or stepped imaging procedures.
FIG. 23 is another preferred embodiment of an imaging system in accordance with the invention.
FIG. 24 is another preferred embodiment of an x-ray imaging system in accordance with the invention.
FIG. 25 is a schematic diagram of a detection structure used for bone densitometry measurements and tissue lesion imaging in accordance with the present invention.
FIG. 26 is a preferred embodiment of the invention in which a dual spaced array is used for digital mammographic imaging and quantitative analysis.
FIG. 27 illustrates another preferred embodiment in which the imaging elements in each linear array positioned at a different angle relative to the patient and the x-ray source.
FIG. 28 illustrates a preferred embodiment in which a large number of imaging elements are arranged in a large array conforming in size to a standard x-ray film cassette which can have three or more spaced linear arrays.
FIG. 29 illustrates a cross-sectional view of a linear array using a common fiber-optic plate and scintillator.
FIG. 30 illustrates a system for translating the array relative to a radiation source.
FIGS. 31A and 31B illustrate the process of a two step imaging sequence.
FIG. 32 illustrates a system for sequential imaging or scanning of tissue in which the array is moved relative to a source.
 In FIG. 1 a preferred embodiment of the invention for performing bone densitometry studies uses a detector 10 and either an x-ray tube 12 or a radionuclide radiation source such as Gadolinium-153. The detector 10 comprises a scintillating plate 20 which is optically coupled to a two-dimensional charge-coupled device 24 (CCD). The CCD is a two dimensional array of detectors integrated into a single compact electronic chip. The optical coupling between the scintillating plate 20 and the CCD 24 is accomplished by an optical grade lens 25. Such a lens should have a low f-number (0.6-1.8) for adequate light collection from the screen. The collection efficiency (E) of light from the scintillating plate emitted in the direction of the CCD can be computed by the equation:
 t: Transmission factor of light through the lens
 m: magnification from the Scintillating plate to the CCD
 f: f-number of the lens
 In an alternate approach, the optical coupling between the scintillating plate and the CCD can be performed with a fiber optic reducer.
 Referring to FIG. 2, a bone densitometry apparatus 10 has an x-ray tube 12 which delivers a beam of x-rays 14 towards the body of a subject 16 being examined. The x-ray tube is capable of emitting x-ray radiation at each of two distinct energy levels. The two energy levels are used to obtain two distinct x-ray images of the patient, as is discussed later. Note in comparison to FIG. 1, the source can be positioned above the patient and the detector below the table.
 When the subject 16 is irradiated with the x-ray energy, a percentage of the x-rays reaching the subject 16 is absorbed by the subject's body, the amount of absorption depending on the density of bone or tissue upon which the x-rays are incident. Since x-rays generally travel in a straight line, the x-ray energy exiting the subject's body on the side of the body away from the source 12 is a spatial representation of absorption in the subject's body, and therefore of relative tissue and skeletal densities.
 To receive the x-rays passing through the subject's body, a scintillation screen 20 is provided on the side of the patient away from the x-ray source 12. The scintillation screen 20 is a fluorescent material sensitive to x-rays, and when it receives x-ray energy it reradiates visible light. The spatial intensity patterns of the radiation emitted from the scintillation screen is proportional to the spatial intensity pattern of the x-ray radiation received by the screen 20. Thus the scintillation screen 20 provides an image in the visible spectrum, or alternating in the ultraviolet or near infrared, which is regionally proportional to the x-ray image reaching the scintillation screen 20.
 A lens 22 is positioned between the scintillation screen 20 and a CCD sensor 24. The CCD sensor 24 is an array of photosensitive pixels using closely spaced MOS diodes which convert photons to electrons and thereby generate a discrete electronic representation of a received optical image. The lens 22 faces the scintillation screen and focuses the visible light emitted from the scintillation screen 20 through the lens 22 and onto the surface of the CCD sensor 24. In order to prevent ambient light from reaching the CCD sensor, a shade surrounding the region between the scintillation screen 20 and the lens 22 is provided in the form of a photographic bellows 26. The shading of bellows 26 serves to reduce the optical noise level of the image signal reaching the CCD sensor 24.
 Although the scintillation screen 20 absorbs most of the x-rays incident upon it, some may still be transmitted through the screen 20 and interfere with the optical image signal of the scintillation screen 20. The direct interaction of x-rays with a CCD sensor produces very bright pixels resulting in a “snow” effect in an optical image detected by the sensor. In addition, prolonged direct x-ray irradiation of a CCD sensor can increase its dark current. For these reasons, an optical grade lead-glass or lead acrylic filter 28 is positioned between the scintillation screen 20 and the lens 22 or alternatively, between the lens and the CCD. The lead-glass filter 20 absorbs most of the stray x-rays and prevents them from reaching the CCD sensor 24. An anti-scatter grid 29 is used between the patient and scintillation screen for preventing scattered x-rays from reaching the screen.
 During a typical examination, the subject 16 is placed between the x-ray source 12 and the scintillation screen 20. The x-ray source is then activated for a short time interval, typically one to five seconds. As x-rays are differentially transmitted and absorbed through the body of the subject 16, they interact with the scintillation screen 20. Upon interaction, the screen 20 emits light in the visible part of the electromagnetic spectrum. In the present embodiment, the scintillation screen is a terbium-activated material and emits light in the region of 540 nm.
 The light emitted from the scintillator is transported to the CCD sensor via the lens 22. Upon interaction with the CCD sensor 24, light energy is converted into electrons which are stored in each pixel of the CCD sensor 24. The CCD sensor 24 of the present embodiment consists of 512×512 pixels, but such sensors come in a number of different sizes. The CCD sensor “integrates” the image signal from the scintillation screen in that it senses the optical image and stores charge during the entire x-ray exposure interval. After termination of the x-ray exposure, the discrete representation in the CCD 24 is read out by CCD controller 30. The CCD controller 30 reads the image representation from the CCD sensor 24 pixel and organizes it into a digital array. The digital array, representing spatial position and x-ray intensity, is then output to a memory or image store 32. From the image store 32, the image can be accessed by a data processor 34 for performing image processing techniques. A cathode ray tube (CRT) 36 is also provided to allow the image to be displayed before or after processing by data processor 34.
 Unlike other conventional detection schemes, such as film screen radiography, CCD-based imaging provides a linear quantitative relationship between the transmitted x-ray intensity and the charge generated in each pixel of the CCD. After the first high energy x-ray exposure is acquired, the resulting image is stored in image store 32 and a second exposure with a low energy x-ray beam is acquired with the subject 16 in the same position. During this exposure, a low energy x-ray beam is used which is typically at about 70 kVp with a tube current at about 1 mA. The tube is capable of accelerating electrons at 40 kVp and up to approximately 140 kVp. Note that the tube potential and current are controlled by the computer menu. The low energy x-ray image is then stored in image store 32 with the high energy exposure. Each image provides quantitative information about the relative transmission of x-rays through soft tissue and bone.
 Once both images are obtained, comparative processing techniques of dual photon absorptiometry are applied to determine quantitative density measurements of those body regions scanned by the x-rays. The correlation of two images generated by x-rays of two different energy levels over a short time interval results in the substantial reduction in the likelihood of systematic pixel-by-pixel errors caused by instability of the x-ray tube output.
 Because the present embodiment of the invention is concerned with an area detector as opposed to a scanning detector, the measurement time necessary for a densitometry examination is greatly reduced. Rather than scanning across the region to be examined in a rectilinear fashion, the entire region is irradiated simultaneously and the resulting image processed simultaneously. Typically, the entire procedure using the present dual photon technique lasts 30 to 60 seconds, depending on the power of the x-ray tube and processing speed of the supporting electronics.
FIG. 3 shows an alternative embodiment to that of FIG. 2. In this embodiment, the x-ray tube source 12 of FIG. 2 is replaced with radionuclide source 40. The radionuclide source is gadolinium-153. Gadolinium-153 emits photons simultaneously in two energy bands, a lower energy band of 44 keV and an upper energy band of 100 keV. Thus, the gadolinium source is a dual photon radiation source. In order to allow the images from the two different energy levels to be obtained separately, an x-ray filter 42 is placed between the source 40 and the subject 16. In the present embodiment, the filter 42 is copper or a K-edge filter, and eliminates nearly all of the low energy (44 keV) emission from the beam. Removal of the filter restores the beam to its dual energy nature. The filter 42 is implemented as an electromagnetic shutter which may be opened and closed in the line of the x-ray beam. A high energy image is acquired first with the filter shutter closed, after which an image is obtained using the dual energy beam with the shutter open.
 Both electronic images are stored, and an image representative of the transmission of only the low energy photons is obtained by electronically subtracting the high energy image from the dual energy image with the data processor 34. Once both images are obtained, comparative dual photon processing techniques are used to make quantitative density calculations.
 An additional feature of the embodiment of FIG. 3 is the replacement of the lens 22 of the FIG. 2 embodiment with a fiber optic reducer 44. The fiber optic reducer 44 is a focusing device consisting of a large array of optical fibers packed tightly together, and leading from the scintillating screen 20 to the CCD sensor 24. Near the CCD sensor 24, many of the fibers can be fused together, thus combining the signals present on individual fibers. The effect is a compression of the image from the input of the reducer 44 at the scintillation screen 20 to the reducer output at the CCD sensor 24. In this manner, the reducer 44 effectively focuses light from the scintillating screen 20 onto the CCD sensor 24 without the necessity of a lens for the focusing region.
 Although they are shown together in FIG. 3, it is not necessary to use the fiber optic reducer 44 with the radionuclide source 40. Either element can be substituted into the configuration of FIG. 2 individually. The x-ray filter 42, however, should be used with the radionuclide source 40 to provide a dual photon discrimination capability. Note, however, that pulse height analysis can be performed in conjunction with the embodiment of FIGS. 10 & 11.
FIG. 4 shows an alternative to the scintillation screen 20 of FIGS. 2 and 3. The screen 48 depicted by FIG. 3 is a scintillating fiber optic plate. The plate 48 is a fiber optic faceplate consisting of scintillating fibers 50 running through the plate. The fiber optic plate is optically interfaced to the CCD in essentially the same way as the scintillation screen 20 of FIG. 2, but the fiber optic plate 48 allows for greater quantum efficiency due to increased x-ray stopping capability.
 Shown in FIG. 5 is a representation of the pixel array of the CCD sensor 24. The array shown in FIG. 5 is only 10×10 for illustrative purposes, and the actual array can be of different dimension. Each pixel in the array is an individual photosensitive element which contributes to the overall image detected by the array. A feature of the CCD sensor of the present embodiment is the capability of the pixels of the sensor 24 to be “binned” together. The binning of the pixel array refers to the ability of the sensor electronics to combine groups of pixels together to form “super pixels” which are then identified as single picture elements.
 Charge is binned by combining charge packets contained in two or more adjacent potential wells into a single potential well during charge readout. Serial and parallel binning can be combined to perform two dimensional binning from any rectangular group of wells or detector elements.
 The dark lines in the binnable array of FIG. 5 illustrate where individual pixels might be grouped together. For example, the four upper left hand corner pixels 50 can be binned together through control of the CCD sensor 24 to form a super pixel. The super pixel is then identified by the CCD electronics as a single pixel, the light intensity reaching each pixel 50 being averaged across the surface of the entire super pixel. In this manner, the dimension of the array can be electronically controlled. As can be seen in FIG. 5, if groups of four pixels are binned together across the 10×10 array, the overall array dimension becomes 5×5. Although the binning of the CCD sensor 24 reduces the resolution of the pixel array, the relative percentage of noise is also reduced, thus providing an improved signal to noise ratio.
 The following x-ray data acquisition approach is an alternative to the one described above. In this approach, an image is acquired at high energy and the CCD is read in the normal non-binned mode. Due to the high penetration of the high energy beam through the body, the x-ray fluence exiting the body is high as compared to that of the low energy beam. Therefore, the resulting charge signal per CCD pixel is relatively strong. This image is stored as the high energy image. Also, this image is used in order to compute the area of the bone to be measured by manual selection of the region of interest or by automatic edge detection. Therefore, we take advantage of the high resolution image for greater accuracy in the measured bone area. Previously, the accuracy and precision of bone density measurements are limited to a great extent by suboptimal spatial resolution. The next image which is acquired with low energy is read out by the pixel binning approach, e.g., using a 2×2 pixel binning. The transmission of the low energy beam through the body is low as compared to the high energy beam. Therefore, in order to record a strong signal in each CCD pixel we must increase the radiation dose.
 Alternatively, the binning technique can be used for the low energy in order to increase the signal to noise ratio and to decrease the radiation dose. This dual mode acquisition procedure is a very powerful tool for improving the signal to noise ratio and lowering the radiation dose to the patient.
 Although the arrangement of optical elements as shown in FIGS. 2 and 3 represent preferred embodiments, the functionality of the system is not dependent upon such an in-line type of optical transmission. FIG. 6 shows an alternative arrangement of optical elements where the CCD sensor 24 is set at an angle relative to scintillation screen 20, and mirror 52 is used to reflect the radiation given off by the scintillation screen toward the CCD sensor 24. Lens 22 is shown between CCD sensor 24 and mirror 52 and focuses the image onto the CCD sensor. However, the focusing of the scintillation screen image can take place before or after the image reaches mirror 52. In fact, the mirror itself may be shaped to provide focusing of the image from the scintillation screen 20.
FIG. 7 shows another alternative arrangement of optical components. In FIG. 7 the subject 16 is suspended by a support 54 which is transparent to x-rays. The support 54 keeps the subject 16 elevated a distance above scintillation screen 20. As the x-rays reach scintillation screen 20, the screen 20 reradiates image data from the same surface upon which the x-ray radiation is incident. Mirror 52 is now aligned to reflect this image towards CCD sensor 24 which collects the image as focused through lens 22 to be processed by the CCD controller 30.
 As with the arrangement of FIG. 6, the focusing of the image from the scintillation screen 20 may take place before or after it is reflected by the mirror 52, or may be focused by the mirror 52 itself. In addition, any of the optional elements previously discussed may be substituted into the arrangement of FIG. 5 or FIG. 7. This includes the x-ray absorbing screen 28, the anti-scatter grid, the fiber optic reducer 44, and the fiber optic faceplate 48.
 A very effective, radiation dose-efficient approach for reducing x-ray scatter and increasing the dynamic range of electronically acquired x-ray images is the use of a slit-scan method. In this approach, a fan beam of x-rays is scanned over the patient and a linear array of detectors is used to detect the transmitted radiation. In typical applications the length of the detector restricts the width of the area that can be covered with one pass. Also, many small linear CCD or photodiode arrays are used to form a line of detection. This results in a rather complex detector assembly. If cooling of the detector assembly is required, it is difficult to accomplish for such an extended detector. Also, image intensification by using an electronic intensifier becomes difficult and very costly.
 An alternative embodiment for dual energy bone densitometry takes advantage of the merits of slit-scan geometry without using a linear CCD or photodiode array. This approach is illustrated schematically in FIG. 8. An area CCD sensor 64 is used in conjunction with a line-to-area fiber optic converter 62. This converter can be made of flexible or rigid optical fibers with cladding of lower index of refraction than the core material. As shown in FIG. 8, the CCD 64 is divided into a number of rows and a fiber optic ribbon is optically coupled or bonded to each row. The coupling of the CCD 64 to the converter 62 can be accomplished using the various systems described in connection with other embodiments. An extramural absorber can be used to prevent light crossing from one fiber to another. The other ends of each ribbon are arranged in tandem to form a linear sensor. In front of the linear sensor (input end), an x-ray converting scintillator 60 is used such as gadolinium oxysulfide activated with terbium (GOS:Tb). Alternatively, a scintillating fiber optic plate can be used for improved quantum efficiency at higher energies. A linear x-ray sensor with a very compact area detector is employed with the slit-scan embodiment.
 A typical linear detector of this type comprises a few ribbons in tandem along the length of the slit, and from one to a multitude of ribbons across the width of the detector slit.
 In a typical example, consider a 512×512 pixel CCD where each pixel has an area of 20×20 microns. A fiber optic bundle with individual fibers of 60 microns in diameter is used for the embodiment. On the CCD each fiber will cover an area of approximately 3×3 pixels. Perfect alignment between each pixel and fiber is desirable but it is not essential for this application. Close packing of the fibers will result in an array of 170×170 or a total of 29,127 fibers covering the entire area of the CCD. Each ribbon of fibers corresponds to one row consisting of 170 fibers and covering approximately 512×3 pixels on the CCD. If all ribbons emerging from the CCD were arranged in tandem, the linear sensor would be approximately 175 cm in length. Alternatively, the ribbons can be arranged with a small number in tandem and a small number across the width of the slit. Using the above CCD, a 15.3 cm linear detector can be made with approximately 15 ribbons in tandem thus using only a small fraction of the CCD area.
 Full use of the CCD area can be made by stacking the ribbons in groups of 15, (one ribbon per CCD row), thus creating a quasilinear detector consisting of an array of 2,550×11 fibers optically coupled to an x-ray scintillator. The dimensions of this slit detector will be 153×0.66 mm with a total sensing area of 1.0 cm2. It is important to note that the total sensing area of the slit must be approximately equal to the total area of the CCD and the linear dimensions of the CCD. A wider or longer slit will result in a larger area at the output end. In this case, a larger CCD can be used for a fiber optic reducer optically bonded between the fiber optic converter and the CCD. Alternatively, the converter itself can be tapered to match the size of the CCD. For higher spatial resolution the fiber optic converter is made with optical fibers of smaller diameter (5-6 microns).
 If higher signal amplification is required for some high detail low dose applications, a proximity focused image intensifier can be optically bonded between the fiber optic taper and CCD or between the fiber optic converter and fiber optic taper. The image intensifier can be a proximity diode type or a microchannel plate device, both commercially available. Alternatively, an integral assembly of CCD and intensifier can be used commonly called an “intensified CCD”. Another approach is to use a lens coupling between the output surface of the fiber optic converter and the intensified or non-intensified CCD.
 Cooling of the CCD can be accomplished easily by a thermoelectric cooler. Cooling is required only when very high contrast resolution is required and the image acquisition time is relatively long. If the CCD is read out at 500 kHz (5×105 pixels/sec), an area of 150 mm×150 mm of the subject can be scanned in approximately 114 seconds (approximately 2 minutes). Faster scanning is attainable by increasing the readout rate of the CCD.
 Alternatively, a frame transfer CCD such as the one illustrated in FIG. 10 can be used for faster scanning. This device uses one half of its sensing area for storage and not for sensing. In this way the transfer of the image from the sensing area 91 to the storage area 93 is accomplished in a few milliseconds. A smaller CCD such as a 128×128 or a 64×64 element could be used for this purpose in a similar arrangement as with the 512×512 CCD. Also, larger area CCDs can be used for this purpose. Pixel binning as described previously can be applied in this detection approach. A Gadolinium-153 (Gd-153) radiation source can be used as described in previous sections in place of an x-ray tube. The Gd-153 source is a small pellet or a collimated line source parallel with the long dimension of the detector.
 The line to area conversion design enables us to remove the CCD from the direct path of the x-ray beam, thus it allows for easy shielding of the CCD from direct x-ray interactions. This prolongs the useful life of the CCD and it alleviates the “snow” effect which results from direct interactions of x-rays with the sensor. Moreover, this approach allows for greater light transport efficiency between the scintillator and CCD than lenses or fiber optic tapers. Note that the pixel binning approach enables the operator to select the desired spatial resolution and contrast without any mechanical modifications on either the x-ray beam or the detector collimator. The pixel size of the detector which determines resolution and contrast can be controlled by a command from the computer. This x-ray imaging modality can be used very effectively to optimize the scan depending on the patient size, and medical history.
 An alternate approach provides an improved rectilinear scanning method for quantitative x-ray radiography. In this embodiment, a two dimensional CCD optically coupled to a scintillator is used as the detector of x-rays in a rectilinear scanning mode. The CCD may be a full frame or a frame transfer device. The frame transfer CCD will enable faster data scanning and acquisition.
 The CCD scintillator assembly is extremely critical to the performance of the system. Direct optical bonding of a polycrystalline scintillator such as gadolinium oxysulfide with the CCD is possible but this approach is not efficient in shielding the CCD from direct x-ray interactions. If the thickness of the layer is increased the spatial resolution of the x-ray images degrades due to light diffusion. The use of a scintillating fiber optic plate between the polycrystalline scintillator and the CCD provides a solution to this problem.
 A scintillating fiber optic plate is a fiber optic faceplate designed to convert x-rays or U.V. light into green light with peak emission at about 550 nm. This faceplate is manufactured with the extra mural absorber to prevent light diffusion between individual fibers. The area of the scintillating fiber optic plate must cover the CCD completely. The desirable thickness depends on the energy of the x-ray radiation. A thickness of 5 to 10 mm is preferable but a thinner or thicker plate can be used. The use of a very thick scintillating fiber optic plate such as 10 mm or 20 mm will eliminate virtually any undesirable direct x-ray interactions with the CCD. The scintillating fiber optic plate can also be used without the thin layer phosphor. However, the combination of the two will produce better image quality at a reduced radiation dose to the patient. Alternatively a conventional fiber optic plate can be used as a substrate to the scintillating fiber optic plate. The optical coupling of the polycrystalline phosphor on the fiber optic can be accomplished by direct deposition techniques or by using an optical adhesive.
 In an alternate approach, a bent fiber optic bundle can be used between the scintillator and the CCD. The geometry of the bent bundle allows for extremely effective shielding of the CCD from extraneous x-ray radiation. A lens coupling between the CCD and the fiber optic converter can also be used. For improved sensitivity, a proximity focused image intensifier, an image diode or microchannel plate can be used at the input end of the fiber optic or between the fiber optic bundle and the CCD. A preferred approach is to use the intensifier at the input end. A scintillator can be optically bonded to the input of the intensifier or an intensifier with a scintillating fiber optic input plate can be used.
 The x-ray tube is aligned in a C-arm configuration with the detector. The x-ray beam is approximately congruent with the area of the detector which is approximately 1×1 cm at the detector plane. As x-rays are transmitted through the patient, some (20%-60%) are absorbed by the primary polycrystalline scintillator producing visible light. This light is transmitted through the optically transparent fiber optic faceplate in the direction of the CCD. The x-rays not interacting with the primary scintillator will be absorbed by the fiber optic faceplate. If a scintillating fiber optic faceplate is used, these x-rays will be absorbed in the fibers thus producing additional scintillations. Therefore, the scintillating fiber optic plate acts as a light conduction device, x-ray shield, secondary x-ray detector and an x-ray signal amplifier.
 Upon interaction of the x-ray induced light with the photosensitive surface of the CCD an electron charge is generated which is proportional to the number of x-ray interactions in the scintillators. The cumulated charge on the CCD is then read out. However, in this rectilinear scanning mode, each CCD readout will correspond to a small segment of the total image, approximately one square centimeter. Therefore, the entire image is acquired by spatial additional of each image segment. For example, if a 15×15 cm field is covered and the sensor area is 1.0×1.0 cm, 152 (225) segments must be acquired and synthesized. A 512×512 pixel CCD operating at 500 kHz will read out each segment in 0.5 seconds and will require about 2 minutes for the entire scan at a scan speed of about 2 cm/sec. Faster scanning is attainable by increasing both the scanning speed and the readout rate of the CCD.
 A dual-energy scan will be acquired by first scanning the entire area at high tube potential, typically 130 kVp without binning and then repeating the scan at low tube potential, typically at about 70 kVp with binning. An automatic slide mechanism places high aluminum filtration for the high energy beam and less filtration for the low energy beam as described previously. The images of each energy level are stored in the computer for subsequent dual photon analysis. Pixel binned acquisition will be possible at both energies for improved precision. Where both high and low energy images are identically binned, this produces an exact correlation between the images produced. A third high energy-high resolution image can then be used to define the outline of the object being scanned. Note that a gadolinium isotope source with a shutter can be used.
 Alternatively, the energy level of the tube can be switched from low to high for each segment of the acquisition and each segment representing high and low energy is stored for subsequent analysis.
 An alternate approach employs light intensification from the screen to the CCD sensor. In this approach, an electrostatically focused image intensifier (In FIG. 2) is employed as the primary detector in place of the scintillating plate. This intensifier preferably employs Cesium iodide input phosphor with an approximate diameter of 15 cm and thickness of 0.3-0.5 min. The high voltage of the image intensifier tube can be reduced to approximately half the normal value. A reduction in the image intensifier accelerating potential will contribute to an improvement in the image contrast characteristics and dynamic range of the device. The CCD sensor is optically coupled to the output phosphor of the image intensifier by a fast lens with an f-number of about 1:1.0. Due to the high signal intensification, cooling of the CCD is not essential but it can be applied if very low thermal noise levels are desirable. The use of an intensifier allows for the use of a CCD with lower noise performance characteristics, thus lowering the cost and complexity of the instrument.
 Ideally, the detected signal is produced by x-rays that have been transmitted through the body without any scatter interaction. Detection of large amount of scatter events will result in non-linearities and in a reduction in the dynamic range. Effective suppression of scatter is accomplished by using a small field of view, typically 10 cm×10 cm and by using a air gap (approximately 20 cm) between the patient and the scintillating plate. Alternatively a small field of view can be used in conjunction with a linear or crossed antiscatter grid.
 An internal instrument stability control system has been incorporated to provide a means of automatic compensation for any instabilities in the x-ray tube potential and current. The stability control device is not essential for the operation of any of the described techniques but it provides better reliability and precision in the measurement of bone density. A schematic represen-tation of the proposed device is shown in FIG. 9. The output of the x-ray tube 12 is monitored by a pair of x-ray sensors 70 placed at a secondary x-ray beam port 78 adjacent to the main beam port 80 near the tube window. The sensors can be silicon diodes, cadmium telluride radiation sensors or any other solid state x-ray sensor. Alternatively, a pair of compact photomultiplier-scintillators or a photodiode scintillator assembly could be used. Both detectors operate in the charge integration mode and the detected signal is continuously monitored as a function of time during the entire scan for each energy. This time varying signal is digitized and stored in the computer memory. The change in the filtration of the secondary beam with energy is identical to that in the main beam because it is controlled by the same filter changing mechanism. As described further in connection with FIG. 12 the sensor system can be used to normalize the detected information or to control operation of the x-ray source to prevent or reduce unwanted variations in the source output.
 In front of one of the sensors 70 an amount of polymethyl methacrylate 86 is placed to simulate an average thickness of soft tissue. In front of other sensor 70 an amount of bone simulating material 84 is placed in an amount equivalent to that encountered in the spine or femur. Various hydroxyapatite-epoxy mixtures are commercially available for bone simulation in x-ray imaging. Therefore, a secondary detection system with a bone standard of known density and a soft tissue equivalent thickness is provided in this embodiment.
 The signals from each sensor 70 can be used to compute the density of the bone internal standard as a function of time during the scan. Any deviations from a constant density of this standard are due to changes in either the energy or intensity of the x-ray emission. Each value of bone density computed in the patient scan corresponds to a computed value of the bone standard. Therefore, each computation of bone density derived from a pair of high and low energy CCD frame acquisitions can be corrected or normalized by using the deviation from the density of the internal standard. For example, if the value of the bone standard during the rectilinear scan deviated by plus 3% in a given area of the image, the computed bone density of the patient scan must be corrected by that amount in this area. This internal reference approach can be used with all stationary and scanning embodiments described herein.
 In conjunction with the above calibration approach, a number of strips 72 (square rods) of bone simulating epoxy material, or aluminum of equivalent x-ray absorption are placed under the table 73 which run in the direction of the scan for the slit scan approach. Each linear strip has a different thickness or bone equivalent density. As the x-ray tube and detector assembly is scanned over the area to be tested, each set of rods are scanned and their density computed. The consistency of the measured densities of these rods is used to ensure proper operation of the system. This set of standards can be placed anywhere from the x-ray exit port 80 to the edges 74 of the detector 76.
 The imaging of radionuclide distributions in biological tissues or specimens is a routine task performed in virtually all biomedical research laboratories by the well established technique of autoradiography. In this procedure, a thin slice of a specimen is placed in contact with photographic film thus allowing the radiation from the specimen to expose the film. Subsequently, the film is processed by standard chemical development techniques, manually, or by using an automatic processor. Frequently, an intensifying screen is used in order to enhance the absorption efficiency of the image receptor and for a reduction in exposure time. Intensifying screens are especially useful when images of relatively high-energy gamma or x-ray emissions are recorded (20-200) keV. Also they can be useful for high energy electrons.
 Autoradiography produces images reflecting the biodistribution of a radionuclide and it has been established as a powerful tool in many biomedical disciplines. Its major shortcomings relate to problems with quantization of the relative or absolute concentration of radionuclide in an area of interest. This difficulty arises from the non-linearity of photographic film typically used and in reciprocity law failure when intensifying screens are used. Moreover, the development temperature, and in general, the condition of the processing chemicals have an influence on the film fog level and contrast. All these factors render quantization a very difficult and time consuming task which becomes vulnerable to many uncertainties in quantitative autoradiography. Despite this problem, several investigators have digitized film autoradiographs by using microdensitometers or video cameras for both quantization and image enhancement.
 In autoradiography, the image represents areas where the radiotracer has been extracted. The anatomical information on the original tissue slide is not transferred with great detail in the autoradiograph. For proper interpretation, it is necessary to observe the tissue slide and autoradiograph side by side in order to correlate radiotracer distribution with anatomy. Often it is necessary to superimpose the slide with the autoradiograph in order to identify the exact anatomic location of the radiotracer. In this process the accuracy in assigning an anatomic location to the tracer is severely compromised.
 One of the most important problems with autoradiography is the long period of time required in order to expose the film. In most applications this time ranges from a few hours to several days, even weeks in some cases. Therefore, the technician may have to wait for a few days in order to find out whether an exposure has to be repeated.
 Autoradiography does not relate to in vivo imaging of radionuclide distributions in humans or animals. Rather it relates to detecting radioactive distributions in excised samples. All available film-screen image receptors have extremely low quantum efficiencies for most gamma emitters commonly used for this purpose. Moreover, the presence of a large volume of tissue results in enormous amount of gamma ray scatter which will reach the image receptor and degrade the contrast and spatial resolution. The film-screen receptors do not have energy discrimination capabilities, therefore scattered events cannot be rejected. The use of a collimator to suppress scatter will result in a dramatic reduction in geometric efficiency.
 Thus the present invention, in its various embodiments, provides an effective means for performing autoradiography by providing a compact device that performs the data acquisition for autoradiography quickly and can superimpose both emission and transmission studies to correlate the emission image with the anatomical features of the object under examination. The embodiments described in connection with FIGS. 10 and 11 below can be used to perform autoradiography procedures.
 Radionuclide imaging of humans and animals is performed on a routine basis by using the Anger camera, most commonly referred to as a “Gamma Camera”. The gamma camera has a quantum efficiency in excess of 50% for the most commonly used radionuclides and it has the capability of discriminating scatter from primary photons by pulse-height analysis of each detected photon. The intrinsic spatial resolution of the gamma camera is approximately 3.5 mm. The total spatial resolution of the camera, including the degradation due to its collimator, can vary from 5 mm to 12 min. Modem gamma cameras can detect photons at the rate of 25,000 counts per second (cps) without significant dead time losses. At higher count rates, significant deviations are observed between true and detected events. This is due to limitations inherent in the design of both the detector assembly and processing electronics.
 The following presents a further embodiment relating to imaging of radionuclide distributions in tissue samples and in vivo quantitative imaging of humans and animals. This procedure employs a charge-coupled device to detect and process information to provide, in essence, a compact “gamma camera” using a highly sensitive stationary (or scanning) detector to conduct both emission and transmission studies at count rates up to 106 of the object being examined.
 Existing gamma cameras have limited spatial resolution, limited capability to perform in high count-rate conditions and it cannot be used to record x-ray transmission (radiographic) images with any degree of acceptable detail to satisfy radiographic imaging standards. Therefore, the recording of a high quality radionuclide (physiologic) image and a radiographic (anatomic) image with the same detector for accurate correlation of the physiologic and anatomic image remain difficult. Where very high detail is necessary, the gamma camera is generally not capable of producing better than 5 mm resolution even under the most favorable conditions. Therefore, the imaging of small parts of the body or imaging small animals like mice cannot be performed with any reasonable detail using the gamma camera. This also applies for the imaging of tissues containing radioactive materials.
 The following procedures enable the acquisition of high detail radionuclide images and the option of combining them with the x-ray radiographic images with the same detector. This approach employs a novel acquisition scheme that enables imaging spectroscopy of gamma rays, x-rays or nuclear particles by using a CCD. CCDs have been employed in the past without a scintillator for imaging spectroscopy of very soft x-rays, up to the energy levels of about 6-9 keV. However, above this energy, the CCD becomes virtually transparent to x-rays or gamma-rays. Generally, scintillators have not been used in conjunction with a CCD for imaging spectroscopy because it is believed that the conversion from gamma-rays to light will destroy the useful information carried by the interacting gamma-ray or x-ray. Therefore, imaging spectroscopy of gamma-rays or x-rays in the energy range of about 10 keV to 2,000 keV with a CCD has not been explored. Also, alternating the mode of operation from a counting, energy sensing detector to an integrating detector for radionuclide and radiography, respectively, presents a useful procedure for imaging spectroscopy. Note, however, that the counting procedure can also be used in certain x-ray transmission measurements to measure the energy thereof.
 When light interacts with the sensitive surface of the CCD, it generates a charge which remains stored in the pixel where this interaction occurred. As with previous embodiments the magnitude of the charge is directly proportional to the detected intensity of light. Each pixel is represented by its two-dimensional coordinates and by an intensity value. The energy required to produce an electron in the sensitive silicon surface of the CCD is about 3.65 eV.
 This value enables the determination of the energy of detected photons if the system can either detect one photon at a time, or if the number of the photons detected per pixel is known. This provides for imaging of radionuclide distributions with a simultaneous measurement of the energy of the detected events. This procedure is termed “Imaging Spectroscopy” and provides a technique using gamma rays, beta-rays, and x-rays in conjunction with CCD technology.
 The upper energy limit of soft x-ray imaging is between 5-10 keV. At 10 keV, the quantum efficiency of a CCD is approximately 5% and it diminishes rapidly at higher energies. The small fraction of the total number of events interacting with the CCD will result in a high partial energy transfer to the sensor with losses in proportion with the energy and the signal. Therefore, when the CCD is used as the primary detector of high energy photons or particles, it is virtually unusable for performing imaging spectroscopy. The following procedure provides high resolution imaging spectroscopy using a CCD that is suitable for many applications including position emission tomography and nuclear particle imaging.
 A schematic of the device is shown in FIG. 10. An important component of this device is a CCD 98 with low readout noise, high charge transfer efficiency and dark current levels. A CCD with less than 10 electrons/pixel (RMS) readout noise is suitable for this purpose. The dark current can be reduced to less than 0.6 electrons/sec at −40° C. by a compact thermoelectric cooler.
 In one embodiment of this method, a thin scintillator 104 is used as the primary detector of x-rays. One such scintillator can be a layer of gadolinium oxysulfide or thallium activated cesium iodide or any of the commonly available phosphors. The scintillator 104 is bonded to a fiber optic faceplate 106 and the faceplate is bonded to an image intensifier 96. The intensifier is bonded to a second faceplate 106 that is bonded to bundle 102. Optical bonding of this type is well established. To further illustrate this embodiment the sensitive area of the scintillator 104, faceplates 106, image intensifier 96, fiber optic coupler 102, and CCD 98 have, identical dimensions. Note that a collimator 94 can be mounted on the lead enclosure 100 and is used during the transmission study, and depending on its configuration, can also be used during the emission study. Note that the collimator 94 can optionally be removed during emission studies.
 When an x-ray photon within the rays 14 interacts with the scintillator 104, it produces light with intensity which is proportional to the energy of the x-ray. This light is transported through the fiber optic faceplate 106 and interacts with the CCD 98. The interaction of optical photons in each CCD pixel will produce a number of electrons in direct proportion to the number of optical photons and to the energy of the detected x-rays 14 or gamma-rays 92 that are produced by the isotope that has collected in the lesion 90. Isotopes commonly utilized include TC 99m or I-125. The following example as a first order approximation of the expected energy resolution from the detector.
 A 60 key x-ray interacts with the scintillator resulting in 3000 optical photons. Approximately one half of these photons are emitted in the direction of the CCD. Assuming a Lambertian distribution of the emitted photons from the screen, the transmission through the fiber optic plate is approximately 40%. Therefore, 600 optical photons will be arriving at the CCD. The quantum efficiency of the CCD is approximately 40%, therefore only 240 photons will be detected in one pixel.
 It can be shown that the energy resolution can be in the order of 10% which is approximately twice that attained with conventional NaI-crystal spectrometers at this gamma-ray energy.
FIG. 11 depicts an alternative embodiment in which a “pin hole” collimator 112 with shutter 110 is used in performing an emission study of lesion 90 or any selected organ. The emission from the lesion or organ impacts the scintillator 104, into housing 100, through the fiber optic reducer 116, coupled to the intensifier 118, and than directed off mirror 124, lens system 120, and onto a cooled CCD 120.
 This procedure produces radionuclide scintigraphy with spatial resolution in the order of about 1 millimeter or less, and transmission images with resolution in the order of 0.2 millimeters. The spatial resolution and sensitivity of the detector will be selectable for both emission and transmission modes via pixel binning. The detector operation will be selectable for pulse-height analysis or integration. For x-ray transmission imaging, the integrating mode of operation is preferred. Note that during x-ray transmission imaging, the pin hole collimator will be removed. Emission imaging of thick tissues requires a collimator, either a multihole type or a pinhole collimator. Very thin specimens can be imaged without a collimator by placing them very close to the scintillator.
 This camera has the capability of detecting very high count rates. In conventional gamma cameras, each x-ray photon interaction occupies the entire scintillator and electronics for a period of time of 1 to 8 microseconds after it is detected. In the present method, due to the multiple detectors, higher count rates can be handled due to the multiple detectors, and higher count rates can be handled without using a scintillator with short decay time. Count rates up to 106 counts per second can be acquired with very low probability (less than 1 %) of detecting 2 gamma ray events in one pixel when operating in the pulse-height analysis mode.
 Note the scintillator can be bonded directly on the fiber optic bundle without the use of an image intensifier. Also, the scintillator can be bonded directly on the CCD without the use of a fiber optic bundle. A frame transfer CCD is a preferred approach, but a full frame CCD can be used.
 The following “shutter” methods can be used (a) a frame transfer CCD; (b) a gated image diode, or microchannel intensifier; or (c) a liquid crystal shutter with very thin window or fiber optic window. The liquid crystal shutter can be positioned between the fiber optic bundle and the scintillator.
 Note that the system has applications for small animal imaging, skeletal imaging, monitoring of fracture healing, thyroid scintigraphy, Bremsstrahlung imaging of beta emitters within the body (radiation synovectomy), intraoperative imaging probe, radionuclide angiography, small parts imaging, and pediatric nuclear imaging.
FIG. 12 illustrates in schematic form several methods that can be used in performing quantitative imaging in accordance with the various embodiments of the invention.
 Note that one can use either a stationary source and detector to project radiation 130, or a scanning source and detector assembly to scan the object being examined 132.
 Both stationary and scanning embodiments utilize a CCD detector that transfers the detected information to a memory 140. The information can be binned or processed 142 to accomplish various tasks. This processing can include the application of software modules to correct for non-uniformities in the source or collection components, or to identify events where light from one gamma-ray interaction has spread to a number of neighboring pixels. Clusters of pixels with high intensity can be identified as primary events and low intensity clusters can be identified as scattered radiation and be eliminated by a filter.
 Quantified information such as an intensity histogram (i.e., a pulse height spectrum) can be generated 146 and a display of the object can be generated 144 with the unwanted pixels removed.
 After each set of data is produced in both the stationary and scanning embodiments, the conditions for operation can be modified 138 to produce an image at a different energy level, to perform an emission or transmission study, or to rotate the source and detector assembly relative to the object under study to produce three dimensional images or two dimensional images at different angles.
 The emission and transmission studies can be displayed alone or superimposed. Due to the binning capability of the system a one to one correspondence exists between both emission and transmission images that was previously not possible. This high resolution image can be color coded to distinguish between the emission and transmission images.
 Another preferred embodiment is illustrated in FIG. 13 where a full frame or frame transfer cooled CCD 150 with a transparent scintillator 152 bonded on the sensitive surface of the CCD, or to an image intensifier 154, as shown. The scintillator 152 is preferably emitting anywhere from the UV blue to the red regions of the spectrum upon stimulation with x-rays or gamma-rays. The preferred scintillator is one emitting in the green such as CsI (TI) or Cadmium tungstate, or alternatively a gadolinium based ceramic scintillator available from Hitachi Corporation. This scintillator has about twice the density of sodium iodide or CsI(TI) and has higher efficiency. A fiber optic plate (straight or reducing) can be incorporated between the CCD and scintillator. Alternatively, an electrostatic image intensifier 154, or image diode intensifier, can be incorporated between the scintillator and the fiber optic plate. The scintillator 152 can be optically transparent plate or comprise a fiber optic array with fibers ranging in diameter from 0.006 mm to one or more millimeters. The thickness of the plate can be in the order of 0.5 mm to 5 mm.
 Another preferred embodiment employs a CCD of the type described above but in conjunction with an electrostatic demagnifying image intensifier. The optical coupling of the CCD is accomplished by a fast lens at the output end of the image intensifier or by a fiber optic plate between the output screen and the CCD.
 The process of obtaining a desired image includes the initiation of acquisition with the CCD for about one second or at a desired binning configuration, typically coarser than 2×2 pixels. Shorter acquisition time will be required for high count-rates and longer acquisition time is tolerated for low count rates. The optimal acquisition time for a particular application can be determined empirically by acquiring a few test frames and search for coincident events within individual pixels. Very short acquisition times (less than 1 millisecond) are easily attainable by using a fast mechanical shutter, an electro-optical shutter, or by gating the image intensifier tube. This enables acquisition with spectroscopy capability even at very high count-rates. Each acquisition “frame” will record from a few hundred to a few thousand counts. After acquisition, each frame is stored in the computer memory for subsequent processing. Depending on the application, the total number of frames for a complete acquisition can vary, for example, from ten to a few hundred.
 Each gamma-ray event in a given frame stored in the computer is represented by its x and y coordinates and by an intensity value (z) which is the number of electrons generated in this area of the CCD. The z value is directly proportional to the energy of the gamma ray (or x-ray). The number of electrons generated from each interaction should be confined to one pixel or group of binned pixels forming a “superpixel”. In a significant percentage of interactions, the electrons generated from a single gamma-ray interaction can be split between two or three pixels or superpixels. These split events form clusters in the image matrix which can be easily identified by the computer software and assigned an x and y coordinate.
 In one embodiment, as shown in the process flow sequences of FIG. 14, pulse height analysis uses the value of these neighboring pixels which are summed to produce the z value for this gamma-ray event. Low z values represent gamma-rays which have been scattered and have lost a portion of their energy. These events are generally not desirable for inclusion in an image because they carry false position information. Therefore, the degree of rejection of each event can be decided by software on the basis of the z value and a spectrum of the number of gamma-rays versus the z value (energy) can be recorded. This filtering process can be repeated for each frame and all the frames can be added together to form the final image. The operator can optionally go back to each original frame, use a different z value threshold and reconstruct the final image using different filter parameters. Variations in the sensitivity of each pixel or superpixel can be mapped and included in the counter for pixel by pixel corrections. The ability to discriminate different radiation sources measured simultaneously or sequentially includes defining filter parameters as selected energy threshold values or ranges.
 In this radionuclide imaging technique, the degree of scatter rejection can be varied after the image acquisition in order to decide on the optimal scatter rejection. This is not possible with the conventional radionuclide imaging technology employing a gamma camera or a rectilinear scanner. A gamma camera or rectilinear scanner is generally incapable of detecting and processing high intensity x-rays which are employed for high quality x-ray radiography.
 If an image intensifier is not used, the scintillator can be in direct contact with the CCD. Alternatively, a fiber optic reducer can be used between the CCD and the scintillator. Typical reduction ratios vary from 1:1 to 6:1 although the present embodiment is not limited to these ratios. Therefore, for a 20 mm×20 mm CCD, and a 6:1 fiber optic reducer, the area of coverage will be about 120 mm. With a gated image intensifier or a shutter, the CCD does not receive any signal during the readout process. In a direct contact configuration, the use of frame transfer CCD as shown in FIG. 10 is preferred.
 In applications utilizing x-ray transmission measurements a single frame is acquired for the recording of the x-rays emerging from the irradiated body of tissue. The CCD is operating in the integrating mode and each pixel or superpixel which accumulates a charge which is proportional to the total number of x-rays in this region without any energy discrimination. The resulting radiographic image can be combined electronically with the radionuclide image to form an accurate representation of both physiologic and anatomic information.
 In the case of thin specimens examined in vitro a light source with wavelength ranging from the ultraviolet to near infrared can be used for the transmission image in the integrating mode. In this approach, the light shield in front of the scintillator is removed and the detector is placed in an enclosure to shield it from ambient light.
 The present invention can thus combine radionuclide emission imaging and x-ray transmission imaging (radiography) using the same area detector with spectroscopic capability in the gamma-ray imaging mode. This camera can be operated utilizing both, the counting pulse-height analysis for gamma-ray imaging, and in the integrating or counting modes for x-ray substantially transmission imaging. This enables exact superposition of the two images for accurate anatomic and physiologic imaging. Also, the operator can change the energy threshold even after the radionuclide image has been acquired. Thus, higher intrinsic spatial and energy resolution are provided than found in the conventional approaches.
FIG. 15 is a schematic illustration of one preferred embodiment of a dual-energy bone densitometry system 200 in accordance with the invention. An x-ray tube 12 emits x-rays 14 which pass through the x-ray transparent patient table 254 and into the patient (not shown). The x-rays 15 which pass through the patient are directed through an x-ray transparent mirror 202 and strike a first scintillator screen 204. The scintillator 204 reacts to low-energy x-rays and generates a light pattern corresponding to the low energy x-ray pattern. The light generated by the scintillator 204 propagates back to the mirror 202 which reflects the light to the lenses 206. The lenses 206 couple the image from the scintillator 204 to an image intensifier 208 having microchannel plates 210. Alternatively, the image intensifier 208 can be a proximity-type intensifier without the microchannel plates 210. The light from the image intensifier 208 is received and detected by the detector 212, which can be a CCD array, a CID array or an amorphous silicon sensor. The detector 212 senses the image which corresponds to the low-energy x-rays and generates an electronic representation of the image in the form of pixel data.
 High-energy x-rays pass through the scintillator 204 to an optional x-ray filter 214. The filter 214 is preferably a copper filter which blocks any remaining low-energy x-rays which pass through the scintillator 204. An optional light block filter 216 can also be included between the scintillator 204 and the x-ray filter 214 to block any stray optical radiation emanating from the scintillator 204 from reaching a second detector 220.
 The high-energy x-rays from the filter 214 strike a second scintillator 218 which is reactive to the high-energy x-rays to generate an optical image which corresponds to the pattern of high-energy x-rays. The optical image is received by a second detector 220, which can also be a CCD or CID array or an amorphous silicon image sensor. The second detector 220 senses the optical image and generates an electronic representation of the high-energy x-ray pattern. An optional x-ray absorbing fiber optic plate 222 can also be included between the scintillator 218 and the detector 220 to absorb any remaining x-rays and thus prevent them from interfering with the detector 220.
 The system 200 of FIG. 15 can be used in either a scanning mode or a stationary mode. In the scanning mode, the x-ray tube source 12 as well as the detection system are moved continuously or in a stepping motion along the region being examined. While the system scans the region, a series of images are obtained having short exposure acquisition times. In the stationary mode, a single exposure is made of the entire region being examined. Time delay integration (TDI) is used in which the CCD stores the total charge for each pixel during a selected x-ray exposure interval. At the end of the x-ray exposure, the discrete representation in each pixel is readout by a CCD controller. Once the data is thus obtained, the comparative processing techniques of dual photon absorptiometry can be used to determine quantitative density measurements of the calcified material such as bone within the body regions exposed by the x-rays.
 In the system 200 of FIG. 15, the image intensifier 208 can be omitted. In that configuration, to ensure that image data for the low-energy x-rays can be accurately collected, the detector 212 can be cooled to increase signal-to-noise ratio.
FIG. 16 is a schematic diagram of another embodiment of a dual-energy bone densitometry measuring system 300 in accordance with the invention. An x-ray tube 12 outputs x-rays 14 through x-ray transparent patient table 254 and into the patient. X-rays 15 directed through the patient strike a first scintillator 302 which is reactive to low-energy x-rays to generate an optical image of the low-energy x-ray pattern out of the patient. The optical image is carried by a coherent fiber optic conduit 304 to a CCD detector 306 which detects the optical image and generates the electronic representation of the low-energy x-ray pattern. The fiber optic conduit 304 is preferably made of plastic optical fibers to facilitate collection of the low-energy image. However, if the distance labeled “x” is selected to be small enough, glass fibers can be used instead. The space labeled 310 is filled with a film material being the same material as that of which the fibers are made.
 The high-energy x-rays pass through the scintillator 302, the fiber optic conduit 304 and the film material 310 and strike a second x-ray phosphor scintillator 312. The second scintillator 312 is reactive to high-energy x-rays and therefore generates an optical image which corresponds to the high-energy x-ray pattern. The optical image generated by the scintillator 312 is detected by a second CCD array 314 which generates the electronic representation of the high-energy x-ray pattern. An optional copper or aluminum filter 316 can be inserted in front of the second scintillator 312 to absorb any remaining low-energy x-rays. Also, an x-ray absorbing fiber optic plate 308 can be inserted between the scintillator 312 and the CCD 314 to prevent x-rays from impinging on the CCD 314.
FIG. 17 is a schematic diagram of another embodiment of a dual-energy bone densitometer measuring apparatus 400 in accordance with the invention. The system 400 of FIG. 17 is the same as the system 300 of FIG. 16 except that the coherent fiber optic conduit 304 in FIG. 16 is replaced with a different conduit 404 in the system 400 of FIG. 17. In the conduit 404 of FIG. 17, the fibers are bent at approximate right angles with small radii of curvature. As in the embodiment of FIG. 16, the fibers are either plastic or glass. Because of the different fiber bending in which the collected radiation is redirected from a first optical path onto a second optical path, the need for the film material 310 shown in FIG. 16 is eliminated.
FIG. 18 is a schematic diagram of another embodiment of a bone densitometer measuring apparatus 500 in accordance with the invention. In this embodiment, scintillator plates 505 and 507 are used to convert the x-ray energy into optical energy. Once again, the x-ray tube 12 directs x-rays 14 through the patient table 254 and the patient. The x-rays 15 emanating from the patient first strike an anti-scatter grid 502 which prevents scattered x-rays from reaching the detectors. The x-rays then strike a first amorphous silicon image sensor 504 which detects low-energy x-rays and generates the data which indicates the low-energy x-ray pattern. The low energy sensor 504 can be thinner than the high-energy sensor 508 to reduce the filtering requirements of the system. Also scintillator 505 can be thinner than scintillator 507 to improve collection efficiency of the system. High-energy x-rays pass through the first sensor 504 and then through a copper, tungsten, gadolinium or aluminum x-ray filter 506 which filters out low-energy x-rays. The high-energy x-rays then strike the second amorphous silicon image sensor 508 which generates the data for the high-energy x-ray pattern. The low-energy x-ray pattern data and the high-energy x-ray pattern data are read out of the amorphous silicon image sensors 504 and 508, respectively, by a detector controller 510.
FIG. 19 is a schematic diagram of an alternative detection structure 550 which can be used with the dual-energy bone densitometry measuring apparatus 500 of FIG. 18. The lower layer of the structure 550 is an anti-scatter grid 552 used to prevent scattered x-rays from reaching the detection structure 550. The next layer is a low energy x-ray scintillator layer 554 which generates an optical image of the low-energy x-ray pattern. An amorphous silicon image sensor 556 detects the optical image from the scintillator 554 to generate the data for the low-energy x-ray pattern. A substrate layer 558 is formed over the amorphous silicon image sensor layer 556. The substrate layer 558 includes a thinned central region 560. The thinned substrate 558 provides for increased transmission to the second scintillator layer 562. The second scintillator 562 is reactive to high-energy x-rays to generate an optical image of high-energy x-ray pattern. The optical image is detected by a second amorphous silicon image sensor 564. The structure 550 is covered by a protective substrate 566, preferably made of glass. A thin layer of lead can be formed on top of the glass to prevent propagation of x-rays beyond the structure 550. Preferred scintillators include Cs(+1), CdW04, or gadolinium oxysulfide.
 The amorphous silicon array sensors and the associated control and processing systems can utilize the binning and other processing capabilities described elsewhere in the present application. Additionally, a plurality of such sensors can be combined to form a single or dual array. The array can be linear, rectangular or square depending upon the particular application. The systems can be used in conjunction with a C-arm assembly where the C-arm 580 rigidly aligns the source 586 and detector assembly 582 as shown in FIG. 20. The C-arm 580 can also be used to rotate the source and detector about the patient on table 584 as indicated at 588 to provide multidirectional viewing of the entire human skeletal structure including the hip and femur. Thus lateral spine imaging and quantitative analysis can be conducted using the present system. The detector assembly 582 includes a CCD sensor as described herein in conjunction with a straight or angled fiber optic coupler and scintillator. The detector assembly 582 can be scanned or stepped along axis 590 and axis 592 that is parallel to the spine of the patient in order to provide a sequence of images for both quantitative and qualitative analysis.
 The detector assembly 582 can include various configurations described elsewhere herein, including the examples illustrated in FIG. 21A, 21B, and 21C. In FIG. 21A a straight fiber optic coupler 602 optically couples the scintillator 604 to the CCD (or CID or amorphous) sensor array 600. An optional cooler(s) 606 can be used in these examples. In FIG. 21B a fiber optic reducer 608 couples the scintillator 610 to the sensor array 600. A proximity type x-ray image intensifier and scintillator can replace scintillators 604 and 610. In FIG. 21C a dual sensor system includes sensors 600 and 612, fiber optic coupler 602, scintillators 618, 620, mirror 616, and lens 614. This system functions in a manner similar to that described in connection with FIG. 15.
FIGS. 22A and 22B illustrate a preferred method of imaging in which the entire imaging field is composed of a series of slightly overlapping individual images 620 that are acquired by a continuous scan or stepped imaging sequence along the rectilinear path 622. Dual energy tissue or bone density measurements can be accomplished by collecting data at two energies at each subfield 620. The x-ray source can be switched or filtered as described previously to generate discrete energy peaks.
FIG. 23 illustrates a fan-beam system in which the x-ray source 586 generates a fan shaped beam 640 that is detected by a detector system 700. System 700 can include a scintillator, fiber optic plate or reducer for each of a plurality of sensors 630 which are aligned in a linear array to collect fan beam 640. Detector system 700 can use a lead slit collimator 702 and can use CCDs, CIDs or a number of amorphous silicon sensors in configurations illustrated, for example, in FIGS. 21A-21C.
FIG. 24 illustrates another preferred embodiment 650 in which a patient 654 is positioned on table 652. X-ray tube 656 directs fan-beam 660 through a scanning slit collimator 658, the patient 654 and a second scanning slit collimator 664. The radiation 660 then passes through mirror 62 striking the scintillator 676. The scintillator emits light that is reflected by mirror 662 towards the sensor 672 as illustrated at 674. Optional lead glass element 666 can be placed at any position between the mirror 662 and the sensor 672. A lens 668 and cooler 670 can also be employed, if necessary. Lead foil 678 can be used to line the enclosure 710 to reduce interactions between the scattered x-rays and the sensor 672. The system can alternatively use a proximity type image intensifiers in the x-ray path before the mirror.
FIG. 25 is a schematic diagram of a detection system 800 which can be used with the systems described above for dual-energy bone densitometry measurements as well as tissue and lesion imaging. The system 800 can include an enclosure 802 having an aperture 804 through which radiation such as x-ray beams 806 enter the system 800. In one embodiment, the x-ray beams 806 pass through an x-ray transparent mirror 808 and strike a first scintillating plate 809. The first scintillator 809 is reactive to low-energy x-rays and produces an optical image corresponding to the low-energy x-ray pattern. The optical image is projected back onto the mirror 808 which reflects the image to lens 812. The lens 812 focuses the light onto a first surface of a CCD array detector 814. The detector 814 can include a proximity-type image intensifier to enhance image detection capabilities. An annular cooler can also be placed around the CCD detector 814 to cool the CCD and therefore improve signal-to-noise ratio.
 High-energy x-rays pass through the top scintillator 809 and strike the lower second scintillator 810 which is reactive to the high-energy x-rays to produce an optical image which corresponds to the high-energy x-ray pattern. The optical image is reflected by a second mirror 816 to a third mirror 818 which directs the light through a second focusing lens 820. The lens 820 focuses the light onto the back surface of the CCD detector 814. The back surface can also include a proximity-type image intensifier. In addition, the annular cooler, if present, cools the front and back surfaces of the CCD detector 814.
 Thus, the system 800 of FIG. 25 produces increased sensitivity by sensing on opposite sides of a single thinned CCD detector 814. Spatial correlation between the two images which are fused to form a single image is greatly improved over the previously described embodiments with separate detection surfaces since the relative locations of the image detection surfaces can be more precisely controlled. High and low energies can be detected on both sides.
 The detection system 800 of FIG. 25 can include plural two-sided CCD detectors to provide the system with a wide field-of-view along with combined electronics and cooling. FIG. 25 illustrates a second detector 834 and associated lenses 830 and 832. It will be understood that more detectors and lenses can be added as needed.
 The dual-energy configuration of the system 800 described above facilitates bone densitometry measurements as previously described. However, the system 800 can also be used for detecting and imaging lesions in patient tissue as described above. In that embodiment, the x-ray beams 806 are replaced with other types of radiation such as in the visible or infrared ranges. The scintillators 809 and 810 and the mirrors 808, 816 and 818 can be used to form images of the tissue to be formed at two different wavelengths on opposite surfaces of the detectors 814 and 834.
 A particular application of the methods and systems described herein for detecting and imaging of soft tissue lesions includes digital mammography using CCDs or similar type silicon-based detectors such as amorphous silicon type detectors described above. These systems are used to detect lesions in the tissue including calcified material within the soft tissue, that can indicate the need for more careful diagnostic procedures and/or treatment of the patient. Slot-scanning approaches using time-delay integration, where the CCD records continuously during a scan as described herein, can be used for digital mammography. However, the continuous recording approach results in certain problems, particularly with artifacts due to the shear distortion of the fiber-optic plates which can be used with such an embodiment. While slot-scan approaches using the continuous record mode can be used, the quality of the images is less than ideal due to the distortion effects.
 Other methods for scanning the breast include the dividing of the image area into four quadrants or even a greater number of segments. Every time it is necessary to take multiple exposures of the breast, the associated problems including increased exposure level and collection time can limit the variety of applications for which the system can be used. It is desirable, therefore, if one needs to acquire the image in a step-wise fashion, that there be no more than two, or at most, three acquisition steps. If one uses a greater number of acquisition steps, the breast has to remain compressed for too long, thus causing extreme discomfort to the patient. Moreover, the x-ray tube power requirements increase significantly. The preferred method for digital mammography applications involving sequential multiple imaging is thus limited to a two image acquisition process. This procedure involves directing x-rays from the source through the tissue to the stationary detector system for about 0.2-5.0 seconds and preferably in the range of 0.5-1 second, the detector system can then be moved to a second position while the first image is read out, then a second exposure is obtained and read out. As many as 2-5 million pixels can be read out in a time interval that is less than the exposure interval.
 A problem associated with the two-dimensional array approach, is its complexity and cost. Although the tiling of 4×3 CCDs, for example, to form an array can be used for digital mammography, it is likely to be too expensive for many common applications. This results from the cost of the CCDs themselves, and with the problems associated with making a seamless joint to three or four sides of the CCDs.
 Referring to FIGS. 26-32, the detector module 900 can consist of from 3 to 5 CCDs in a first linear array 902 and another set of CCDs can be positioned approximately 6 cm apart in a second linear array 904. This embodiment utilizes four CCD elements in each array. Each element can include a scintillator 906 and a tapered fiber optic plate 908. In the embodiment where the CCDs are replaced by amorphous silicon sensors, a single strip silicon sensor can be substituted for each linear array in these embodiments.
 The first set of CCDs can be placed as closely as possible to the chest wall of the patient. The x-ray beam is collimated by using a double slot to provide two fan beams, each fan beam being directed onto a linear array, thus only two areas are irradiated which correspond exactly to each CCD group. After one x-ray exposure and acquisition, the x-ray collimator is translated in synchrony with both CCD banks which are also translated to the next position. Another exposure is taken and the signal is read out. A small amount of overlapping of the fields, about 1-3 mm can be desirable. With the use of a micro-stepping translation stage the successive fields can be aligned to within a few microns with or without overlap. The images can then be joined and will be substantially seamless with less than a 5-10 micron difference between the region of the body and the joined images of the region.
 The sensing surface does not have to be on a plane. As shown in FIG. 27, the CCDs 910 can be arranged on a curving or non-planar surface. This is an extremely important embodiment because it provides for the use of straight (non-tapering) fiber-optic plates 912 which dramatically reduces the cost and contributes to better image quality. Please note that the CCDs can be cooled or non-cooled and can be operated in the pixel binned or non-binned mode. Additionally, an anti-scatter grid can be used between the breast and the detectors. Each element 910 in the array is generally equidistant from the x-ray source in order to reduce distortion across the entire field of view of the array. This arced linear array can be used for many different applications as described elsewhere herein.
 This approach is preferable as current manufacturers can readily make CCDs which are buttable on two sides. It remains difficult and expensive to make CCDs buttable on three or four sides. In the illustrated embodiment there are only six joints required between the CCDs, unlike a large area cassette which has many more. A typical CCD for this application can have an area of 6×6 cm but for economy reasons, one can use a larger number of CCDs, such as 3×3 cm elements. For example, if one uses a 3×3 cm device, each CCD linear array 902, 904 incorporates eight CCDs for a total of 16 CCDs. This can also be used to provide a larger area of coverage comparable to a standard large film cassette. FIG. 28 illustrates a four line separated array 916 covering a 24×30 cm planar area with 10 of the 3×3 cm devices in each line 918. In FIG. 29, a partial cross sectional view of one of the lines 918 in FIG. 28, illustrates a preferred embodiment in which each CCD 914 is butted against one or two adjoining CCDs in each line with each line coupled to a scintillator 915 and a fiber optic plate 917. The two-step acquisition is preferable relative to the narrow slot-scanning approaches which typically use a slot width of about 1.5 cm, and the larger area imaging approaches which are effective but can be extremely costly.
 As illustrated in FIG. 30, an x-ray source 922 and a double or multiple slot collimator 924 can be used to generate and align the x-rays 928 with the translating CCD modules. An actuator or motorized system 920 is used to translate both CCD arrays 902, 904 without altering the distance between the rigidly aligned CCD arrays. The system 920 can be connected to a controller or personal computer as described previously so that the user can control array position along the direction of translation 926, which in this embodiment, is towards or away from the chest wall.
 As shown in FIGS. 31A and 31B, the arrays 902, 904 are positioned to image two parallel regions 930, 932. The detectors 902, 904 are then translated from the first position to a second position to image and analyze two further parallel regions 934, 936 to provide full image of compressed breast 925. The relative spacing between the two linear arrays can also be controlled to increase or decrease overlap. A preferred embodiment, however, retains the two spaced arrays in a rigid position relative to each other. This particular embodiment moves the detectors towards or away from the chest wall of the patient.
 Shown in FIG. 32 is an embodiment 940 in which the array 916 of FIG. 28 in which the direction of scan 942 is along the chest wall. The collimator 944 is also moved along the same axis as the array to direct x-rays 928 onto the CCDs 948 and not onto the spaces 946.
 While the invention has been particularly shown and described with reference to a preferred embodiment thereof, it will be understood by those skilled in the art that various changes in form and details can be made therein without departing from the spirit and scope of the invention as defined by the appended claims.