JP2003265457A - Method and system for x-ray imaging - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To compose an image from measurement data in a scan beam type digital X-ray radiograph with low dose. <P>SOLUTION: An X-ray beam transmitted through an object to be imaged is formed, and X-ray intensity is measured by a detector including detector arrays DET<SB>x</SB>*DET<SB>y</SB>. From the measurement data of the respective detectors, an equivalence of data array (IMAGE) with a size of DET<SB>x</SB>*DET<SB>y</SB>and an equivalence of four-dimensional array (PIXEL) of DET<SB>x</SB>, DET<SB>y</SB>, AP<SB>x</SB>, AP<SB>y</SB>containing IMAGE data are stored in a memory. FOCUS having a numeral from 0.1 to 10.0, for OUTIMAGE having the value of the following formula 1, an output value for forming an equivalence of two-dimensional array with a size of IP<SB>x</SB>*IP<SB>y</SB>is created and output. <P>COPYRIGHT: (C)2003,JPO

Description

Detailed Description of the Invention

[0001]

BACKGROUND OF THE INVENTION 1. Field of the Invention The present invention relates to a diagnostic X-ray imaging apparatus. More particularly, the invention is segmented with multiple collimation grids with openings.
A real-time scanning beam digital radiography system with improved resolution and reduced x-ray radiation dose by incorporating an x-ray detector array.

[0002]

2. Description of the Related Art With the development of therapeutic techniques, real-time radiography in medical procedures is increasingly desired. For example, many electrophysiological procedures in cardiology,
Peripheral vascular, urological and orthopedic procedures rely on real-time radiography. Unfortunately, current clinical real-time x-ray machines deliver high levels of x-rays to both patients and participating staff. US Food and Drug Administration (FD
A.) reports anecdotal evidence of acute radiation sickness in patients and the problem of occupationally overexposed physicians (Radi).
logical Health Bulletin,
Vol. 26, No. 8, August 1992).

Conventionally, many real-time X-ray imaging systems are known. These are equipped with a system based on X-ray fluoroscopy, which irradiates the object under examination with X-rays,
The shadow caused by the X-ray opaque body in the object is displayed on the X-ray fluoroscope located on the opposite side of the X-ray source with respect to the object. Scanning X-ray tubes have been known in connection with fluoroscopic techniques since at least the early 1950s. Moon, "Amplification and Enhancement of Fluoroscopy with Scanning X-Ray Tubes," S
science, October 6, 1950 issue, 389-39.
See page 5.

Scanning beam digital radiography systems are also well known in the art. In this system, an X-ray tube is used to generate X-rays. An electron beam is generated inside the X-ray tube and focused on a small spot on a relatively large anode (transmission target) of the tube to generate X-rays from this spot. The electron beam is deflected (electromagnetically or electrostatically) into a raster scan pattern across the anode. A small X-ray detector is arranged at a predetermined distance from the anode of the X-ray tube. The detector converts the X-rays impinging on it into an electrical signal proportional to the detected X-ray flux. When the object is placed between the X-ray tube and the detector,
X-rays are attenuated and dispersed by the object according to the X-ray density of the object. If the X-ray tube is in scan mode,
The signal from the detector is modulated in proportion to the X-ray density of the object.

An example of a conventional scanning beam digital X-ray system is shown in Albert US Pat. No. 3,943.
9,229, Albert U.S. Patent No. 4,032,787, Albert U.S. Patent No. 4,057,745, Albert U.S. Patent No. 4,144,457, Albert U.S. Patent No. 4,149,076,
Albert U.S. Patent No. 4,196,351, Albert U.S. Patent No. 4,259,582, Albert U.S. Patent No. 4,259,5
83, Albert U.S. Pat. No. 4,288,697, Albert U.S. Pat. No. 4,321,473, Albert U.S. Pat.
4,323,779, Albert U.S. Pat. No. 4,465,540, Albert U.S. Pat. No. 4,519,092, and Albert U.S. Pat. No. 4,730,350.

[0006]

[Patent Document 1] US Pat. No. 3,949,229

[Patent Document 2] US Pat. No. 4,032,787

[Patent Document 3] US Pat. No. 4,057,745

[Patent Document 4] US Pat. No. 4,144,457

[Patent Document 5] US Pat. No. 4,149,076

[Patent Document 6] US Pat. No. 4,196,351

[Patent Document 7] US Pat. No. 4,259,582

[Patent Document 8] US Pat. No. 4,259,583

[Patent Document 9] US Pat. No. 4,288,697

[Patent Document 10] US Pat. No. 4,321,473

[Patent Document 11] US Pat. No. 4,323,779

[Patent Document 12] US Pat. No. 4,465,540

[Patent Document 13] US Pat. No. 4,519,092

[Patent Document 14] US Pat. No. 4,730,350

[Non-Patent Document 1] FDA, Radiological Health Bul
letin, Vol. 26, No. 8, August 1992

[Non-Patent Document 2] Moon, Science, October 6, 1950, pages 389-395.

[0007]

In a typical prior art scanning beam digital X-ray system, the output signal from the detector is applied to the Z-axis (luminance) input of the video monitor. This signal adjusts the brightness of the screen. The x and y inputs to the video monitor are derived from the same signal that deflects the x-ray signal of the x-ray tube. Therefore, the brightness of a point on the screen is inversely proportional to the absorption of X-rays that reach the detector from the X-ray source through the object.

Medical x-ray systems operate at the lowest levels of x-ray dose that meet the resolution requirements for the procedure. Therefore, radiation dose and resolution are limited by the signal-to-noise ratio. As used herein, the term "low radiation dose" means an x-ray exposure level of about 2.0 R / min or less that reaches the patient during operation.

The time and area distributions of X-ray photons follow the Poisson distribution and have unavoidable randomness. Randomness is expressed as the standard deviation of the mean radiant flux and is equal to its square root. The signal-to-noise ratio of an X-ray image under these conditions is therefore equal to the mean radiant flux divided by the square root of the mean radiant flux, with a noise of ± 10 for an average radiant flux of 100 photons.
For photons, the signal to noise ratio is 10.

Therefore, the spatial resolution and signal-to-noise ratio of an X-ray image produced by a scanning radiography system are highly dependent on the size of the sensitive area of the detector. As the detector aperture area is increased, more divergent light is detected, the effective sensitivity is increased and the signal to noise ratio is improved. However, at the same time, as the "pixel" size (measured in the plane of the imaged object) increases, the large detector aperture reduces the adjustable spatial resolution. This is inevitable because most objects imaged in the medical field (eg, the internal structure of the human body) are some distance from the X-ray source. Therefore, in the prior art, the detector aperture size had to be chosen to balance resolution and sensitivity, and it was not possible to maximize both resolution and sensitivity at the same time.

In medical imaging technology, the amount of irradiation to the patient, the frame rate (the number of times per second at which the object is scanned to update the image) and the resolution of the image of the object are key. Is the parameter. High X-ray flux easily provides high resolution and high frame rate,
The patient and the participating staff will be exposed to unacceptably high X-rays. Similarly, the image is not visible and updated (r
Even if the efresh) rate is insufficient, the irradiation dose can be reduced by allowing it. A preferred medical imaging system should simultaneously achieve low dose, high resolution, and a sufficient update rate of at least about 15 images per second all at the same time. Therefore, a system such as the scanning beam digital radiography system described above keeps the x-ray dose received by the patient at a minimum, although the exposure time is relatively long and the same can be said for real patients at all times. Not suitable for diagnostic medical procedures that must be done.

[0012] Therefore, it is an object of the present invention to provide a scanning beam type digital X-ray imaging system which can be used in a medical diagnostic procedure which can be safely received by a patient.

Another object of the present invention is to provide a high resolution image at a sufficient frame rate and to minimize the irradiation amount of X-rays on an object to be inspected.
It is to provide a radiography system.

Yet another object of the present invention is to provide a scanning beam digital radiography system having improved resolution at sites spaced from the plane of the x-ray source while reducing the x-ray flux level.

[0015]

A scanning beam digital X-ray imaging system (SBDX) according to the present invention comprises an X-ray tube having an electron beam source and a target anode. A circuit is provided for focusing the beam and directing or scanning the beam through the target anode into a predetermined pattern. For example, the predetermined pattern may be a raster scan pattern, a curved or S-shaped pattern, a spiral pattern, a random pattern, a Gaussian distribution pattern centered at a predetermined point on the target anode, or a short task. Other patterns that are effective for

A collimating element (preferably formed in a grid) may be interposed between the X-ray source and the object to be irradiated with X-rays. The collimator has, for example, a diameter of about 25.4 centimeters (10 inches) and 500 lines 5 in the center of the collimator.
It comprises a circular metal plate having an array of 00 openings.
The collimation element is preferably arranged immediately before the light emitting surface of the X-ray tube. Other shapes of collimating elements can also be used. In one preferred embodiment of the invention, each of the apertures of the collimation element is arranged to be directed towards (or towards) a detection point on a surface located at any distance from the collimation element. ing. The distance is selected so that the object to be irradiated with X-rays is arranged between the collimation element and the detection point. The function of the collimation element is to form thin beams (arranged like pixels) of X-rays, all directed from one point on the anode of the X-ray tube towards the detector.

An array of detector elements (preferably DET _X
The center of a segmented detector array having an array of areas such as a × DET _Y rectangle or square, and more preferably a relatively round array, is located at the detection point.
The detector array preferably comprises a plurality of closely packed X-ray detectors. Such arrays, in accordance with the present invention, are designed, positioned, and used to provide high sensitivity without compromising resolution, and thus have resolutions comparable to or better than conventional x-ray equipment. However, it is possible to realize an X-ray system having an irradiation amount that is at least one order of magnitude smaller than the irradiation amount of the conventional X-ray system. This feature of the invention has important implications in medicine and other fields. The dose to the patient or participating medical staff in the current procedure is reduced. The risk of violence makes possible treatments not currently possible.

The output of the detector array is the intensity value for each element of the detector array as the x-ray beam is emitted through the apertures in the grid. Since each aperture is located at a spatially different point with respect to the object under examination and the detector array, the detector array provides a different output for each aperture through which the X-ray beam passes. The output of the detector array can be converted into an image in various ways. One method is to simply combine the array outputs, ie, sum the intensity values of the array elements corresponding to each scanned aperture and further normalize. The output array is then
It can be used to drive a video or other display. Even more preferred are the multiple image compositing and multi-output compositing methods described below, which provide improved visual output.

The scanning beam digital X-ray imaging system is also capable of stereoscopic imaging using a collimation grid having two groups of apertures. In this case, the first of the openings
The swarm is directed to a first detection point where the split first detector is located, while the second swath of apertures is directed to the second detection point where the split second detector is located. .
A stereo image is created by forming the two images with two segmented detectors and using conventional stereo display methods.

Scanning beam digital radiography systems can also produce enhanced images of materials that exhibit different x-ray transmissions at different x-ray photon energies. Therefore,
For example, the precursor of breast cancer, microcalcification, can be imaged. By configuring the grid and / or the anode to emit two or more groups of X-ray beams, each with a different X-ray energy spectrum, and directing each group to a detector array (multiple detector arrays can also be used) Since it is possible to image that the transmittance of the inspection object is different for various X-ray photon energies, it is possible to emphasize only the portion showing the different X-ray transmittance inside the inspection object. Optimized for calcium detection, for example, the imaging system is a powerful tool for early detection of breast cancer and other abnormalities.

By utilizing a split array that blocks the collimated total X-ray beam and imaging the array output, the resolution obtained using a small surface area detector is not sacrificed. Maximum sensitivity can be provided. A non-segmented detector of the same size as a segmented array has the same sensitivity but low resolution.

Further, the sub-sampling method may be used to process the data from the array detector, providing substantially the same image quality while reducing system complexity, required processing speed and energy consumption. To do.

The system described here is 199
US Patent Application No. 08/00 filed January 25, 3rd
"X described in No. 8,455 (CAM-003).
It may be used with a catheter equipped with a line-sensitive photosensor detector, the contents of which are incorporated herein by reference. This US Patent Application No. 08 / 008,455
Issue is owned by the assignee of the present application.

[0024]

BEST MODE FOR CARRYING OUT THE INVENTION Embodiments of the present invention will be described below with reference to the accompanying drawings. In the drawings, the same reference symbols indicate the same or equivalent parts.

Device Overview FIG. 1 shows a scanning beam digital radiography system according to a preferred embodiment of the present invention. Scanning X-ray tube 1
0 is used as the X-ray source. As is well known in the art, a power supply of approximately 100 KV to 120 KV is used for the X-ray tube 10.
Used to activate the. Power supply of 100KV is 1
An X-ray spectrum extending to 00 KeV is provided. Here, 100 KV X-ray refers to this spectrum. X-ray tube 10
Scan generator 30 as is well known in the art.
The deflection coil 20 is controlled by the following. X-ray tube 1
The electron beam 40 generated in 0 scans from one side of the grounded anode 50 to the other side in a predetermined pattern in the X-ray tube 10. For example, the predetermined pattern is a raster scan pattern, a meandering or S-shaped pattern, a spiral pattern, a random pattern, a Gaussian distribution pattern centered on a predetermined point of the anode of the target,
Or any other pattern that is convenient for near future tasks. Presently preferred are serpentine or sigmoidal patterns that eliminate the need for "retrace" in raster scans.

The electron beam 40 is directed to the anode 5 at a point 60.
At 0, a cascade of X-rays 70 is emitted and travels out of the X-ray tube 10 towards the object 80 examined by X-ray. To maximize the operation of the instrument, cone-shaped X-ray photons are generated that radiate to just cover the detector array 110. This is preferably accomplished by placing a collimation grid between the anode of the scanning x-ray tube and the detector. Therefore, the collimation grid 90 includes the object 80 and the X-ray tube 1.
It is arranged between 0 and. The collimation grid 90 is designed so that only X-rays 100 directed to the detector 110 pass through it. The collimation grid 90 does not move with respect to the detector array 110 while the device is operating. Thus, whenever the electron beam 40 scans the anode 50, there is only a single x-ray beam 100 passing from the anode 50 to the detector array 110.

FIG. 2 shows the X-ray distribution when there is no collimation grid. The output of the detector array 110 is processed and displayed on the monitor 120 as a brightness value at the x, y position on the monitor 120 corresponding to the x, y position of the anode 50. This is accomplished by using the same scan generator that drives the x, y position of the electron beam 40 and the electron beam within the video monitor 120. Also, image processing techniques can be used to create computer-operated images on any suitable display or photographic medium.

The inventive device disclosed herein generally measures from about 0.15 R / min at an update rate of 15 frames / sec to 30 frames / sec, measured at the portion incident on the patient. It is a low radiation dose device that delivers doses up to about 0.33 R / min to patients. With this device, 30 frames /
An update rate of seconds would be about 0.50 R / min for the whole body. The actual range of radiation dose at the patient surface for the operation of the invention disclosed herein is from 0.15 R / min to 2.0.
Up to 0 R / min.

X-Ray Tube FIG. 3 shows an enlarged structure of the grid and anode. The anode 50 is preferably formed on a beryllium anode support 130 and assembled with a target layer of material that has good vacuum properties and the ability to withstand high temperature and electron bombardment. Additionally, aluminum or other material that is relatively transparent to X-rays can be used for the anode support 130. Presently preferred target layers are, in order of preference: (1) a second layer of tantalum sputter-deposited at a thickness of approximately 5 microns on a first layer of niobium sputter-deposited at a thickness of approximately 1 micron on the anode support (this The structure is such that niobium has a coefficient of thermal expansion intermediate between that of beryllium (anode support 130) and tantalum, thus suppressing microcracking due to thermal cycling of the target anode between the on and off states of the tube. (2) approximately 5 microns thick tantalum sputter deposit layer, (3) approximately 5 microns thick tungsten-rhenium sputter deposit layer, and (4)
It is a sputter deposited layer of tungsten approximately 5 to 7 microns thick. Tantalum, tungsten and tungsten-rhenium are preferred for the anode 50 because they have relatively high atomic numbers and densities and readily emit X-rays when irradiated with an electron beam. The fact that tungsten has a high melting point of 3370 ° C. and good vacuum characteristics are suitable for the high temperature and ultra-high vacuum state in the X-ray tube. Tantalum and tungsten-rhenium have similar properties as known to those skilled in the art. The thickness of the anode layer is 100 KV
It is chosen to be approximately equal to the distance required to effectively convert the electrons into x-rays. Beryllium is preferred for the anode support 130 because it is strong and does not significantly reduce or dissipate the X-rays emitted from the anode 50.
The thickness of the beryllium anode support 130 is about 0.5c.
It is preferably m. The anode support 130 is subject to the physical constraint that it must be strong enough to withstand a pressure gradient with an external atmosphere of 1, but should be as thin as possible.

The collimation grid 90 comprises an array of apertures 140, each oriented or directed toward the detector array 110, according to a preferred form of the invention. That is, the apertures in the collimation grid 90 are not parallel to each other, and in use such as, for example, chest x-ray applications, with respect to the front surface 260 of the collimation grid 90, from 0 ° in the center of the collimation grid 90 to the grid 90. The angle at the end can be between about 20 °. In use of the present invention in a breast examination application, the grid 90 may be assembled such that the openings make an angle with the front surface at an end of the grid in the range of 45 °. The number of openings 140 in the collimation grid 90 is, for example, 500 × 500 to 10 at the center of the round collimation grid 90.
It can correspond to 24 × 1024 pixels and can determine the resolution of the device. Also, apertures with less than the number of pixels can be used in connection with the subsampling technique described below. The thickness of the grid 90 and the dimensions of the apertures 140 are such that the distance from the x-ray tube 10 to the detector array 110 (here preferably 91.4 cm (36 inches)) and any that are not directed to the detector. Because of the need to reduce X-rays, detector element 1 of detector array 110
60 (not shown in this figure). The front surface 2 is not important in the present invention.
The openings 140 viewed from 60 are preferably designed in a rectangular column and row pattern with circular boundaries having a diameter of 25.4 cm (10 inches). The aperture array can be a convenient layout relative to each other using the detection and convolution techniques outlined below for analyzing an image of the object 80. This array of apertures is called the "circular active area". At the center of the circular active area, the total number of openings is, according to a preferred embodiment of the invention, preferably 500 × 500. The unopened portion 150 of the collimation grid 90 has a misaligned X
It is designed to absorb stray X-rays so that the rays do not illuminate the object 80. This means that the X-rays that impinge on the non-aperture portion 150 have a "1/2 value" (at the energy of this device, here 100 KeV, the X-rays that impinge on it are 1/2).
This is done by assembling the grid to show at least 10 times the amount of material required to reduce The displaced X-rays irradiate the object and staff with X-rays but do not provide useful information to the image. As shown in FIGS. 3A and 3B, the collimation grid 90 has multiple sheets of X-ray absorbing material 14 having apertures 140 therethrough for passing the X-ray beam 100 to a detector.
3, 144 can be formed. The collimation grid 90 is preferably assembled by stacking 50 thin sheets of 0.0254 cm (0.010 inch) molybdenum and holding them together. Molybdenum is preferred because it stops the x-rays generated in the x-ray tube 10 that are not directed to the detector 110 before they have a useless and harmful impact on objects 80, including of course human patients. Lead or similar X-ray dense material can also be used.

The opening 140 of the collimation grid 90
Is preferably square in cross section for maximum packing density and to match the preferred square shape of detector array element 160. Other shapes, especially hexagons, can also be used. The square aperture 140 is about 1 of the cross-sectional area of a conventional collimator used in fluoroscopy.
0.0381 cm (0.
A size of 015 inches) × 0.0381 cm is preferable. This tighter collimation achieves a smaller beam width for the x-ray beam 100. This means that the cross-sectional area of the surface of the detector can be correspondingly smaller than in conventional devices. As a result, the X-rays scattered by the object are not detected by the detector, and it is possible to obscure the image as the scattered X-rays acted on in conventional devices using detectors of relatively large surface area. Absent.

The preferred method of assembling the collimation grid 90 is by photochemical milling or etching. Photochemical milling is preferred because it is cost effective and accurate. According to this method, 0.0254
A set of 50 photomasks is made to etch holes and gaps in 50 sheets of thin sheet material, which is 0.010 inch thick. The etched sheets are stacked, aligned and held together to form a grid assembly having a number of stepped openings each having a predetermined angle for each sheet. FIG. 3A shows a modification of the original collimation grid 90.
This variant includes a number of X-ray absorbing sheets 143 having individual openings of constant cross section (but not necessarily constant cross section). The resulting opening 14 is
As shown, the stepped but x-ray beam 100 is passed to a detector array 110. The modification shown in FIG. 3B is quite similar to that shown in FIG. 3A except that the individual openings formed in the X-ray absorbing sheet 144 are themselves stepped. These stepped openings can be made by milling or chemical etching from each side of the sheet 144 with slight staggers to the shape shown, as will be apparent to those skilled in the art. The shape of FIG. 3B has a collimation grid 90.
Further, since there is less absorption of X-ray energy in the stepped aperture 140 of the X-ray beam 140, and consequently the X-ray flux at the ends of the X-ray beam 100 is not reduced to the same extent as the variant shown in FIG. 3A. preferable.

One preferred method of holding the etched sheet (s) forming the grid 90 is shown in FIG.
4 is shown. Each etched sheet 91 (preferably 50) is provided with an alignment hole or alignment opening 94, respectively. Alignment peg 95
Are placed in respective alignment openings 94 to align the etched sheet 91. Sheet 91
Aluminum ring 35 for peg 95 assembly
Located within 9. The aluminum ring 359 is provided with a vacuum port 370 that can be sealed with a pinch-off 375. And an aluminum sheet 365 with a thickness of 0.1 cm
Are bonded and sealed to the upper surface 380 of ring 359 with a vacuum adhesive. Aluminum sheet 360 is similarly adhered to lower surface 385 of ring 359. Port 3
A partial vacuum is pulled through 70, and port 370 is sealed at pinch-off 375 as is known in the art. In this manner, the relatively X-ray transparent aluminum sheets 360, 365 provide a clamping action that helps hold and orient the etched sheets 91 as a grid assembly 90.

The opening 14 farthest from the center of the grid 90
0 has a stepped surface and is preferably square in cross section. X-rays are generally unaffected by the unevenness of the passageway due to the stepped surface, and even if they scatter, they will not measurably affect the resultant beam. As mentioned above, the material used for the collimation grid 90 is currently preferably molybdenum, brass, lead, or copper with molybdenum. The preferred error for the hole position is ± 0.00127 cm (0.0005 inch) from center to center, excluding cumulative error, and the hole size error is ± 0.00254 cm (0.005 cm).
01 inches).

Other methods of making usable collimation grid 90 include electron beam machining, drilling, mini-machining, laser drilling. Drilling and laser drilling have the drawback of producing round holes rather than square holes. Circular openings work well but are presently not preferred.

More details of the preferred scanning x-ray tube 10 are shown in FIGS. The electron gun 161 is located on the opposite side of the front surface of the X-ray tube 10 from about -100 KV-
Operates at potentials up to 120 KV. The grounded anode 50 is located in front of the tube and the electron beam 40 is
1 and the anode 50. The grounded electron aperture plate 162 is located near the electron gun 161 and
It has an opening 163 at the center where 0 passes. The magnetic field focusing lens 164 and the deflection coil 20 position the beam spot on the anode 50 using automatic focusing well known in the art. The tube is 25.4 cm (10 cm) across which the electron beam 40 intersects the anode 50 with the electron beam being deflected by about 30 ° at the end of the circular active area.
Assembled to have a circular active area of diameter (inch). When the beam is not "fired" through an aperture, it is preferable to stop producing the beam, resulting in power savings of up to about 25%.

FIG. 5 shows a cross-sectional view of the front portion of a suitable x-ray tube 10. The inside of the X-ray tube 340 that is maintained in vacuum is behind the anode 50. Anode 50 is a coating of anode material as described above. The front of the anode 50 is a 0.5 cm thick beryllium anode support 1
Thirty. The front of the beryllium anode support 130 is a cooling jacket 350, which is preferably 0.4 cm thick and is adaptable for flowing water or forced air. The aluminum grid support 360, 365 is 0.1c
It is etched to a thickness of 1.27 cm (0.
Helps support a 5 inch thick collimation grid 90.

When the X-ray tube 340 is used, only one opening 140 in the collimation grid 90 will pass a significant amount of X-rays at any given moment.
According to a preferred embodiment, the electron beam 40 can be stopped when the electron beam 40 is not located directly in front of the aperture 140. Thus, the x-ray tube can be effectively operated in scan pulse mode to reduce power consumption and reduce wear and tear on the target anode 50.

Stereoscopic Radiography In FIG. 6, according to another preferred embodiment of the present invention, a grid with a plurality of focal points can be provided so that a stereoscopic radiography can be obtained. For example,
If all other rows of apertures in grid 90 are directed to focus F1 (92), the remaining apertures are focused to focus F2 (9).
If directed to 3), it is possible to scan the aperture in a raster or serpentine pattern by placing the first sensor array at F1 (92) and the second sensor array at F2 (93). Yes, thereby creating a "line" of data for the first sensor array and a "line" of data for the second sensor array. Repeating this, two completed images can be assembled so that they can be seen from two different points F1, F2 in space, and thus in a conventional stereoscopic display device that provides a stereoscopic X-ray image. You can display them. Figure 3
C shows how to construct a three-dimensional collimation grid from within the layer 144 of X-ray absorbing material.
In this case, the apertures 140A, 140B are "V""legs" for the X-ray beams 100A, 100B, as shown.
It can actually be shaped like a "V", providing separate passages along. However, although it is not a requirement that the openings 140A and 140B be connected as shown in the figure, the advantage of the "V" -shaped opening in which X-rays are incident at the apex of the "V" is that both detections are The instruments are illuminated at the same time and the "V" is to act as a signal separator for the X-rays going to F1 and F2. This halves the power required for the beam and deflection current. The price is lower, but the scatter and the resulting image haze is increased.

Array Detector In order to achieve a resolution of several lines / mm on the object plane,
As required in some medical applications, spatial resolution limits are primarily determined by detector size.
This is because in today's x-ray tube technology it produces very high power levels which would be required to get very well directed x-ray radiation, and in the combined x-ray direction. Because neither of the means of attachment is possible.

If the detector is made smaller than the area that intersects the cone of emitted X-rays, as shown in FIG. 7A,
A large proportion of the X-rays emitted by the source 50 are not detected by the detector 250. In practice, such industrial beam scanning digital X-ray inspection apparatus is designed in this way, where the irradiation dose is normal and not a problem. As a result, the irradiation dose is increased to maintain the desired resolution.

Thus, resolution is improved with the use of a smaller detector, but the area of the detector is limited to the detector plane 270.
The X-ray dose is minimized when it is equal to or exceeds the area defined by the cone of radiant X-rays intersecting with.

The resolution of the scanning X-ray imaging apparatus is the object plane 2
It is determined by the cross-sectional area of the detector element projected at 80 (the plane perpendicular to the line between the center of the anode 50 and the center of the detector 110 where the object 80 is located). Thus, as shown in the front view of the detector array in FIG. 8, if a large area detector is divided into smaller array elements, a large capture area of the detector assembly is maintained while at the same time. , Keep the image resolution proportional to the size of each individual small detector element 160.

The resolutions defined by the individual detector elements 160 sum up from each of the individual elements, with each address or pixel distributing the reading to a memory buffer corresponding to a particular location on the object plane 280. Maintained by. The X-ray beam 100 has a collimation grid 90 located in front of the X-rays that are emitted to the anode 50.
As the output of a given detector element is summed changes. The imaging geometry is shown in FIGS. 7B and 7C. FIG. 7B shows how a single beam position is divided into 5 pixels. Figure 7C shows how successive beam positions are added to each other within a single pixel.

In other words, the signal for each detector element is stored in the image buffer at a memory area corresponding to a very small specific area of the object plane 280, ie a single pixel. Thus, the memory storage address for each detector element changes to the position of the scanning x-ray beam in a configuration where each pixel of the memory is arranged to contain the sum of the radiation that has passed through a particular spot on the object plane 280. In this way, the resolution of the device is determined by the size of a single detector element, while the sensitivity of the device is maximized because substantially all X-rays reaching the detector plane 270 are recorded.

A further advantage of the imaging geometry of this array detector is that the object plane 280 is narrowly defined. Structures in front of it or behind it become hazy (out of focus). First opening 141 and second
X-rays from the opening 142 pass through the object surface 280 at a distance SO from the openings 141, 142, and the openings 141, 1
X-rays that have passed through surface 281 at a distance SO from 42 are shown in FIG. 7D. As can be easily seen, the resolution obtained with twice the SO falls to about 1/2 of the resolution obtained with SO. This feature provides improved placement and visualization of detailed structures on the surface of interest 280, while providing sufficient depth of field that can be modified by the geometry of the device.

The array of the presently preferred embodiment is 96 of 0.135 cm square detector elements arranged within a circle of 0.72 inch diameter.
It is an element pseudo-circular array. It does not have to be as large as this; the side length of one detector, here 0.135 cm.
It is also possible to have more than two detectors arranged so that they are not all aligned in a circle of radius equal to.

X-Ray Detector Conventional image intensifier technology has the fundamental limitation of system sensitivity. The thickness of this usable scintillator material is limited by its light transmission properties. Typically it is made thick enough to trap about 50 percent of the incident X-ray photons. Only half of the emitted photons reach the photocathode. At the photocathode, only about 10 percent of the incident photons produce photoelectrons. Thus, only the incident X-ray photon energy of 2.
Five percent (0.5 x 0.5 x 0.1) is maintained in the image intensifier system. In addition to this limited conversion factor, photons are scattered vertically by the scintillator material, causing blurring that reduces the resolution of the system at a given dose level.

One of the underlying aims of the invention is to ensure that the object under examination is exposed to the lowest possible x-ray level while achieving the appropriate image quality required for the procedure to be performed. It is to provide an SBDX imaging system. This means that the system used to detect the X-ray photons emerging from the object must have the highest photon-to-electrical signal conversion efficiency. To achieve this, the material used in the detector must be long enough to ensure that the length in the direction in which the photons travel does not enter the photons from the edge far away from the incident X-ray. I won't. For example, photon energy must be properly dissipated in the material in order to maximize the detector output. There are many types of detectors that can be used in the SBDX system described here. Preferred is X
A scintillator in which line photon energy is converted into visible energy and light intensity is converted into an electrical signal by a photomultiplier tube, photodiode, CCD or such device means. The scintillator material must have high responsivity and minimum afterglow time because each pixel of the SBDX image must occur in a very short time of about 140 nanoseconds. Afterglow is a phenomenon in which the scintillator continues to emit light after the excitation incident X-ray emission is completed. A plastic scintillator such as polystyrene mixed with an organic substance is suitable in that it has the necessary high responsivity, but it has a relatively small X-ray photon interaction cross section and also has a small linear X-ray absorption coefficient. It is a value. As a result, a considerable thickness is needed to stop the X-ray photons. 1
For a 00 kV x-ray, a typical plastic scintillator is 28 cm to capture 99% of the incoming x-ray
A thickness of (11 inches) is required. More preferably here (in a preferred choice): (1) cerium-doped YSO (yttrium oxy-orthosilicate, available from Airtron (Litton), Charlotte, NY); (2) cerium-doped L.
SO (lithium oxy-orthosilicate, available from Schlumberger), (3) BGO (bismuth
Germanite, available from Lexington Components of Beachwood, Ohio). YSO and LSO are excellent in that they can be used at room temperature.
BGO must be heated to about 100 ° C. to achieve a suitable light output decay time on the order of 50 nanoseconds. These scintillator materials do not require the same length as plastic scintillators and are effective at lengths of 0.10 cm.

In accordance with the presently preferred embodiment of the present invention, the SBDX array detector 110 includes x-ray sources 50-9.
Placed at a distance of 1.4 cm (36 inches),
It consists of an array of 12x12 pseudo-circles of 96 closely packed individual x-ray detectors 160. (5x5, 3x
The 3 array is also considered as a non-square array with square detectors (filling the circle around the x-ray target). For example, see the bottom of Table 1. ) The shape of the collimation grid aspect ratio, the distance from the X-ray source 50, and the size of the X-ray source 50 are such that the total included angle of about 2.23 cm (0.9 inches) in diameter at the entrance of the detector array 110. Give a quadrangular pyramid that is 1.46 °. Each scintillator 1
70 should therefore be about 0.06 inch from center to center in the plane of the detector. If the scintillator 170 had parallel sides, X-rays incident near the edges would not be able to travel the required distance without hitting the scintillator wall. Therefore, these x-rays pass to nearby scintillators, and if they are not shielded, they produce an apparent output from the wrong spatial location, resulting in a consequent deterioration of the image quality of the object. To avoid this effect, as shown in FIG. 9A, each scintillator according to a preferred embodiment of the present invention is preferably tapered and the interface 173 is at the extreme end of the incident X-ray 100 '. It has an included angle equal to the angle. This is especially useful for long plastic scintillators. In the preferred example cited above, each scintillator 170 has a diameter of 0.28.
It is a preferred quadrangular pyramid 28 cm long with a 5 cm entrance face 172 and a 0.37 cm diameter photodetector end face 174. Each of the 81 detector bundles therefore has polyhedral ends, each face adjoining the surface of a sphere centered in the X-ray source 50.

A further improvement in scintillator detection efficiency is achieved by the scintillator taper at an angle greater than that of the incident X-ray beam 100. Near the edges of the scintillator, photoelectrons and scattered x-rays generated by the interaction between incident x-rays and scintillator atoms can be lost to the shield material that preferably separates adjacent scintillators. The lost photoelectrons do not generate any light and do not behave as light emitted from the scintillator.
Therefore, such extinction reduces the efficiency of the scintillator. The maximum distance traveled by a photoelectron depends on its energy and the material it travels. In the interaction of electrons with 100 kV X-rays in a plastic scintillator,
No photoelectron can travel a distance longer than about 0.01 cm. If the four-sided frustum of the scintillator is the X-ray beam 10
If made with an included angle greater than 0, its size is a short distance, 2 compared to the detector length (28 cm).
It becomes larger than the beam enclosed by × 0.01 cm, and the reduction in efficiency due to lost photoelectrons is minimized. In this case, the center of the sphere tangent to the surface no longer coincides with the x-ray source 50 and is closer to the detector array 110.

Scattered photons travel longer distances than photoelectrons. To prevent these from escaping to adjacent scintillators, the value of the scintillator pyramid taper is set to be greater than that required for complete photoelectron capture to maximize trapped scattered photons.

Referring now to FIG. 9, in accordance with a preferred embodiment of the present invention, each scintillator element 170 is contacted to optically attach each scintillator element 170 to a corresponding photomultiplier tube 190 or solid state detector. A mating light pipe or fiber optic cable 180 is provided. The alternative scintillator 170 may be placed in physical proximity to the particular photodetector.

FIG. 10 shows the preferred shape of the detector element 160. The X-ray impermeable sheet 200 having the openings 210 corresponding to the respective detectors 160 is used as the detector array 11.
It is placed before 0. Each detector element 160 is surrounded by an enclosure 220 that does not leak light and is impermeable to X-rays. Light blocking window 230, which is preferably formed of an aluminum plate
Is placed in front of the light tight enclosure 220. The light blocking window 230 transmits X-rays. In the light-tight enclosure 220, the scintillator element 170 has a preamplifier 240
It is located near the photomultiplier tube 190 electrically connected to. Preferably, the analog signal from preamplifier 240 is converted into a digital signal in a conventional manner by further processing.

Alternatively, the scintillator can be placed directly on or near an array of photodiodes, phototransistors, or solid state imagers (CCDs) to obtain a more crude and compact detector. Cooling with a Peltier cooler can be used to increase the signal-to-noise ratio of the device, especially when a solid-state device such as a CCD is used.

Alternatively, a scintillator array may be placed at or near the photomultiplier tube which, as well as its size, identifies a position comparable to the light source and detects one or more positions that provide an output signal. can do.

In another preferred embodiment, the sensor array may consist of an array of aligned pencil-type detectors 285, for example as shown in FIG. In FIG. 11, the tapered scintillator 290 is aligned with the path of the x-ray beam 100, and the scintillator corresponding to a particular cross-sectional area of the beam 100 absorbs all x-rays within that cross-sectional area. The optical amplifying tube 300 is provided physically close to the scintillator 290, and an electric signal is generated corresponding to absorption of X-rays by the scintillator 290. A solid-state device can also be used instead of the optical amplification tube 300.

According to this preferred embodiment of the invention,
The scintillator is covered along the length and at the entrance surface with a light-reflecting material such as silicon dioxide to prevent light from escaping (or entering) and to aid internal reflection within the scintillator. According to another preferred embodiment of the present invention, each scintillator element 179 is separated from an adjacent scintillator element by a thin film 171 of a highly radiopaque material such as gold or lead. Membrane 171 is preferably about 0.0102 cm (0.004 inch) to 0.0127 cm (0.005 inch) thick. The location of the membrane 171 between the scintillators 170 is shown in FIG.
Shown in 2.

As shown here, the circular active area of the collimation grid 90 is one of the detector arrays.
Wider than 10 areas. In this way, all X-ray pencil beams emitted from the respective apertures 140 of the collimation grid 90 are collected at the detector array 110. On the other hand, each independent X-ray beam 100 diverges and spreads like a flash beam.

Image Processing An important aspect of the present invention relates to the application of image processing systems to further reduce the required exposure dose. As a practical matter, the signal from the detector is usually not directly used for the "Z" or luminance input of the video monitor. Instead, the digitized intensity of each pixel is stored at a separate address in the "frame store buffer". Pixel addresses in the buffer are randomly accessible and numeric intensity values are treated mathematically. This feature has many image extension algorithm applications that can be applied,
Allows pixel allocation of data from independent parts of the detector array.

According to a preferred embodiment of the invention, the SBD
The X image consists of 500 columns and 500 rows, about 250,0
00 or more pixels (corresponding to 500 columns and 500 rows of the opening at the center of the collimation grid 90). For the purposes of the examples below, it is assumed that the scanning x-ray source is at a center P on a pixel at 100 columns and 100 rows of collimation grid 90 at a given moment. Further, for this embodiment, detector array 110 has nine segments 179.
Assume that each segment 179 is sized to block all X-ray radiation associated with a single pixel. Obviously, other array geometries may be used as detailed herein.

The digitized numbers from the individual segments of the detector array 110 are assigned the following pixel addresses. Segment 1-99 column, 99-row segment 2-99 column, 100-row segment 3-99 column, 101-row segment 4-100 column, 99-row segment P-100 column, 100-row segment 6-100 column, 101-row segment 7 -101th column, 99th row segment 8-101st column, 100th row segment 9-101th column, 101th row The same data allocation pattern is repeated by passing the scanning X-ray beam through all pixels.

In the displayed image, the value of each pixel is equal to the sum of the "n" parts. Where "n" is
The number of segments 179 of the array 110 (where n =
9).

When configured as shown here, the detector array 110 has a fixed working distance for optimum focus and is a conventional undivided (non-divided) detector array SBDX.
It has the effect of providing a plane of optimum focus not available in image systems.

The following parameters must be taken into account in the detector design. 1. The size and shape of the collimated beam from the x-ray source (anode target) 50. 2. Distance between X-ray source 50 and detector array 110; "SD" 3. 3. Distance between the x-ray source 50 and the center of the object of interest 80; "SO" 4. 4. desired resolution of object of interest 80, or pixel size In medical applications, the entire area of the array must be large enough to block X-rays from the collimation grid 90.

SBD for the preferred embodiment of the invention
In the X system, the distance between the X-ray source 50 and the exit 260 of the collimation grid 90 is about 2.271 cm.
(0.894 inches) (see FIGS. 3 and 5). Aperture 140 is 0.0381 cm (0.015 inch) x
It is a square of 0.0381 cm. The spot size of the electron beam 40 on the anode 50 is about 0.0254 in diameter.
cm (0.010 inch). Detector array 110
Is 36 inches from the anode 50. Thus, the beam width of the X-ray beam 100 is 2 *
ARCTAN ((spot diameter / 2) / ((aperture width /
2) + (spot diameter / 2)) * 2.271 cm (0.
894 inches), or 1.6 °. Anode 5
At a distance of 0 to 91.4 cm (36 inches), the irradiated X-ray beam diameter is 91.4 * TAN (1.6
°) cm. Therefore, the detector array 110 is
In the aspect of this preferred embodiment, about 2.54 cm.
Should be (1 inch). For example, the object to be photographed is 22.86 cm (9 inches) from the anode and the desired pixel size is 0.0508 cm for the object.
(0.020 inches), the distance from the X-ray source to the detector is 91.4 cm (36 inches), and the size of the optical detector array is 2.54 cm (1 inch) square The size of the illuminated pixel at the vessel plane 270 is simply (SD / SO) * pixel size at the object, or 0.2032 cm (0.080 inch). 2.5
Dividing 4 cm (1 inch) by 0.2032 cm (0.080 inch) will give the desired resolution using a squared detector with 12 to 13 segments on the surface. Obviously, many other geometries could be used depending on the circumstances in which the SBDX system is used.

Outside the surface SO (280 in FIG. 7D) of optimum resolution, 0.5 × SO, 2 × SO (281 in FIG. 7D)
At, the resolution deteriorates to half. This allows a reasonable depth of focus for many applications. In some applications, such as images of the human heart, degraded focus outside this depth range may be seen as an advantage. Blurring of details outside the region of interest increases the perception of detail within the region of interest.

Many methods can be used to obtain a usable image from the data obtained as described above.
As mentioned above, simple convolution can be used, but in this case the resolution is not optimized at all.
The two additional methods are preferred to obtain maximum resolution and sensitivity from the data obtained. These are called multi-image rotation and multi-output rotation. In both cases the following is assumed.

AP with an opening on the collimation grid 90
_{There are X} columns and AP _Y rows of openings. The intersection of each row and column is a "pixel". Those pixels outside the circular active area of the collimation grid 90 are treated as if there were no measured intensity in the image, eg as “dark”. Pixels that are not illuminated by the x-ray beam during scanning are also treated as if they were "dark", as if there was no measured intensity in the image.

Referring to FIG. 15, the sensor element 16 of the detector array is added to the pseudo circular sensor array 110.
There is a maximum in the DET _X column of 0 and a maximum in the DET _Y row of sensor element 160 of detector array 110.

ZRATIO is a real number between 0 and 1. If ZRATIO = 1, the focus is set on the sensor surface. If ZRATIO = 0.5, the focus is set at the midpoint between the X-ray source and the sensor surface. PIXE
LRATIO is the number of pixels for the physical distance between neighboring sensors in a column or row. For example, if the object plane 2
If the distance between the pixel centers at 80 is 0.01 cm,
The distance between the sensors on the detector plane 270 is 1.0 cm, PIXELRATIO = 10, FOCUS = ZRA
It becomes TIO x PIXELRATIO.

IMAGE is a DET _X × D containing intensity information for a particular scan and corresponding to a particular pixel.
A data column of dimensions ET _Y. PIXEL is a DET _X × D obtained by scanning all (or part of) the aperture.
AP _X , AP _Y , D including IMAGE data sequence of ET _Y
It is a four-dimensional array with dimensions of ET _X and DET _Y. PIX
The EL is refreshed after each scan according to one preferred embodiment of the invention. As the beam is scanned across the anode surface, it is effectively placed before the center of the selected aperture 140, "fired" and then reset. Thus, for each irradiation, an IMAGE sequence of data is obtained. While these images are assembled into some directly accessible displayable images, high resolution, high intensity is obtained by combining them. The preferred method of combining the first images is called the multi-image rotation method. In this multi-image turning method, AP _X × A that can be displayed on a CRT or similar display means
The OUTIMAGE sequence of the intensity in the P _Y dimension is OUTIMA
GE (y, x) is formed by assigning the value of the following equation (1).

[Equation 2]

A second currently preferred method of combining the APX × APYIMAGE data sequence into a useful image is called the multi-output rotation method. In this case, DETX ×
DETX x D together with the sensor array of the DETY sensor
There are ETY digitizers (or equivalent or multiple) and the same number of pixel summing circuits. The digitized value from each sensor is called SENSOR (j, i). The final OUTIMAGE column is calculated as follows. Output image sequence OUTIMAGE (y, x) [y
= 1 to AP _Y , x = 1 to AP _X ], D
ET _X x DET _Y source image SENSOR (j,
One pixel from each of i) is summed [j = 1 ...
DET _Y , i = 1 to DET _X ], the target image pixel O
UTIMAGE (y-j * FOCUS, x-i * FOC
US). Then, each element is DET _X * DET _Y
The OUTIMAGE sequence is standardized by being divided by.

A further refinement of these techniques is the FOCU
It can be obtained by performing linear interpolation based on the fractional part of the S element.

The advantage of the multi-image rotation method over the multi-output rotation method is that, in the former case, the plane of the optimum focus can be selected by software after the data is acquired.
The latter is not possible. However, the latter allows for faster operation when time is at the limit.

Reconstruction of 3D Images from SDBX Data The scanning beam digital radiography system described herein uses a set of continuous planar images used to form a tomographic 3D image of an object 80. Can be used to make The set of images can be analyzed to re-analyze the data set at different FOCUS values to produce a three-dimensional image consisting of a series of images at different depths. The natural FOCUS value used is n / DET _X or n / D
ET _Y , where n is an integer from 0 to DET _X or DET _Y , respectively. Typically, the focus value is analyzed to match the plane of interest within the object 80. For example, in the scanning beam digital radiography system shown in Table 1 below, the plane of focus is 22.8.
It is placed near the normal focal plane of 6 cm (9 inches) (optimal focal plane) and spaced approximately 2.54 cm (1 inch) apart.

The following equation (2) shows that a series of plane images are located in terms of the distance from the anode 50.

[Equation 3] Where Ft (FOCUS) is the distance from the anode to the particular focal plane of interest, and Fd is the detector to focal plane distance (anode to detector machine distance,
Is smaller than Ft) and λt is the distance between the centers of the adjacent openings of the collimation grid, and λd
Is the distance between the centers of adjacent detectors 160 in detector array 110.

When using the sub-sampling technique, the calculation remains the same, "not skid".
"pped)" data from the openings in the collimation grid are handled. However, λt is the same even if there are no intervening collimation grid holes.

Negative Feedback X-Ray Flux Control The BDX imaging system shown in FIG.
Negative feedback path 305 is used to control the 00 X-ray flux. Preferably, the negative feedback from the sensor array controls the x-ray flux so that the sensor array always sees approximately the same x-ray flux level. Thus, when soft tissue (relatively transparent to X-rays) is scanned, the X-ray flux falls, reducing the overall dose to the patient (object). By using negative feedback flux control, contrast and dynamic range are improved. According to this embodiment, the differential amplifier 310 has an adjustable reference level 320 that is user configurable. Negative feedback loop 3
Reference numeral 05 gives an X-ray flux negative feedback to the X-ray tube.

Time Domain Scan Mode The time domain imaging system can also be implemented using the principles described herein. In such a system, the time to reach the pre-measured x-ray flux from various pixels can be calculated and mapped. Negative feedback control can then be used to remove or reduce the x-ray flux from the apertures corresponding to pixels that have reached a predetermined flux level for the scan period being handled. In this case, the information collected is the time to flux level and the mapped and imaged information corresponds to time rather than intensity. The ability of such a system gives a much larger signal to noise ratio, improves contrast, and dramatically reduces the X-ray dose to the object under examination,
Allows to improve the dynamic range.

Multi-Energy X-Ray Imaging Mode According to one embodiment of the invention, two or more groups of X-ray beams 100 are directed towards one or more detector arrays. The first group of X-ray beams has a first characteristic X-ray energy spectrum. The second group of X-ray beams has a different second characteristic X-ray energy spectrum.
By comparing the measured transmission of the X-ray beams of the first group and the second group, the presence of an object in the object under examination can be detected. The basic concept of differential radiography has been known in the past and is described in US Pat. No. 5,185,77.
No. 3 ("Methods and Apparatus for Non-Destructive Selective Determination of Metals"), incorporated herein by reference.

The two groups of X-rays can be generated in a number of ways. In one method, a special anode comprising a first material or first thickness of material near the first group of openings and a second material or second thickness of material near the second group of openings. Perform manufacturing. Thus, the first
The aperture associated with the group emits X-rays having a first characteristic energy spectrum and the aperture associated with the second group emits X-rays having a second characteristic energy spectrum. Alternatively, K-filtering (K-edge filtering) can be used by placing a filter material (such as molybdenum) within the portion of the opening 140 that produces a similar effect. In this case, the first group of apertures comprises a first filter inserted therein and the second group of apertures comprises a second filter inserted therein. The second filter does not have to be a filter. As in the previous case, two groups of X-rays with different characteristic energy spectra are associated with the two groups of apertures.

If at least two groups of apertures are associated with different characteristic energy spectra, then a broadband X
It makes it possible to detect microcalcifications (breast cancer) and other abnormalities not normally visible on the line. For example, scanning the first group of apertures forming the first image,
Scan a second group of apertures forming an image,
Microcalcifications and other abnormalities can then be detected in real time using a low dose scan radiography system by enhancing their ratio by dividing the image. Similarly,
The multi-detector array device can be used with a first group of openings towards the first detector array and a second group of openings towards the second detector array.

Another embodiment of multi-energy imaging is described next. Since the magnitude of the electric pulse from the detected X-ray photon is proportional to the energy (kV) of the photon, 2
It is possible to separately count the pulses coming from the photons in the above energy bands. The pulses are separated by intensity, then counted and processed separately to produce two or more separate images. These images can be displayed as a ratio.

It is also possible to change the selected energy level in the fly in order to distinguish different density areas in the object. The advantage of this embodiment is that it is more flexible than the previous embodiments and does not require any special collimation grid, anode material or double detector.

While a number of preferred embodiments have been described above for various configurations of the present invention, the following description describes a preferred SBDX imaging system according to the present invention.

[Table 1]

Thus, this SBDX imaging system, which utilizes a segmented detector array, has high resolution and sensitivity, while at the same time providing a low irradiation dose to the object under examination. The system can also set the optimum focus at any point between the source 50 and the detector array 110, providing an effective working depth of field.

Beam Subscan Technique The following description relates to a particularly preferred embodiment of the present invention, which employs a technique of beam subsampling to reduce computational processing overhead.

The standard video image is 640 × 480.
Pixels, and updated at 30 Hz. This is about 12
It requires a pixel sample rate of MHz. X at this speed
Positioning the high voltage electron beam of the tube exactly behind 250,000 different apertures requires high precision and relatively high power consumption. Digitizing signals from a large array of X-ray detectors at a rate of 12 MHz is likewise effective and power consuming. Therefore, reduction of the pixel sample rate below 12 MHz without significant reduction in the spatial or temporal resolution of the SBDX is due to the cost of the initial unit, operating costs due to power consumption, and waste heat from the X-ray tube. Useful to reduce the cooling requirements for.

Therefore, a mechanism has been developed that provides substantially the same spatial and temporal resolution while reducing the pixel sample rate. This mechanism is referred to as subsampling and is obviously applicable to other configurations of SBDX, but is best implemented in the SBDX embodiments described herein. The effect of this embodiment is to reduce power consumption, a simpler circuit for deflecting the electron beam in the X-ray tube, lower production cost of the collimating grit 90, and calculation required to decompose the image of the object 80. , And other effects that will be apparent to those skilled in the art.

In this embodiment, the collimation grid 90 is made with a number of openings less than 500, preferably APX = APy = 166 (although other numbers could obviously be used). . The effect of this reduction from a computational point of view becomes clear below. However, the effect from a manufacturing point of view is a much simpler construction, about 1/9 the number of openings required for manufacturing. Manufacturing the grid with a higher deflection angle (ie, the angle that the openings make with respect to the front surface 260 of the collimation grid) causes the problem of having openings that intersect adjacent openings because the number of openings is reduced. Not easier.
This is because when a solid grid is manufactured, the adjacent apertures in the solid grid point towards different detector arrays and therefore require wider physical separation than the non-solid grid to prevent aperture crossing. , Especially useful.

Opening of collimation grid (plural)
Are arranged in a circle of APX row and APy column in the circle of maximum size. For computational purposes, this is outside the circle that does not contribute to information, ie it is always “dark” or X
It can be treated as a rectangle with dimensions of APX rows and APy columns with elements not illuminated by rays.

Sensor 160 of X-ray detector array 110
Shows DETx rows and DEs with maximum dimensions, as shown in FIG.
They are arranged in a circular array of Ty rows. The pixel sample rate can be reduced by illuminating less than all of the collimation grid apertures, or subsampling. Preferably, a collimation grid with no unilluminated openings is used. To make an image with this detector array, only every DETx th collimator hole in each row and the DETy th collimator hole in each row need be illuminated, and thus the image is each DETx It can be assembled from image tiles of pixels, with dimensions of pixels and DETy pixels. This is DETx x DE
Although it corresponds to a Ty subsampling ratio, subsampling does not correspond to a 1 × 1 sampling rate. Therefore, this sub-sampling ratio can be adjusted from 1 to DETx in the x direction (row) and from 1 to DETy in the y direction (column). According to this preferred embodiment, FIG.
As shown in, DETx = DETy = 12.

When a 12 × 12 detector is used and the subsampling ratio is 12, the image is much like the David Hockney light mosaic,
Created from multiple non-overlapping images that are substantially stuck together. Since the actual scintillator and detector are not quite perfect and do not respond exactly the same, the x-ray pencil beam is not perfectly uniform and the aperture of the collimation grid is ideal. It is not exactly the same as the actual area. And since a circular detector is used rather than a square detector, some degree of overlap is very favorable to allow averaging the detector non-linearity and noise.

If the sub-sampling ratio is less than the size of the detector in pixels (ie, less than 12 in the preferred embodiment), the images are assembled from overlapping "tiles" which are summed or Have to be averaged. If the sub-sampling ratio (s) are not multiple of the detector size (in pixels), or if the detector array is not a rectangular parallelepiped, then a different number of samples is added to each pixel and different divisions are performed. Is needed to average each pixel. Techniques for dealing with numbers less than the ideal environment are well known to those skilled in the art and will not be described here as it does not complicate the description.

In the calculation below, SS_XIs in the X direction (row)
Represents the sub-sampling dimension in_YIs in the Y direction
Shows the subsampling dimensions in (column). For example,
SS _X= SS_Y= 1, subsampling
Rather, the process in the other embodiments of the invention described above
Is exactly the same as. Similarly, if in this embodiment
Like SS_X= SS_YIf = 1, use pixel average
Return to "Light Mosaic". If SS_XAnd SS_YIs 3
And the active area of the circle is 500x500
The 166 x 166 aperture is scanned and scanned in the x direction.
1/3, 1/3 in the y direction, and the number of data obtained is
It is reduced by a factor of 9. If 1/9 opening is used for the whole time
If there is no need for these calculations, collimation
The grid does not have to be included.

Therefore, in order to create the image, the original code
1 / (SS in remation grid_X* SS_Y)
Only openings (500x500 openings) need to be used
That is, by the electron beam for X-ray irradiation
Needs to be illuminated. If the frame speed is constant,
That is, if it is kept at 30 Hz, drive the electron beam
The number of electron beam motions, such as the frequency response of the circuit
Is SS_X* SS_YIs reduced only. Electron beam moves
Total distance (and number of scan lines) is 1 / SS_YIs
And therefore the average through the target anode
Beam speed is 1 / SS_YIs reduced only. Reconstruct the statue
Pixel speed depends on the collimation aperture speed (aperture
Or the speed at which the light is emitted), as well as 1
/ (SS _X* SS_Y) Is reduced.

According to this method, the number of samples averaged to each display pixel is (DDTx / SS _X ) * (DDT
y / SS _Y ). With the maximum sampling rate (SS _X = DDTx and SS _Y = DDTy), only one digitizer sample is averaged for each display pixel (“light mosaic” mode). Sample averaging is important to smooth out non-uniformities in the beam, scintillator, detector and amplifier. Subsampling size (SS _X and SS _Y), for the conditions presented in order to ensure the quality of the acceptable image, must be set to an appropriate level. This can be set by the user for the fly depending on the user's preference for image quality and the conditions submitted by a particular set of environments.

The detector array 110 shown in FIG. 15 is preferably an array of 96 individual detector elements 160 approximately arranged in a circular area having a diameter of about 1 inch. Twelve detectors (D
DTx), 12 in the horizontal row and column at the center of the array.
There are detectors (DDTy). Scintillator crystals
Preferably, it is cut into square horizontal sections and is supported by an "egg crate" construction consisting of a 0.005 inch thick strip of stainless steel. Circle 40 of FIG. 15 with all scintillator crystals (hatched) located therein.
0 is preferably about 0.800 inches in diameter.

The scintillation crystal length in the detector array 110 is preferably 0.10 cm,
The input surface in front of it is preferably 0.135 cm x 0.1
It is 35 cm. The scintillation crystal is preferably YSO, LSO or BGO, but as mentioned above, other materials can be used. BGO has an appropriately reduced decay time (50%) for its light output in this application.
nS) must be heated to about 100 ° C.
Therefore, the resistance heating element may be provided.

FIG. 17 illustrates a detector assembly 402 according to the preferred embodiment of the present invention. X-ray is from above
It passes through the X-ray window 404 and enters the lead shield 406. X-ray collimation grid 90 while attenuating scattered X-rays
The x-ray window 404 is preferably circular and has a diameter of 1.91 cm to allow it to exit the apertures and strike the detector array 110. The light shield 408 is
It is provided to shield the detector from stray light. The light shield 408 can be manufactured from a sheet of aluminum or beryllium selected to attenuate light without substantially attenuating x-rays.

Detector array 110 is located near a suitable heating element 410 for use with a BGO scintillator. The heating element 410 may be a resistive heating element designed to maintain the detector array 110 at an operating temperature of about 100 ° C. A photon emerging from the fiber optic imaging taper 412 from the bottom 414 of the detector array 110 is converted into a 96-channel photomultiplier tube (PMT) 416.
Turn to. The detector assembly 402 is enclosed within a light tight outer housing 418 to prevent stray light from generating noise. 3 shoulder screws 42
Zero and three centering screws 422 are provided for planar and linear arrays, as is well known to those skilled in the art. Rotational arrangement includes outer housing 418 with PMT mount 426
Achieved by rotating with respect to. The Fiber Optic Imaging Taper 412 is commercially available from Collimated Halls, Inc., Campbell, Calif., With a 2.03 cm diameter circular input aperture and a 3.38 c diameter.
m circular output apertures. Taper 412 is PMT
Fit the pitch dimension (0.06 inch) of each scintillator crystal to the 416 dimension (0.10 inch). That is, the taper has a magnification of 1.667 times. The highly viscous optically coupled liquid (type 200) commercially available from Dow Corning has a refractive index approximately equal to that of glass, from scintillator crystal 160 to taper 412,
In order to maximize the efficiency of light transfer from the taper 412 to the PMT input surface 424, it is used as the optical coupling medium on two surfaces of the taper.

The photomultiplier tube 416 is a 96 channel tube (one channel corresponds to each scintillation crystal) commercially available as Model XP1724A from Phillips. The photomultiplier tube 416 comprises a fiber optic faceplate so that the spatial arrangement of the scintillation array is accurately performed on the PMT photocathode located at the PMT on the other side of the optic faceplate. X-ray photons striking one scintillation 160 produce a light pulse that is coupled to the PMT photocathode. This produces a corresponding electron pulse on the photocathode, which pulse is
1 channel of MT diode structure, 1,000,000
It is amplified up to 0 times.

This PMT output pulse is coupled to the input of an amplifier in the 30 MHz band, the output of the amplifier is in the range of 0.5 to 5.0 volts, so that the pulse is differentiated,
It is about 30 nsec long. This allows DC restoration (r r to keep the baseline reference voltage as the pulse rate changes.
Estorer) eliminates the need for circuitry.

The output of the amplifier is input to the comparator, and the comparator outputs an output pulse of a constant size regardless of the size of the input. The reference voltage for the comparator is set to a value just above the noise output level so that the amplifier is not triggered at the noise output level. The chain of amplifiers is repeated 96 times, once for each scintillation crystal in the detector array. The output pulse of the comparator becomes the new data for the data acquisition and image reconstruction system. As tests have shown, the prototype system can randomly count X-ray photons up to a rate of about 10 MHz.

While the embodiments and uses of the invention have been described above, those skilled in the art will appreciate that many more variations than those set forth above may be made.
Obviously, it is possible without departing from the inventive concept. Therefore, the present invention is not limited within the spirit of the claims.

[Brief description of drawings]

FIG. 1 is a diagram showing basic elements of a low-dose scanning beam digital radiography system.

FIG. 2 shows the distribution of X-rays from an SBDX system without a collimation grid.

FIG. 3 is an enlarged view of an X-ray tube grid and anode for a low-dose scanning beam digital X-ray imaging system.

FIG. 3A is a partial sectional view of a collimation grid effective for the device of the present invention.

FIG. 3B is a partial sectional view of a collimation grid effective for the device of the present invention.

FIG. 3C is a partial sectional view of a collimation grid effective for the device of the present invention.

FIG. 4 is a diagram of an X-ray tube for a low-dose scanning beam digital X-ray imaging system.

FIG. 5 is a sectional view showing the structure of an X-ray tube for a low-dose scanning beam digital X-ray imaging system.

FIG. 6 is a diagram of a stereoscopic scanning beam type digital X-ray imaging system.

FIG. 7A is a diagram of an X-ray source with an aperture cooperating with a non-segmented simple detector.

FIG. 7B is an X-ray view from one aperture of an apertured X-ray source cooperating with a split detector array.

FIG. 7C is a diagram of X-rays from multiple apertures of an apertured X-ray source cooperating with a split detector array.

FIG. 7D is a view of X-rays from two apertures of an X-ray collimation grid cooperating with the inspected object and the split detector array.

FIG. 8 is a view of the exposed surface of a 5 × 5 detector array for a low dose scanning beam digital radiography system.

FIG. 9: Diagram of a 5 × 5 detector array for a low-dose scanning beam digital radiography system.

FIG. 9A shows a scintillator element according to one preferred embodiment of the present invention.

FIG. 10 is a diagram of detector elements for a low dose scanning beam digital radiography system.

FIG. 11 shows an array of pencil-type detector elements for a non-planar detector array.

FIG. 12 is a diagram of a 3 × 3 detector array for a low dose scanning beam digital radiography system.

FIG. 13 shows the basic elements of a low dose scanning beam digital radiography system utilizing negative feedback to control the x-ray flux.

FIG. 14 is a perspective view showing a grid seal device.

FIG. 15 is a layout of a 96-element detector array according to a preferred embodiment of the present invention.

FIG. 16 is a diagram showing the interaction between a collimation grid and a detector array.

FIG. 17 shows a preferred embodiment of the detector device.

[Explanation of symbols]

10 X-ray tube, 20 beam, 30 beam,
161 electron beam source, 50 target anode (X-ray source), 80 object to be irradiated,
90 Collimator element (aiming element), 100 X
Ray beam, 110 X-ray detector array, 140 aperture, 164 beam, 270 detector plane,
280 focal plane.

─────────────────────────────────────────────────── ─── Continuation of front page (51) Int.Cl. ⁷ Identification code FI theme code (reference) G01T 1/00 G01T 1/00 B G03B 42/02 G03B 42/02 B G06T 1/00 400 G06T 1/00 400B G21K 1/02 G21K 1/02 G 5/02 5/02 X H01J 35/24 H01J 35/24 H04N 5/225 H04N 5/225 C // A61B 6/00 300 A61B 6/00 300C G01T 1/20 G01T 1/20 G (72) Inventor John William Willent USA 95003 Victoria Lane, 139th, California Lane 139 (72) Inventor Jack Wilson Mooreman United States 95032 California, Los Gates, Pinta Court 136th ( 72) Inventor Brian Su Recorn USA 94087 California, Sunnyvale, The Dallas Avenue 898 (72) Inventor Peter Joseph Fiekowski USA 94024 California, Los Altos, Springer Road 952 F term (reference) 2G088 EE01 FF02 GG13 GG14 GG18 GG19 GG20 JJ01 JJ05 JJ12 JJ22 KK32 KK33 MM04 2H013 AC01 AC06 AC08 4C093 AA30 CA02 CA03 CA34 EA03 EA04 EA07 EA11 EA12 EB02 EB12 CB16 AC08 AC21 AC08 AC21 AC01 AC06 AC08 4C093

Claims (25)

[Claims] 1. An electron beam is traversed by an AP across an X-ray source._x
× AP_yFollow the step pattern of stepwise,
DET_x× DET_yImage before crossing the detector array
Create an X-ray beam that penetrates the object to be converted, X-rays detected by each detector constituting the detector array
Measure the intensity and detect each step of the electron beam
Contains information that is an indicator of the intensity measured by the vessel
DET_x× DET_ySize data array (IMAG
Store the equivalent of E) in memory, IMAGE data for each step of electron beam
Including the DET _x× DET_y× AP_x× AP_yof
Size equivalent of 4D array (PIXEL) in memory
Store and FOCUS is a number between 0.1 and 10.0, and
formula [Equation 1] With a value of IP_x* IP_y-Dimensional array O of size
Output to form the equivalent of UTIMAGE (x, y)
Create strength, An X-ray imaging method for outputting the output value for analysis. 2. The method of claim 1, wherein the x-ray beam is transmitted through a plurality of x-ray transparent areas of the collimator before being transmitted through the object to be imaged. 3. The method of claim 2, wherein the plurality of x-ray transparent regions are cylindrically shaped and each of their axes intersects the detector array substantially at the center of the array. 4. The FOCUS is a factor of the proportional distance between the plane of the object to be imaged and the x-ray source and the ratio of the distance between the centers of adjacent detectors and the distance between electron beam steps. The method according to claim 1, corresponding to. 5. An electron beam is traversed across the x-ray source to the AP.
Proceed stepwise with the step pattern of _x , AP _y ,
Creating an X-ray beam that penetrates the object to be imaged before intersecting the DET _x × DET _y detector array and measuring the X-ray intensity detected by each detector in the detector array, the X-ray intensity value from each of the detectors DET _x × DET _y digitizing, DET _x × containing the digitized information indicative of the measured intensity by each detector for each step of the electron beam Two-dimensional matrix of size DET _y (SEN
SOR) equivalent is stored in memory, j is varied from 1 to DET _y and i is varied from 1 to D
Change to ET _x and sum SENSOR (j, i),
Target pixel array DEST (y-j * FOCUS, x-i *) containing FOCUS that is a number between 0.1 and 10.0
FOCUS), obtain an output value, and output the output value for analysis. 6. The method of claim 5, wherein the output values are normalized by dividing each value by DET _x × DET _y . 7. The method of claim 5, wherein the x-ray beam is transmitted through the plurality of x-ray transparent areas of the collimator before being transmitted through the object to be imaged. 8. The method of claim 7, wherein the plurality of x-ray transparent regions are cylindrically shaped and each of their axes intersects the detector array substantially at the center of the array. 9. The FOCUS of the proportional distance between the plane of the object to be imaged and the x-ray source and the ratio of the distance between the centers of adjacent detectors and the distance between the steps of the electron beam. The method according to claim 5, which corresponds to a factor. 10. A plurality of X-ray beams emitted from different points on an X-ray source are generated, and each of the generated X-ray beams transmits an X-ray transmitted through an object.
The rays received by the detector array, and from each x-ray beam, the x received by each detector.
Determining information indicative of the intensity of the line, accepted by at least two detectors intersected by a line defined by different points on the x-ray source and the pixel for which the imaging information is sought An imaging method for obtaining imaging information about a pixel in an object by combining the information that is used as an index of the intensity of the X-ray. 11. The method comprising: producing the x-ray beam, scanning an electron beam across the x-ray target, and producing the x-ray beam at each of a plurality of points on the x-ray target. 10. The method according to 10. 12. The method of claim 10, wherein the pixels are all coplanar. 13. Determining information indicative of the intensity of X-rays received by each detector from each X-ray beam comprises counting the number of photons of X-rays received by each detector. 11. The method of claim 10, wherein 14. The X-ray beam is transmitted through an X-ray of a collimator before the X-ray beam is transmitted through an object to be imaged.
The method of claim 10, wherein the array of AP _x × AP _{y in} the linearly transparent region is transmitted. 15. Imaging information is provided to a monitor for producing an image having an IP _x , IP _y pixel array, wherein the ratio of the total number of X-ray transparent regions to the total number of pixels is greater than 1. 15. The method of claim 14, which is small. 16. Imaging information is provided to a monitor for producing an image having an IP _x × IP _y pixel array, wherein the ratio of the total number of X-ray transparent areas to the total number of pixels is greater than 1. 15. The method of claim 14, which is large. 17. The first of AP _x × AP _y of the X-ray transmissive region of the collimator before the X-ray beam is transmitted through the object to be imaged and detected by the first detector array. A second AP _x × AP _y of the X-ray transparent region of the collimator is transmitted through the array and before the x-ray beam is transmitted through the object to be imaged and detected by the second detector array. The array is transparent and is received by each of the detectors in the first array crossed by one of the lines defined by each x-ray transparent region in the first array and the pixel for which imaging information is sought. The first imaging information about the pixels on one surface in the object is obtained by combining the information that is an index of the number of photons of the obtained X-rays, and each X-ray transparency in the second array is obtained. Area and pixel for which imaging information is sought By combining information indicative of the number of photons of X-rays received by each of the detectors in the second array crossed by one of the lines thus defined, on one face in the object 11. The method of claim 10, wherein the first and second imaging information for outputting the second imaging information about the pixels of the respective pixels and performing the processing for providing to the left and right eyes of the user, respectively. . 18. AP _{x in} the X-ray transparent region of the collimator
The first array of × AP _y each includes one longitudinal axis, each of the axes being DET substantially at the center of the first detector array.
_A second array of AP _x , AP _y in the X-ray transmissive region of the collimator includes one vertical axis each, and each of the axes is substantially parallel to the first detector array of _x x DET _y . so that the cross the second detector array of DET _x × DET _y at the center of the second detector array, the method of claim 17. 19. The method of claim 17, wherein the x-ray transparent regions of the first and second x-ray transparent region arrays are each adapted to diverge from a common x-ray transparent region. 20. An X-ray source, an X-ray detector array, a monitor and a scan generator, the X-ray source comprising an electron beam generator and an X-ray source.
Control including two anodes, the anode including one target layer that emits photons of X-rays when excited by the electron beam, the scan generator controlling the electron beam produced by the electron beam generator. A circuit, the x-ray detector array including a plurality of detector elements,
The detector elements are arranged in an array forming a normally flat detector surface having one center, the anode and the x-ray detector array being arranged such that the electron beam is at a first point. When the target layer is excited, photons of the emitted X-rays strike the detector elements of the X-ray detector array and are emitted when the electron beam excites the target layer at a second point. X-ray photons are arranged to impinge on the detector elements of the X-ray detector array, the scan beam generator comprising:
Positioning circuitry for continuously positioning the electron beam at selected points in a predetermined pattern, each of the detector elements measuring the intensity of photons of the X-rays impinging on the detector element and A means for generating an electrical signal indicative of the intensity, wherein the electrical signal from each of the detector elements is output in digital form from the X-ray detector array to provide a group of digital electrical outputs. An X-ray imaging system that includes processing circuitry that is output for positioning an electron beam at successive points on the target layer and wherein the monitor produces a visible image based on the electrical output in digital form. . 21. A collimator disposed between the target layer and the detector array proximate to the target layer, the collimator including a plurality of x-ray transparent regions. Item 21. The X-ray imaging system according to Item 20. 22. The x-ray transmissive regions each include one longitudinal axis, each of the axes being substantially centered across the detector array. Two
The X-ray imaging system according to 1. 23. An X-ray source, an X-ray detector array, a monitor and a scan generator, the X-ray source including an electron beam generator and an electron beam generator. Electrons comprising one anode and one collimator, the anode including one target layer that emits photons of X-rays when excited by the electron beam, and the scan generator includes electrons generated by the electron beam generator. A control circuit for controlling a beam, the collimator being disposed between the anode and the X-ray detector array proximate to the anode, the X-ray detector array including a plurality of detector elements;
The detector elements are arranged in an array forming a normally flat detector surface with one center, the collimator being composed of one X-ray absorbing material, The absorptive material includes a plurality of apertures, the axis of each of the apertures substantially across the detector surface at the center of the X-ray detector array and across the target layer, the anode and the anode. Collimator and above X
A line detector array, the emitted X-rays when the electron beam excites the target layer at a first point at the intersection of the first of the axes of the aperture and the target layer. Of photons pass through the first aperture and strike the detector element of the X-ray detector array, and the electron beam is directed at the intersection of the second one of the axes of the aperture and the target layer. Is arranged such that, when exciting the target layer at a second point, the emitted photons of X-rays pass through the second aperture and strike the detector elements of the X-ray detector array; The scan beam generator includes a positioning circuit for continuously positioning the electron beam at the intersection of the axis and the target layer in a predetermined pattern, each of the detector elements being the detector. Hit the element A conversion means for effecting energy conversion from the photons of the x-rays into electrical signals, the electrical signals from each of the detector elements being output in digital form from the x-ray detector array, a group of digital forms Electrical power is output to position an electron beam at successive intersections of the axis and the target layer, and the monitor includes processing circuitry for generating a visible image based on the electrical output in digital form. Delivering X-ray imaging system. 24. The x-ray imaging system of claim 23, wherein the detector elements are arranged in an array that approximates a circular shape. 25. A scanning beam X-ray tube for use with anode voltages up to about 120 kV, comprising an anode disposed on an anode support, the anode support being a relatively X-ray transparent material. The anode includes a first layer of niobium provided on the anode support and a second layer of tantalum provided on the first layer, and the anode is made of niobium. One layer has a relatively uniform thickness of about 1 μm and the second layer of tantalum is about 5 μm.
Scanning beam x-ray tube having a relatively uniform thickness of.