CN1041237C

CN1041237C - X-ray detector for a low dosage scanning beam digital X-ray imaging system

Info

Publication number: CN1041237C
Application number: CN94192143A
Authority: CN
Inventors: J·W·韦伦特; J·W·穆尔曼; B·施基里科恩; P·J·菲考斯基; V·韦伦特
Original assignee: NexRay Inc
Current assignee: NexRay Inc
Priority date: 1993-04-05
Filing date: 1994-04-05
Publication date: 1998-12-16
Anticipated expiration: 2014-04-05
Also published as: EP0693225A1; JP4229859B2; WO1994023458A3; JPH08508431A; CN1127553A; IL109143A; IL109143A0; WO1994023458A2; AU6700694A; JP3569526B2; JP2004261605A; EP0693225A4; JP2003265457A

Abstract

A scanning beam digital x-ray imaging system according to the present invention includes an x-ray tube (10) having an electron beam source (161) and a target anode (50). Circuitry is provided for focussing the beam (164) and scanning the beam (20, 30) across the target anode (50) in a predetermined pattern such as a serpentine scan pattern. A collimating element (90), preferably in the form of a perforated grid containing an array of apertures, is interposed between the x-ray source (50) and an object to be x-rayed (80). The apertures (140) are oriented to form x-ray beams (100) converging at a detector array (110) on a plane (270) located a selected distance from the collimating element (90). That distance is selected to allow placing the object to be x-rayed in between the collimating element (90) and the detector array (110). A segmented x-ray detector array (110) containing a rectangular matrix of detector elements (170) is located in the detection plane (270). A focal plane (280) is created which gives optimal resolution at a particular, selectable distance from the x-ray source (50).

Description

The X-photo-detector technical field of the present invention that is used for low dosage scanning beam digital X-photoimaging system

The present invention relates to X-photodynamic diagnosis imaging device.More particularly, the present invention relates to a kind of real time scan bundle X-light digital imaging system, this system is by being used in combination the emission that porous alignment grating and subregion X-photo detector array have improved resolution and reduced X-ray-ray.

Background technology

Along with the progress of treatment technology, medical approaches is growing for the needs of real-time X-photoimaging.For example, the many electronics-physiological diagnosis method of cardiology, the capillary diagnostic method, urology diagnostic method and orthopaedic surgery method all rely on real-time X-photoimaging.

Regrettably, present clinical employed real-time X-light device all has the x-ray radiation of high dose to patient and medical worker on the scene.FDA (FDA) has reported about the atomic disease that causes serious patient and the excessive occupational radiation strong evidence for doctor's influence.(radiation and healthy proceedings, XXVI volume, 8 fascicles, in August, 1992).

The real-time X-photoimaging of many kinds system has been arranged in the prior art.Comprise the system that adopts the fluoroscopy method in these systems, in this system, with X-rayed-object, in the object comparatively speaking not the shade that material produced of saturating X-light just on video screen, show with respect to the opposite side of this object with the X-light source.It is reported that scanning the X-light pipe as far back as the beginning of the fifties at least just is used in combination with the fluorescence detection technology.Can be about this point referring to the article of MOON, " utilizing scanning X-light pipe to amplify and the enhancing fluorescent image ", science magazine, October nineteen fifty, 389-395 page or leaf.

Scanning beam X-light digital imaging system is also known in the art.In these systems, adopted the X-light pipe to produce the X-optical radiation.In the X-light pipe, produce a branch of electron beam and it is focused on the point on the bigger comparatively speaking anode (transmission target) of X-light pipe, from then on the some place causes the X-optical radiation.Thereby electron beam is formed grating scanning pattern with electromagnetism or electrostatic methods deflection on whole anode.A little X-' photo-detector is placed on apart from X-light pipe anode a distance.The X-light that this detector will be radiated it converts the electric signal that is directly proportional with the X-luminous flux that is detected to.When being placed on an object between X-light pipe and the detector, X-light is decayed and scattering by object, and the degree of decay and scattering is proportional to the X-optical density on this object.When the X-light pipe was scan pattern, the signal of detector was modulated by the X-optical density that is proportional on the object.

The example of scanning beam digital X-photosystem is included in those systems that put down in writing in the following United States Patent (USP) of authorizing Albert in the prior art: US-3,949,229, US-4,032,787, US-4,057,745, US-4,144,457, US-4,149,076, US-4,196,351, US-4,259,582, US-4,259,583, US-4,288,697, US-4,321,473, US-4,323,779, US-4,465,540, US-4,519,092, US-4,730,350.

In the exemplary embodiments of the scanning beam digital X-of prior art photosystem, the output signal of detector is imported into Z-axle (brightness) input end of video monitor.The brightness of this Signal Regulation screen.Giving the X of video monitor and Y input signal all is that same signal from the X-light signal deflection that influences the X-light pipe draws.So the brightness of a point is inversely proportional to the light emitted from X-on the screen,, arrive the degree of absorption of the X-light of detector through object.

Medical science X-photosystem is to work under the possible minimum x-ray dose that adapts with the desired resolution of diagnostic method.So radiation dose and resolution are all limited by signal to noise ratio (S/N ratio).

Employed technical term " low dosage " refers at equipment patients gateway place among the application, the X-light width of cloth amount of penetrating during system works less than or approximate 2.0R/min (roentgen/minute).

The time of X-light photon and space distribution are that Poisson (Poisson) distributes, and have relevant randomness inevitably.This randomness can be expressed as the standard deviation of average flux, and equals its root-mean-square value.So the signal to noise ratio (S/N ratio) of X-light image equals the root-mean-square value of average flux divided by average flux under these conditions, promptly when average flux be 100 photons, noise is+during/-10 photons, signal to noise ratio (S/N ratio) equals 10.

Therefore, the spatial resolution of the X-light image that is produced by scanning X-photoimaging system and the size that signal to noise ratio (S/N ratio) depends on the sensitive area of detector to a great extent.If increase the aperture area of detector, then can detect more distribution ray, thereby improve sensitivity effectively, and improve signal to noise ratio (S/N ratio).But because elemental area (what measured by the object plane of imaging) change is big, bigger detector aperture makes accessible spatial resolution reduce simultaneously.Because most of object of being made a video recording in medical applications (for example, various tissues in the human body) all has certain distance with the X-light source, so this is inevitable.Thereby in the prior art, the pore size of detector must be through selecting taking into account resolution and sensitivity, but can not make resolution and sensitivity all reach best simultaneously.

In medical imaging was used, patient's absorption dosage, frame rate (number of times that number of times that the object per second is scanned and image upgrade) and the resolution of image all were key parameters.High X-luminous flux can easily obtain high resolution and high frame rate, but also patient and medical worker on the scene has been produced unacceptable high X-optical radiation dosage.Equally, low radiation dose then must be with the cost that is reduced to of image quality and renewal rate.The medical image system of a success must provide low dosage, high resolving power and gratifying turnover rate simultaneously, wants per second 15 images at least.So, system in aforesaid prior art the scanning beam digital X-photoimaging system just can't be applicable to many medical diagnostic method, because the exposure time that needs in these methods is longer, and owing to what face is real patient, so the X-light dosage that patient takes in must be remained on minimum value.

So an object of the present invention is to provide a kind of scanning beam digital X-photoimaging system that can be used in for patient's medical diagnostic method.

Another object of the present invention provides and a kind ofly can generate high-resolution picture under enough frame rate, and has reduced the scanning beam digital X-photoimaging system to the irradiation of the object checked with X-light to greatest extent.

Another purpose of the present invention provides a kind of in the X-luminous flux that has kept reducing, and has the scanning beam digital X-photoimaging system of the resolution of raising in a distance, distance X-light source plane.

By with reference to the accompanying drawings with following for description of the invention, it is fully aware of that these purposes of the present invention and other purpose and advantage will become for a person skilled in the art.

Of the present invention open

Scanning beam digital X-photoimaging system of the present invention (" SBDX ") comprises an X-light pipe with an electron beam source and a plate target.For make electron beam focus on, directed and with predetermined pattern antianode target scan, disposed electronic structure.For example, predetermined figure can be a grating scanning pattern, spirally or ' S ' shape figure, and the spirality figure, random figure, the Gaussian distribution figure of center on a predetermined point of plate target, perhaps other are suitable for the figure of the work of doing.

At the X-light source and between collimating element, preferably a grating can be set with the object of X-photoirradiation.For example, collimating element can be made of a metal dish, the about 25.4cm of its diameter (10in), and the array in the hole of 500 * 500 ranks is arranged at the center of collimating element.Collimating element preferably directly is placed on the surface of emission front of X-light pipe.The collimating element of other structures also can use.Each hole on the collimating element all has such structure in a preferred embodiment of the invention, promptly is oriented to (or point to) and is positioned at sensing point on the plane at collimating element selected distance place.This distance is can be placed between collimating element and the sensing point through the object of selecting to use the X-x ray fluoroscopy x.The function of collimating element is to form thin X-light beam, and these X-light beams are lined up array as pixel, and some directed towards detector from the X-light pipe anode all.

A subregion detector array that is made of the detector cells of an array (preferably a planar array, such as DETx * DETy rectangle or square, perhaps, preferable is general rounded array) is centered in this sensing point.Detector array preferably is made of the X-photo-detector of one group of fine and close encapsulation.According to the present invention, such array can design by this way, is provided with and uses, promptly under the prerequisite of not losing resolution, obtain high sensitivity, thus the X-photosystem is lacked than X-photosystem of the prior art at least at radiation dose under the situation of an order of magnitude but have can with the analogy of X-photosystem or the The better resolution of prior art.This characteristic of the present invention has great importance in medical science and other field.Radiant quantity for patient and personnel on the scene in present diagnostic method will be reduced.The method that can not use owing to radiation risk might be applied now.

The output valve of detector array is the intensity level of each unit when the X-light beam passes hole on the grating in the detector array on each aspect.Because each hole is arranged on the different point in space, so for each hole that has X-light to pass through, detector array has different output with respect to checked object and detector array.Detector array output can convert image in many ways to.A method is that a simple convolution transform is carried out in output to array, soon is scanned the intensity level adduction of the array element in hole, normalization then corresponding to each.This output array just can be used for driving a video display or other displays afterwards.Preferable is how visual convolution transform and many output convolution transform methods, and as describing hereinafter, it can provide the image output of enhancing.

The SBDX imaging system can also be carried out three-dimensional imaging, and alignment grating wherein will be used two groups of holes.In this case, one group of hole constitutes points to first sensing point that is provided with the first subregion detector, and second group of hole is configured points to second sensing point that is provided with the second subregion detector.By forming two images, and use conventional stereo display method, can constitute a stereoscopic image by two subregion detectors.

The SBDX imaging system can also make those work as X-light photon energy, and not have the imaging of material of different X-transmitances simultaneously outstanding.So as the early stage form of breast cancer, promptly Microcalcification can be made a video recording.Make it to launch the X-light beam that two or more sets have different X-energy spectrum by structure grating and/or anode, and make each organize all directed towards detector arrays (also can use a plurality of detector arrays), then just can be transformed into an image by the radioparent difference of the object of multiple X-light photon energy irradiation, thus only outstanding behaviours have those materials of different X-light transmissions in the inspected object.Such imaging system can be preferred for surveying calcium, and for example it is strong instruments for the early stage inspection of breast cancer and other tissue diseases.

Utilization by intercepting and capturing whole collimation X-light beams the subregion array and handle array output and obtain image maximum sensitivity is provided, but not have to sacrifice the resolution that when the little surface area detector of use, obtains.The not subregion detector array onesize with the subregion detector array may have same sensitivity, but resolution is lower.

In addition, can adopt the data of double sampling technical finesse from detector array, this makes and can lower the complicacy of system and required processing speed and energy consumption under the situation of the practically identical image quality of maintenance.

The described system of the application can be 08/008,455 (CAM-003) in sequence number, " having the conduit to the light activated optical sensor locating device of X-" of being put down in writing in the U.S. Patent application that on January 25th, 1993 filed an application is used in combination.U.S. Patent application 08/008,455 for reference only in this application, it is also had by the application's assignee.

Brief description of the drawings

Fig. 1 represents the building block of low dosage scanning beam digital X-photoimaging system.

The distribution of X-light in SBDX system when Fig. 2 represents not have alignment grating.

Fig. 3 is used for the grating of low dosage scanning beam digital X-photoimaging system and the enlarged drawing of X-light pipe anode.

Fig. 3 A, 3B and 3C are the partial cutaway figure that is used for the alignment grating of apparatus of the present invention.

Fig. 4 represents to be used for the X-light pipe of low dosage scanning beam digital X-photoimaging system.

Fig. 5 represents to be used for the sectional view of the X-light-pipe structure of low dosage scanning beam digital X-photoimaging system.

Fig. 6 represents stereoscanning bundle numeral X-photoimaging system.

Fig. 7 A represents the X-light source with holes that is used with a simple not subregion detector.

Fig. 7 B represents the X-light that sends from a hole of the X-light source with holes that is used with a subregion detector array.

Fig. 7 C represents the X-light that sends from a plurality of holes of the X-light source with holes that is used with a simple not subregion detector.

Fig. 7 D represents the X-light that sends from two holes of the X-ray alignment grating that is used with checked object and subregion detector array.

Fig. 8 represents to be used for the irradiating surface of 5 * 5 detector arrays of low dosage scanning beam digital X-photoimaging system.

Fig. 9 represents to be used for 5 * 5 detector arrays of low dosage scanning beam digital X-photoimaging system.

Fig. 9 A represents scintillator cells according to a preferred embodiment of the present invention.

Figure 10 represents to be used for the detector cells of low dosage scanning beam digital X-photoimaging system.

Figure 11 represents to be used for the array that the pencil-type detector cells of an on-plane surface detector array constitutes.

Figure 12 represents to be used for 3 * 3 detector arrays of low dosage scanning beam digital X-photoimaging system.

Figure 13 has represented to adopt the essential part of negative feedback with the low dosage scanning beam digital X-photoimaging system of control X-luminous flux.

Figure 14 is the skeleton view of grating packoff.

Figure 15 is illustrated in the outward appearance of employed 96 single-element detector arrays in the most preferred embodiment of the present invention.

Figure 16 represents cooperating of alignment grating and detector array.

Figure 17 represents a most preferred embodiment of detector assembly structure.

Embodiments of the present invention

It only is illustrative that those skilled in the art should understand the following description of this invention, rather than restrictive.Other embodiment of the present invention can oneself be provided by these technician.System's overview

Fig. 1 has represented the scanning beam digital X-photoimaging system in the most preferred embodiment of the present invention.Wherein used a scanning X-light pipe 10 as the X-light source.As in the prior art, adopted one approximately-100kV gives the X-light pipe 10 power supplies to the power supply of-120kV.The power supply of 100kV can produce the X-energy spectrum up to 100keV.As used in this application, the X-light of 100kV just refers to this power spectrum.The same with prior art, X-light pipe 10 comprises a deflection coil 20 that is subjected to scanning generator 30 controls.The electron beam 40 that produces in X-light pipe 10 is by with the ground connection anode 50 in a kind of predetermined graph scanning X-light pipe 10.For example, predetermined figure can be a grating scanning pattern, spirally or ' S ' shape figure, and the spirality figure, random figure, the Gaussian distribution figure of center on a predetermined point of plate target, perhaps other are suitable for the figure of the work of doing.Preferably spirally or ' S ' shape figure, it can exempt the needs of grating scanning pattern for " flyback " (fly back).

When electron beam 40 during in point 60 place impinge anode 50, cluster X-ray 70 is launched out, and flies out from X-light pipe 10, and directive is with the object 80 of X-optical test.For the performance that makes system reaches best, must produce a cone-beam shape X-light photon, its degree of dispersing will cover detector array 110 just.This preferably realizes by between anode that scans the X-light pipe and detector an alignment grating being set.So between object 80 and X-light pipe 10, placed alignment grating 90.Alignment grating 90 is designed to only allow those X-light 100 of directive detector 110 by it.When system works, alignment grating 90 does not move with respect to detector array 100.Therefore, when electron beam 40 scans on anode 50,, can only there be a branch of X-light to be transmitted into detector array 110 from anode in any given moment.

Fig. 2 is illustrated in the distribution of the X-light when not having alignment grating.

The output of detector array 110 is through handling, can on the monitor 120 as monitor 120 on x, the brightness value in y orientation shows, the x on itself and the anode 50, the y orientation is corresponding.This can drive the x of electron beam 40 by using same scanning generator, and realize the location of electron beam in the mobile and video monitor 120 on the y direction.Perhaps adopt image processing technique on a demonstration that is fit to or photographic medium, to produce Computer Generated Image.

The disclosed system of the present invention of the application is a low dosage system, accept to diagnose the place, entry position to measure the patient of equipment, to patient's radiation dose scope 15 frame/seconds visual renewal rate, about 0.15R/ assign to visual turnover rate 30 frame/seconds, about 0.33R/ branch.When the visual turnover rate of 30 frames/second, system is approximately the 0.50R/ branch to the radiation dose of whole human body.So, use the present invention at the equipment entrance place radiation dose scope for patient assign to the 2.00R/ branch as 0.15R/.The X-light pipe

Fig. 3 represents the zoomed-in view of grating and anode construction.Preferably with having good vacuum characteristic and can Nai Gaore and the target layer made of the material of the bombardment of electron pair beryllium anode substrate 130, the material that aluminium or other can see through X-light comparatively speaking also can be used as anode substrate 130 to anode 50.More desirable target layer structure, according to the label order be: (1) ground floor niobium of about 1 micron thickness of sputter on anode substrate, then on the ground floor niobium again the second layer tantalum of about 5 micron thickness of sputter (why desirable this structure is, be because the thermal expansivity of niobium between the thermal expansivity of the thermal expansivity of beryllium (anode substrate) 130 and tantalum, thereby reduce or prevented since between the on off state of X-light pipe the caused fine crack of thermal cycle of plate target generation); (2) the sputter tantalum layer of about 5 micron thickness; (3) the sputter tungsten-rhenium layer of about 5 micron thickness; (4) to be suitable as anode 50 be because they have bigger atomic number and density for sputter tungsten layer, tantalum, tungsten and the tungsten-rhenium of about 5 to 7 micron thickness, launches 3370 ℃ high-melting-point of X-light, tungsten and good vacuum characteristic easily and make it can adapt to high temperature and high vacuum condition in the X-light pipe when being shone by electron beam.As is known to the person skilled in the art, tantalum and tungsten-rhenium alloy have similar characteristic.The thickness of each of anode layer becomes the 100kV electronic switch the required distance of X-light through selecting effectively so that it is approximately equal to.Why beryllium is suitable as anode substrate 130 is because its intensity is very high, and can not decay significantly or X-light that scattering emits from anode 50.The thickness of beryllium anode substrate 130 is preferably about 0.5cm.The thickness of anode substrate 130 should be thin as much as possible within bodily form restriction, and said bodily form restriction is meant that it must have enough intensity, thereby is able to take to act on an atmospheric pressure gradient thereon.

Alignment grating 90 preferably is made of the array in hole 140, according to a most preferred embodiment of the present invention, and, each hole all is orientated or directed towards detector array 110.In other words, each hole on the alignment grating 90 is not parallel each other, and in order to be used in combination with for example Chest X-rays X-light device, they are all angled with the front surface 260 of alignment grating 90, its scope is that the middle section in alignment grating 90 is 0 °, fringe region maximum to grating 90 can be 20 °, and when the present invention was applied to breast disease diagnosis equipment, grating 90 can be made and make hole and front surface angulation scope reach 45 ° in the grating edge zone.The quantity of grating 90 mesopores 140 can be equivalent to the quantity of pixel, for example at preferably circular grating 90 middle bodies 500 * 500 to 1024 * 1024 holes is arranged, and the resolution that the quantity in hole to a certain extent can decision systems.Another kind method is to combine with double sampling technology discussed below and use the hole of lacking than pixel number.The aperture in the thickness of grating 90 and hole 140 is determined by the distance between detector array 110 and the X-light pipe (in this application, desirable value be 91.4cm (36in)), all sizes that do not enter the probe unit 160 of the requirement of X-light of detector and detector array 110 (not expressing) in this figure that decay.Although be not to be strict with so for the present invention, when front surface 260 was observed, hole 140 preferably presented the rectangle ranks form of rule and has the circular boundary that diameter is 25.4cm (10in).The hole array can be the detection any conventional profile relevant with the convolution switch technology of resolution object 80 images summarized below using.This hole array is called as " circular effective coverage ".Be preferably 500 * 500 according to the quantity of a most preferred embodiment of the present invention in the middle body hole of circular active zone.The imperforate section 150 of alignment grating 90 is designed to absorb invalid X-light, so that their irradiating objects 80 not.This can finish by the manufacturing to grating, even the X-ray of bump imperforate section 150 is subjected at least 10 times of stop (so-called " 1/2 value " are meant and can make by system capacity, be the X-ray attenuation 1/2 necessary quality of materials of the bump restraining mass of 100keV) to " 1/2 value " here.Invalid X-ray can make patient and medical worker be subjected to the irradiation of doses, but does not comprise any to the significant information of image.Shown in Fig. 3 A and Fig. 3 B, alignment grating 90 can be by some material layers 143 that absorb X-light, 144 constitute, many holes 140 are arranged on these layers of absorbent material, can allow X-light 100 therefrom pass to arrive detector .. alignment grating 90 preferably to use the thick molybdenum of 50 layers of 0.0254cm (0.010in) stacked and be fixed together and make.Why desirable molybdenum is, is because it absorbs the X-ray easily, thereby makes what those were produced by X-light pipe 10, the X-light of right rather than directive detector 110, at they unhelpful and irradiating objects 80 nocuously, this object may be exactly a human body certainly, just is blocked before.Plumbous or other similar X-light compactness materials also can use.

Hole 140 its xsects of alignment grating 90 are preferably foursquare obtaining the maximum density of arranging, and match with the preferred square shape of detector array elements 160.Other shapes also can be used, particularly sexangle.The size of square hole 140 is 0.0381cm (0.015in) * 0.0381cm preferably, like this its cross-sectional area approximately be generally used for fluoroscopy systems the collimating apparatus cross-sectional area 1/100.Owing to adopted strict collimation mode, so can obtain the X-light beam 100 of less beamwidth.The cross-sectional area that this means detector surface is much smaller than in the conventional system correspondingly.So just do not entered detector by the X-light of object scattering, do not cause image blurring the relatively large surface area detector thereby can not resemble to use in the conventional system to have.

A kind of desirable method of making alignment grating 90 is that light-chemistry erosion is cut or etching.Why desirable photochemical etching is, is because it is effective and accurate.According to this method, made 50 photomasks of a cover and on the thick layer material of 50 0.0254cm (0.010in), etched hole or space.Then the multilayer material of etching is stacked and align, be fixed together to constitute a grating assembly again, many step-like holes that are are arranged on this grating, each hole all keeps predetermined angular relationship with respect to each layer.Fig. 3 A represents the change of shape of alignment grating 90 of the present invention.This shape comprises that many layers can absorb the material layer 143 of X-light, and every layer has some long-pending holes 14 (still, xsect needs not to be constant) of constant section that have respectively.As shown in the figure, formed hole 14 is stair-stepping, but X-light beam 100 is passed through and arrival detector array 110.Shape is very similar shown in shape shown in Fig. 3 B and Fig. 3 A, itself is exactly step-like absorbing each hole that forms on the X-optical material layer 144 just.Position that these stepped bore can be offset a little on the two sides of material layer 144 adopts erosion to cut or the method for chemical etching processes as shown in the figure structure, and this is obvious for the person of ordinary skill of the art.The structure of Fig. 3 B is desirable especially, because absorbed X-luminous energy is less in the stepped bore 140 of alignment grating 90, so lack than the structure shown in Fig. 3 A in the damping capacity of the X-luminous flux of the marginal portion of X-light beam 100.

With constitute grating 90 a plurality of through the fixing more desirable methods of the material layer of etching as shown in figure 14.All fix the hole or fix hole 94 on each etch layer 91 (preferably 50).Register pin 95 inserts each and fixes in the hole 94 so that each etch layer 91 aligning.Then layer assembly 91 and register pin 95 are put into aluminium ring 359.Aluminium ring 359 has a vaccum exhaust outlet 370, and exhausr port 370 adopts compressing member 375 sealings.Then with the thick aluminium foil of 0.1cm 365 bondings and be sealed in the upper surface 380 of aluminium ring 359 with vacuum glue.Adopt on the lower surface 385 that in the same way aluminium foil 360 is bonded to aluminium ring 359.To be pumped into the partial vacuum in the aluminium ring by exhausr port 370, as well-known in the prior art, exhausr port 370 is adopted compressing members sealing 375 then.In this manner, the aluminium foil 360 and 365 of X-light has played and each etch layer 91 is fixed together and the clamping action that becomes a grating assembly 90 aligned with each other thoroughly.

Have the stepped appearance surface apart from hole 140 farthest, grating 90 centers, its xsect is square preferably.X-light can not be subjected to passage usually because the influence of the roughness that stepped surfaces causes had both made them be scattered, and also can not produce the influence that can measure to resulting X-light beam.As mentioned above, the material that is used as alignment grating 90 can be molybdenum, brass, lead or copper, and wherein molybdenum is the most desirable.The best tolerance of the position in hole is between the Center-to-Center during no accumulated error+/-0.00127cm (0.0005in), for the aperture then be+/-0.0025+cm (0.001in).

Other methods that can be used for making alignment grating 90 comprise electron beam processing, bore system or little processing, and laser bores.The shortcoming that system of boring and laser bore is their generations is circular hole but not square hole.Although circular hole can be worked equally, they are not preferred shape.

The more details of preferred scanning X-light pipe 10 is shown in Figure 4 and 5.Electron gun 161 is positioned at a relative side on X-light pipe 10 surfaces, and its work potential is about-and 100kV is to-20kV.Ground connection anode 50 is positioned at the surface of X-light pipe, and electron beam 40 navigates between electron gun 161 and the anode 50.The electronics orifice plate 162 of a ground connection is placed near electron gun 161, and the heart has a hole 163 to allow electron beam 40 pass therein.Magnetic focusing lens 164 and deflection coil 20 suitably are positioned at the X-light beam spot on the anode 50 by well-known dynamic focusing mode in the prior art.The X-light pipe is made with the circular active zone of 25.4cm (10in) diameter, and electron beam 40 can hit anode 50 in this circular active zone, can reach 30 ° of angles in the deflection of the edge of circular active zone electron beam.If electron beam is by specific hole " ejaculation ", then it preferably be deflected need not, thereby can save 25% energy.

Fig. 5 has represented the anterior xsect of a suitable X-light 10 pipes.The rear portion of anode 50 is the inside that keeps the X-light pipe of vacuum state.Anode 50 is made of the coating of anode material as discussed above.The front portion of anode 50 is the thick beryllium anode support 130 of 0.5cm.The front of beryllium anode support 130 is that preferred thickness is 0.4cm, can hold the cooling sandwith layer 350 of water or gas-pressurized.Aluminum grating support chip 360 and 365 thickness are 0.1cm, and being used for the supplemental support preferred thickness is the alignment grating 90 of 1.27cm (0.5in).

When 340 work of X-light pipe, in any one given moment, alignment grating 90 has only a hole 140 to allow most X-light pass through.According to a preferred embodiment, when electron beam 40 is not that it will be stopped when just in time being positioned at the front in a hole 140.Therefore the X-light pipe can be effectively with the damage of scanning impulse pattern work with minimizing energy consumption and antianode target 50.Three-dimensional X-photoimaging

Refer now to Fig. 6, according to a further advantageous embodiment of the invention, can use X-light, thereby obtain three-dimensional X-light image with multi-focus point.For example, if hole directed toward focal point F1 (92) on the grating 90 in every line, remaining hole directed toward focal point F2 (93), and locate to place the first sensor array at F1 (92), and locate to place second sensor array at F2 (93), just can be with net form or these holes of spirally graph scanning, thus the data line of first sensor array obtained, be the data line of second sensor array then.Repeat this process, can set up two complete images, resemble two different points from the space, F1 and F2 see, and then adopt conventional stereoscopic image display system to show that they are to produce the X-light image of a solid.Refer now to Fig. 3 C, how to the figure shows the alignment grating that the material 144 that can absorb X-light with multilayer constitutes such solid.In this case, hole 140A and 140B in fact can make " V " shape as shown in the figure, have formed two paths that separate of X-light beam 100A and 100B along " two legs " of " V " shape.Yet and do not require necessarily with hole 140A and 140B in conjunction with the advantage that forms illustrated shape " V " shape hole be when X-light when the summit of " V " shape enters, two detectors can be shone simultaneously, this " V " shape has been played the effect of demultiplexer, make a part of X-light directive F1, another part directive F2.This makes for forming the X-light beam and producing the required energy in half of deflection current.And cost just very little but increased acceptable X-scattering of light and the image that therefore causes fuzzy.Detector array

In order to reach the resolution of every millimeter some lines on object plane, as desired in some medical applications, the restriction of spatial resolution mainly is that the size by detector is determined.This is because under existing X-light pipe technical merit, can not produce very high energy or be equipped with relevant X-light orienting device in order to obtain enough strong directed X-light emission.

When detector is made into its area when intersecting area with the taper of the X-light of emission, will can not hit detector 250 by most of X-light of X-light source 50 emissions, shown in Fig. 7 A.In fact a key how designing of this industrial just scanning beam digital X-optical detection system, radiation dose is not the problem that needs are considered in these systems.Consequently Enhanced Radiation Reduced Blast dosage is to keep required resolution.

So, improved resolution by using minimonitor, but when the area of detector equaled or exceeded by the crossing area that forms of the X-light cone of emission and detector plane 270, the X-light dosage became minimum.

The resolution of scanning X-photoimaging system by detector cells be projected in the object plane 280 (promptly placing object 80) with the plane vertical, the center of anode 50 with the line of centres of detector 110 cross-sectional area determined.Therefore, if a large-area detector is divided into many less array elements, shown in the front view of detector array among Fig. 8, then kept the big area of catching of combination detector, simultaneously kept visual resolution again, it is proportional to the size of single little detector cells 160.

In a storage register, distribute and addition by the signal that will from each detector cells, gather, i.e. each address therein, be the certain location of pixel, can keep the resolution that is limited by single detector unit 160 corresponding to object plane 280.When X-light beam 100 is being positioned on the alignment grating 90 of X-light emission anode 50 fronts when mobile discontinuously, the continuous address of output that is used for given detector cells is also changing.The geometric relationship of imaging is shown in Fig. 7 B and 7C.Represented in Fig. 7 B how the position of a branch of light and it are assigned with in 5 pixels.Represented then in Fig. 7 C how the light-beam position that continues and they are superimposed upon on the pixel together.

In other words, one of the signal of each detector cells is corresponding to specific region very little on the object plane 280, and promptly the memory address of a pixel is stored in the image register.Therefore corresponding to the memory address of each detector cells along with changing with the position of the X-light beam of given graph scanning, thereby make each pixel in the storer comprised a specified point by object plane 280 radiant quantity and.Like this, owing to all X-light that arrive detector planes 270 in fact all have been recorded, so the resolution of system is determined that by the big or small of single detector unit it is best that the sensitivity of simultaneity factor also reaches.

Another benefit of this detector array imaging geometry is that object plane 280 is limited very narrowly.Be positioned at before it or structure afterwards will be by obfuscation (beyond the focus).Fig. 7 D has represented that the X-light of launching passes the object plane 280 of the 141 and 142 certain distance SO apart from the hole and the plane 281 apart from SO of 141 and 142 twices apart from the hole from first hole 141 and second hole 142.Be easy to find out, be reduced in 1/2 of the resolution at distance SO place in the resolution of twice SO distance.This feature is used for improving location and the developing in the thin portion structure on the plane of being detected 280, and the enough depth of field is provided simultaneously, and this can be improved by the geometric relationship of system.

The array of preferred embodiment is the quasi-circular array of one 96 unit, and it is to be that the square detector cells of 0.135cm is arranged in the circle that a diameter is approximately 1.93cm (0.72in) and constitutes with the length of side.It does not need so big, can dispose three or more detector and constitute, thereby not every detector cells all equals detector length on one side at radius, is on the line in the circle of 0.135cm here.The X-photo-detector

Conventional image enhancement technology has basic restriction for the sensitivity of the system of qualification.The thickness of operable scintillator material is limited to by its saturating characteristic of light.Usually their are done enough thickly in can catch incident X-ray-photon of about 50%.In the photon that is launched, has only only about half of arrival photocathode.At photocathode, has only about 10% incident photon generation photoelectron.Therefore, have only that about 2.5% (energy of the incident X-photon of .5 * .5 * .1) is changed in the image enhancement system.Except this conversion efficiency that is restricted, photon also can be by the scintillator material transverse scattering, and produces haze, and this has caused the reduction of system resolution under given radiation dose level.

A basic purpose of the present invention provides a SBDX imaging system, and it should be able to guarantee that detected object in that image quality is satisfied under the desired prerequisite of diagnosis of being carried out fully, is subjected to the irradiation of the X-light of possible minimum radiation dose.This means that the system that is used for surveying the X-photon that sends from object must have the highest photon-electrical signal conversion efficient.In order to realize this point, material as detector must have enough length to guarantee not having photon to run away from the far-end of incident X-light on the direction of photon flight, promptly the energy of photon must be dissipated in the material for detector fully so that make the output maximum of detector.There is the detector of several types can be used to said SBDX system here.Wherein preferably the X-photon energy of scintillator-in this scintillator be converted into visible light energy-then by a photomultiplier cell, photodiode, CCD or suchlike device convert light intensity to an electric signal.Because each pixel in the SBDX image must be in the very short time, about 140 nanoseconds form, so scintillator material must have response and the shortest after time fast.Twilight sunset is meant that scintillator continues luminous phenomenon under the situation that incident X-ray has disappeared.Plastic scintillant, as the organic scintillator that contains polystyrene is suitable for, because they have the characteristic of required quick response, but their X-light photon action section is but less relatively, so the value of their the linear X-absorption coefficient of light is also less.Consequently to intercept and capture all X-light photons; Need suitable thickness, for the X-light of the preferred 100keV of the application, catch 99% incident X-light, general plastic scintillant is must about 28cm (11in) thick.At present the most preferably (according to preferred order): the YSO of (1) doped with cerium (oxygen orthosilicic acid yttrium, Airtron (litton) of Charlotte company produces); (2) LSO of doped with cerium (oxygen lutetium orthosilicate, Schlumberger company product; (3) BGO (bismuth germanium oxide, Rexon Components, Inc.ofBeachwood, OH company product).The advantage of YSO and LSO is that they can be used for room temperature.BGO must be heated to about 100 ℃ with reach 50 nanosecond the order of magnitude the light output attenuatoin time that is fit to.These scintillators do not need to resemble the plastic scintillant long, and length is that 0.10cm is just very effective.

According to a preferred embodiment of the invention, 1 of being made up of the discrete X-photo-detector 160 of 96 intensive encapsulation of SBDX detector array 110 takes advantage of 12 director circle array to constitute, and this array and X-light source 50 are placed at a distance of 91.4cm (36in).(5 * 5 and 3 * 3 arrays also plan to constitute a non-square array, promptly are filled in the circle of X-light target face by foursquare detector: the table 1 below for example visible).The aspect ratio of alignment grating, it is 1.46 ° that the interval of it and X-light source 50 and the geometric parameters such as size of X-light source 50 form a total angle, and intersecting length on the sensitive surface of detector array 110 is the X-light pyramid of the square sectional of 2.23cm (0.9in).So, must be between each scintillator 170 Center-to-Centers in the detector plane at a distance of about 0.152cm (0.06in).If scintillator 170 has parallel edges, then will bump against the wall of scintillator from the X-light of its edge incident desired distance of not flying as yet.If so the not shielding of adjacent scintillator, these X-light may pass adjacent scintillator, make it produce one, thereby cause the reduction of image quality as if from the signal output of the locus of the mistake of object.Shown in Fig. 9 A, for fear of this effect, in a preferred embodiment of the invention, each scintillator all makes cone-shaped, thereby make its border surface 173 have angle α, this angle equals the angle between the X-light beam 100 ' outermost of incident, and this is useful especially for long plastic scintillant.In the preferred embodiment of being quoted, each scintillator 170 is preferably all made the long prismoid of 28cm in the above, its sensitive surface (172) length of side 0.285cm, and the length of side of photo-detector end surface 174 is 0.37cm.So whole 81 detectors of a small bundle of straw, etc. for silkworms to spin cocoons on have the end face that is made of a plurality of facets, each facet is all tangent with the sphere that with X-light source 50 is the center.

Further improvement for the scintillator detection efficiency can realize greater than the angle of incident X-light beam 100 by the interior angle that makes the scintillator pyramid.In the shielding material that the X-light possible loss of incident X-light and photoelectron that produces near the scintillator atomic interaction at scintillator edge and scattering is being used for adjacent scintillator is separated.The photoelectron of these losses no longer produces any light, so they are exported without any contribution the light of scintillator.So their loss has reduced the efficient of scintillator.The ultimate range of photoelectron flight depends on its energy and its material of flying therein.For in plastic scintillant with the X-light of the 100keV of atomic interaction, photoelectronic flying distance can be greater than 0.01cm.If the angle of scintillator pyramid frutum is greater than the angle of X-light beam 100, thereby its size is greater than X-ray envelop face 2 * 0.01cm, therefore in short distance (28cm) lining of ratio detection device length, because the reduction of the efficient that the loss photoelectron causes will be reduced to minimum level.In this case, no longer consistent with the center of the tangent sphere of facet with X-light source 50, but nearer apart from detector array 100.

The photon that is scattered will be longer than photoelectron flying distance; So, escaping in the adjacent scintillator in order to prevent them, the tapering of scintillator pyramid can be greater than catching the required tapering of photoelectron, so that the capture rate of scattered photon reaches maximal value fully.

Referring now to Fig. 9,, according to a preferred embodiment of the invention, what link to each other with each scintillator elements 170 is photoconductive tube or fiber optic cable 180 by corresponding photomultiplier 190 or solid probe and 170 optically-coupled of each scintillator.Also scintillator 170 directly can be placed on and the approaching place of the photo-detector that is fit to.

Figure 10 has represented the preferred structure of a detector cells 160.Placed the X-light shadow shield 200 that has with each detector cells 160 corresponding holes 210 in the front of detector array 110.Each detector cells 160 is encapsulated in one also in the close shell 220 of light of not saturating X-light.Be provided with a visible light shielding window 230 of preferably making by flake aluminum in the front of the close shell 220 of light.Light shield window 230 is for only transmission of X-.In the close shell of light 220 the insides a scintillator unit 170 is arranged, it is close to a photo-multiplier 190, and photo-multiplier then is electrically connected with a prime amplifier 240.Preferably will convert digital signal to be further processed with conventional method from the simulating signal of prime amplifier 240.

Perhaps, scintillator also can with photodiode, photistor or a charge-coupled image sensor (CCDs) array directly or closely contact placement, to make firmer and compact detector.Using solid-state devices, particularly during CCDs, can use cooling medium, for example Peltier type cooling valve or person like that are to increase the signal to noise ratio (S/N ratio) of device.

Perhaps, with scintillator arrays and one or more location sensitive, can produce the position coordinates of determining light source and light source intensity output signal photo-multiplier directly or closely contact place.

In another preferred embodiment, sensor array can also be to be made of one group of pencil-type detector 285 array, for example, and as shown in figure 11.The scintillator 290 of taper is arranged on the path of X-light beam 100 in Figure 11, thereby will be absorbed in X-light in this transverse cross-sectional area fully corresponding to the scintillator in the certain cross section zone of X-light beam 100.Photo-multiplier 300 is close to scintillator 290 and is provided with, thereby the absorption of 290 pairs of X-light of response scintillator will produce an electric signal.Can replace photo-multiplier 300 with solid-state devices.

According to a preferred embodiment of the invention, on the length direction surface of scintillator and input surface is gone up and is covered a kind of reflectorized material, and silicon dioxide for example is to prevent light loss (perhaps entering) and to help the reflection of light in scintillator inside.

According to a further advantageous embodiment of the invention, between each scintillator unit 179 and the scintillator unit 170 adjacent with it with a kind of extremely not material of saturating X-light, for example gold or plumbous thin layer 171 isolated opening.The thickness of thin layer 171 preferably is about 0.0102cm (0.004in) to 0.0127cm (0.005in).The position of thin layer 171 between scintillator 170 as shown in figure 12.

As shown in the figure, the area of the circular effective coverage of alignment grating 90 is greater than the area of detector array 110.Although therefore single X-light beam 100 is dispersed or expanded as the flash of light bundle, the pencil beam of the X-light of launching from the corresponding hole 140 of alignment grating 90 all converges to detector array 110.Image processing

An important improvement of the present invention relates to uses an image processing system with the desired radiation dose of further minimizing.In fact, be not directly inputted to " z " or the luminance input of video monitor usually from the signal of detector.But the intensity data of the process digitized processing of each pixel is stored in each address in " frame memory buffer register ".Can use more than such buffer register in some applications.Pixel addresses can be chosen randomly in the buffer register, and digitized intensity level can carry out mathematical computations to it.This function is used when using multiple image enhancement algorithm, and can be used for distributing from the pixel of the data of the various piece of detector array.

According to a preferred embodiment of the present invention, a SBDX image can comprise about 250,000 pixels at most, is arranged in 500 row, 500 row (holes of taking advantage of 500 row corresponding to 500 row of the central authorities of alignment grating 90).In order to explain following example, suppose the X-light source certain concentrate in a flash alignment grating 90 the 100th the row, 100 column position places pixel P on.Suppose that further detector array 110 in this embodiment is arrays 110 of one 3 * 3, comprise 9 district 179 (Figure 12), and the size in each district 179 can be returned and received the whole X-optical radiation that are associated with a pixel.Describe in detail as the application, other array structure obviously also can use.

The numerical value that obtains from the measurement result digitizing in each district of detector array 110 is assigned to pixel addresses as described below:

1 district, 99 row, 99 row

2 districts, 99 row, 100 row

3 districts, 99 row, 101 row

4 districts, 100 row, 99 row

P district 100 row, 100 row

6 districts, 100 row, 101 row

7 districts, 101 row, 99 row

8 districts, 101 row, 100 row

9 districts 101 are capable, and 101 broomrapes scanning X-light beam repeats same data allocations form when passing through all pixels.

In shown image, the digital value of each pixel equal " n " individual part and, wherein n be subregion 179 in the array 110 number (in this example, n=9).

When the structure of detector array 110 is shown in the application, has the fixing effect of the operating distance of making, can obtain best focal length in this distance, and can be created in not subregion detector array SBDX imaging system of the prior art the focal plane of the best that can not obtain.

When the design detector, must consider following parameter:

1. the thickness and the shape of X-light beam through collimation of launching from X-light source (plate target) 50;

2.X-the distance " SD " between light source 50 and the detector array 110;

3.X-light source 50 and by the distance " SO " between the center of the object 80 of imaging;

4. desired resolution is perhaps by the pixel size on the object 80 of imaging;

5. in medical applications, the total area of array must enough receive all X-light that send returning greatly from alignment grating 90.

In the SBDX of a preferred embodiment of the present invention system, the distance between the bright dipping side 260 of X-light source 50 and alignment grating 90 is about 2.271cm (0.894in) (seeing Fig. 3,5).The size in hole 140 is that 0.0381cm (0.015in) * 0.0381cm (0.015in) is square.The diameter of the luminous point of electron beam 40 on anode 50 is about 0.0254cm (0.010in).Detector array 110 is 91.4cm (36in) apart from the distance of anode 50.Therefore the beamwidth of X-light beam 100 is 2 ^*ARCTAN ((spot diameter/2)/((width in hole/2)+(spot diameter/2)) ^*2.271cm (0.0894in), perhaps equal 1.6 °.Distance at distance anode 50 is that 91.4cm (36in) locates, and the X-beam diameter of projection is 91.4 ^*TAN (1.6 °) cm.So for preferred embodiment, detector array 110 length on one side should be 2.54cm (1in).For example, if by the object of imaging apart from anode (50) 22.86cm (9in), the size of required pixel on object is 0.0508cm (0.020in), the X-light source also is 91.4cm (36in) to the distance of detector, it is square that detector array has optimum dimension 2.54cm (lin), then the projected size of pixel on detector plane 270 be pixel size on the object (SD/SO) doubly, perhaps 0.2032cm (0.080in).Divided by 0.2032cm (0.080in), we see can obtain being equivalent to the required resolution that every limit has the square subregion detector array in 12 to 13 districts with 2.54cm.Obviously, the environment that uses according to the SBDX system can also adopt other structure.

At optimum resolution planar S O (280 among Fig. 7 D) in addition, at 0.5 * SO place with at 2 * SO place (among Fig. 7 D 281) resolution with drop by half.This has had the rational depth of field for great majority are used.In some applications, such as to the imaging of human heart the time, the decline of resolution is considered to be good outside field depth.The fuzzy identification that can increase of the details beyond institute's interesting areas to the details in institute's interesting areas.

There are many methods can be used for from above-mentioned obtained data, obtaining a useful image.As mentioned above, a kind of simple convolution transform can be used, but best resolution can't be all reached in this case.There are other two kinds of preferable methods can be used for obtaining best resolution and sensitivity in this application from the data of being obtained.They are called as how visual convolution transform method and many output convolution transform methods.In both cases, all done following supposition:

Capable and the APy row hole of APx is arranged on alignment grating 90.Row is that one " pixel " do not contributed any measurable brightness as them for image at the pixel of the circular active zone outside of alignment grating 90 with each intersecting area of row, and promptly they are handled as " black ".The pixel that is not shone by X-light 100 in scanning process does not contribute measurable brightness to handle as them for image too,, promptly they are used as " black " processing.

Refer now to Figure 15, in detector array 110, have maximum DETx row detectors unit 160 and maximum DETy row detector unit 160 to form a quasi-circular detector array 110.

ZPATIO be one 0 and 1 with real number.If ZRATIO=1, then focus is set at detector plane.If ZRATIO=0, focus is set at X-light source plane.If ZRATIO=0.5, focus half distance between X-light source plane and detector plane then, or the like.PIXELRATIO is the number of image pixel in each section actual range between the detector adjacent in a row or column.For example, if between the pixel center be 0.01cm at interval on the object plane 280, between the detector in detector plane 270 is 1.0cm at interval, then PIXELRATIO=10.FOCUS＝ZRATIO ^*PIXELRATIO.

IMAGE is the data matrix of a DETx * DETy dimension, is wherein comprising once specific scanning and corresponding to the monochrome information of a particular pixels.PIXEL is the four-matrix of an APx * APy * DETx * DETy, is wherein comprising the DETx * DETy pictorial data matrix that obtains by all (perhaps part) holes of scanning.According to a preferred embodiment of the invention, PIXEL is updated after each scanning.

When beam scans on anode surface, be actually the front, center " emission " that earlier beam is positioned at selected hole 140, and then the location.So, all obtain an IMAGE data matrix for emission each time.When these images can be configured the displayable image with some direct purposes, by they comprehensively can be obtained higher resolution and sensitivity.The first kind of preferable methods that is used for synthetic image is called how visual convolution transform method.In how visual convolution transform method, by give matrix element OUTIMAGE (x, y) down train value can constitute the luminance matrix OUTIMAGE of an APx * APy that can show on CRT or similar display device:

The second kind of preferable methods that in this application the data matrix IMAGE of APx * APy is attached to a useful picture is called as many output convolution transform methods, in this case, corresponding to the detector array that constitutes by DETx * DETy detector, need DETx * DETy Aristogrid (or its equivalent, multipath conversion) and same number of pixel adduction circuit.The value that each detector is digitized be called as SENSOR (j, i).Final OUTIMAGE matrix calculates in the following manner-for output image matrix OUTIMAGE (y, x) each pixel in [to y=1 to APy and x=1 to APx], with each DETx * DETy source image SENSOR (j, pixel in i) and the pixel OUTIMAGE (y-j in the purpose picture matrix ^*FOCUS, x-i ^*FOCUS) addition [to DETy, i=1 is to DETx to j=1].Pass through each matrix element divided by DETx then ^*DETy is to the normalization of OUTIMAGE matrix.

Further improvement for these technology can realize by the fraction part of the FOCUS factor being carried out the linearity insertion.

The advantage that many visual convolution transform methods are compared with many output convolution transform method is that the former can select best focal plane with software after obtaining data, and the latter can not.But a kind of method in back can be handled under the situation of limited time system quickly.Utilizing the SBDX data to carry out three-dimensional image rebuilds

SBDX system described in the application can be used to produce a series of plan view images, and these images can be used to constitute the tomography or the 3-D display of object 80 again and can analyze image sets.By the various values of using FOCUS image data set is analyzed again and can be generated a three-dimensional image that is included in a series of images under the various degree of depth.Used FOCUS value naturally is respectively n/DETx or n/DETy, wherein n be one from 0 to DETx or the integer of DETy.Usually only analyze corresponding to object 80 interior by those FOCUS values on interested plane.For example, table 1 (below) in the described SBDX system, there is the interval of about 2.54cm (1in) focal plane and the focal plane of locating near regular 22.86cm (9in) (best focal plane).

Following formula is to have represented the position of serial plan view image to the distance of anode 50. Wherein Ft (FOCUS)=from anode to the distance of interested specific focal plane

The distance of Fd=from the detector to the focal plane (anode to the distance of detector less than Ft)

Interval on the λ t=alignment grating between the Center-to-Center of adjacent holes

Interval in the λ d=detector array 110 between the center of adjacent detector 160.

When using the double sampling technology, account form does not change the data that only processing is not obtained by the hole of " jump " on the alignment grating.But, the hole dislocation on the alignment grating was arranged.λ t also keeps same value.The control of negative feedback X-luminous flux

See Figure 13 now, wherein represented to use a negative feedback paths 305 to control the SBDX imaging system of the luminous flux of X-light beams 100.Preferably utilize the negative feedback control X-luminous flux that is connected with detector array, thereby make detector array always receive essentially identical luminous flux.According to the method, when scanning, reducing the X-luminous flux, to reduce to the total radiation dose of patient's (or object) to human body soft tissue (they are than being easier to see through X-light).Utilize the control of negative feedback flux can also improve contrast and dynamic range.According to this embodiment, differential amplifier 310 have one can by the user be provided with adjustable with reference to level 320.Feedback loop 305 feedbacks are connected to X-light pipe 10 with control X-luminous flux.The time domain scan pattern

Utilize the disclosed principle of the application can also realize a kind of time domain X-photoimaging system.In this system, the time that the X-light of outgoing from each pixel reaches predetermined measured X-luminous flux can calculate and draw.So can use negative feedback control to block or reduce the X-luminous flux that from hole, emits corresponding to the pixel that in the scan period of being considered, has reached the predetermined flux level.In this case, the information of being gathered will be the information of time to amount of flux, draw or the information of imaging will be corresponding to time rather than brightness.A kind of like this system has the signal to noise ratio (S/N ratio) that can provide very high, improved contrast, greatly reduces the X-light dosage of detected object and the potentiality of improving dynamic range.Multipotency X-photoimaging pattern

According to a preferred embodiment of the present invention, can or organize X-light beam 100 more and point to one or more detector array two groups.First group of X-light beam has the first feature X-energy spectrum.Second group of X-light beam has the second different feature X-energy spectrum.Whether by the transmitance of first and second groups of more measured X-light beams, can detect has certain class material to exist in detected object.The key concept of difference X-photoimaging technology is known in the prior art, for example be disclosed content in the U.S. Pat-5185773 of " the harmless selectivity pick-up unit and the method for metal " in denomination of invention, this patent documentation is quoted as a reference in this application.

These two groups of X-light beams can produce with many methods.One of them method is to realize by making a kind of special anode 50, and this anode has certain material of adjacent with the first group of hole first kind of material or first thickness and certain material of adjacent with the second group of hole second kind of material or second thickness.Like this, the X-light that emission is had the first feature power spectrum with first group of relevant hole, and the X-light that emission is had the second feature power spectrum with second group of relevant hole.Perhaps, can use K-optical filtering technique (or K-edge optical filtering technique), promptly in the part in hole 140, place filter (for example, molybdenum) to produce same effect.In this case, comprise insertion first light filter wherein in first group of hole, comprise insertion second light filter wherein in second group of hole.Second light filter can be not have light filter at all.Described in situation in front, two groups of X-light with different characteristic power spectrum are associated with two groups of holes.

As long as be correlated with different feature X-energy spectrum, just might detect Microcalcification (breast cancer is early stage) and other use wide power spectrum X-light to be difficult for observed unusual condition by at least two group holes.For example, by scanning first group of hole to form one first image, scan second group of hole then to form one second image, decompose these images and increase their contrast, just can utilize low dosage scanning beam X-photoimaging system to detect this type of unusual condition of Microcalcification and other in real time.Equally, can use multi-detector array, allow first group of hole point to first detector array, and allow second group of hole point to second detector array, or the like.

Introduce another embodiment of multipotency imaging now.Because the electric pulse amplitude that the X-photon of being surveyed produces is proportional to the energy (KV) of photon, photon pulse is separately counted so can be with at two or more.Pulse separates by intensity, and technology is also handled respectively then, thereby constitutes the image that two or more separate.These images can show with ratio.

The energy value through selecting that can also change emission with differentiation of objects in different density areas.The advantage of this embodiment is than those cleverer drawing together described above, does not need special alignment grating, anode material, perhaps twofold detection device.

Many embodiment of various structures of the present invention have been discussed above, and following explanation will be explained preferred SBDX imaging system of the present invention:

1A. ：：25.4 ( 10 ) ：0.0508 ( 0.020 ) ：500 ( 166 ) ：506.45 ( 78.5 ) ：196350 ( 21630 ) ：：0.0381 ( 0.015 ) ：0.0127 ( 0.005 ) ：2.5cm ( 0.98in ) B.-：91.4 ( 36 ) ：22.86 ( 9 ) C.：30D.X-：70-100kVE.：： ( 15 ) ：：0.135×0.135 ：1296 ：1.83 ( 0.72 ) ：15.8° ：19.05 ( 7.5 ) ：0.038 ：0.152 ：13/

So far, shown and introduced a SBDX imaging system that adopts the subregion detector array that this system has high resolving power, high sensitivity and simultaneously to the low X-optical radiation dosage of detected object.This system also allows pinpointed focus is arranged on any point between light source 50 and the detector array 110, and has the effective work depth of field.Bundle double sampling technology

Following content relates to a particularly preferred embodiment of the present invention, and this embodiment has adopted bundle double sampling technology to reduce the energy consumption of Computer Processing non-productive operation and scanning beam digital X-photosystem.

The image of normal video quality adopts 640 * 480 pixels and with 30Hz frequency new data more.This requires the pixel sampling rate to be about 12MHz.With such speed the high-pressure electronic bundle of X-light pipe is accurately positioned in the different back, hole that 250000 orders arrange and needs high degree of accuracy and quite high energy consumption.With the big X-photodetector array of 12MHz speed collection signal digitalized be expensive and power consumption equally.Therefore under the prerequisite of the space of not obvious reduction SBDX system or temporal resolution, the pixel sampling rate is reduced to help reducing the original equipment cost below the 12MHz, because the running cost that the used heat that electric energy consumption and cooling are produced by the X-light pipe causes.

Therefore, developed and a kind ofly be used to reduce the pixel sampling rate, and the mechanism of identical in fact room and time resolution is provided simultaneously.This mechanism is called as double sampling by journey, and preferably and the embodiment in the SSBDX system described in this joint be used in combination, although it obviously also can be used to the SBDX system of other structures.The advantage of this embodiment comprises and cuts down the consumption of energy and simplify the circuit that is used at X-light pipe deflection beam, reduce the cost of making alignment grating 90, reduce the required complexity of calculation of object analysis 80 images and other those advantages that obviously exist to those skilled in the art.

According to this embodiment, the decreased number in alignment grating 90 holes of manufacturing although can be other numbers, be preferably APx=APy=166, rather than 500.The viewpoint that the advantage of this minimizing is calculated from the following stated becomes fairly obvious.But, only need this structure in hole of manufacturing 1/9th quantity simpler from the viewpoint of making.Because the minimizing of the number in hole, easier manufacturing has the grating (that is, the hole is with respect to front surface 260 angulations of alignment grating) than high deflection angle, and the problem that hole and adjacent holes intersect can not occur.This is very useful when making 3 D grating, because adjacent hole is to point to different detector arrays in 3 D grating, therefore requires more substantially to separate to avoid intersecting of hole than between non-3 D grating mesopore and the hole.

The hole of alignment grating be arranged on out to out be APx capable * circle of APy row in.For computation purpose can with its as APx capable * the square processing of APy row, but the part beyond circle is not contributed any information, always promptly " black " or not by the X-rayed.

Show as Figure 15, the detector 160 of X-photodetector array 110 be arranged to out to out be DETx capable * circular array of DETy row.Than the hole that alignment grating hole sum lacks, promptly, can reduce the pixel sampling rate by irradiation by double sampling.Preferably adopt and do not have the not alignment grating in irradiated hole.In order to constitute an image by detector array, have only each DETx collimator holes of each row and each DETy collimator holes of each row to need illuminated, so can constitute this image with the pixel unit of image, the size of each image is DETx pixel * DETy pixel.This is equivalent to the double sampling ratio of DETx * DETy, although do not have double sampling corresponding to 1 * 1 double sampling ratio.Therefore the double sampling ratio can be adjusted from 1 to DETy (row) in the Y-direction in X-direction (OK) from 1 to DETx.According to this preferred embodiment, DETx=DETy=12, as shown in figure 15.

Using 12 * 12 detectors and double sampling than being that this image is to be actually " gluing " not superimposed image formation together-extraordinary image one width of cloth David Hockney by one group to splice photo (photomosaic) under 12 the situation.Because scintillator in the reality and detector can not fully accurately and as one man respond, X-light cone shape pencil neither be very even, hole on the alignment grating neither be accurately in same zone, and owing to used circle, rather than square detector, so extremely need to a certain degree overlapping, thereby the non-linear and noise of detector is on average neutralized.

If double sampling is than less than the detector yardstick of representing with pixel (in other words, in this preferred embodiment less than 12), image must constitute with overlapping " tile ", and they must be added or quilt is asked average certainly.If double sampling than be not detector yardstick (representing) with pixel if even-multiple or detector array be not square, then on each pixel, will add the sampled value of different numbers, and each pixel will be asked the different divisor factor of average needs.The technology that is used to handle these pixels that are less than ecotopia is known for a person skilled in the art, do not need open here, to avoid making instructions too complicated.

In the calculating below, SSX is illustrated in the double sampling yardstick of X-direction (OK), and SSY is illustrated in the double sampling yardstick of Y-direction (row).For example, if SSX=SSY=1 does not then carry out double sampling, processing procedure and other embodiment of the present invention discussed above are just the same.Equally, in the present embodiment, if SSX=SSY=12 then presents " the photo shove joint " that do not have pixel average again.If SSX and SSY are 3, the size of circular active zone is 500 * 500, will have 166 * 166 hole to be scanned so, promptly the X-direction 1/3 and in 1/3 of Y-direction, thereby obtained data are reduced to 1/9.Note,, also just do not need them, thereby also just there is no need on alignment grating, to make these holes if only use 1/9 hole all the time.

So in order to form the X-ray that can generate an images, (500 * 500 holes) has only 1/ (SSX in former alignment grating ^*SSY) hole need be used or be shone by electron beam.If frame rate remains constant, 30Hz for example, then the number of times of electron beam scanning reduces to 1/SSX ^*SSY doubly shows as and drives the circuit correspondent frequency of electron beam.Total flying distance (and number of scanning lines) of electron beam reduces to 1/SSY doubly, thereby the average Shu Sudu on the target anode reduces to 1/SSY doubly.Image reconstruction pixel speed identical with alignment grating hole speed (speed that the hole is scanned or shines), and also be reduced to 1/ (SSX ^*SSY) doubly.

According to this scheme, the number of samples that on average arrives each display element is (DETx/SSX) ^*(DETy/SSY).When using the maximum secondary sampling scale, when SSX=DETx and SSY=DETy, each display element on average has only a digitizing sample value (" photomosaic " pattern).Sample value is important for the heterogeneity of smooth beam, scintillator, detector and amplifier on average.The quantity of double sampling (SSX and SSY) must be set in the level that adapts with surrounding enviroment, to guarantee to obtain qualified image quality.This can be adjusted for the requirement and the specific environment state of image quality according to the user in operation by the user.

Detector array 110 shown in Figure 15 preferably is arranged on the array that constitutes in the rounded basically zone that diameter is about an inch by 96 single detector cells 160.12 detectors (DETx) are arranged on the vertical row of array center, 12 detectors (DETy) are also arranged on the horizontal line of array center." egg basket " support structure that scintillator crystals preferably cuts into square horizontal section and made by 0.005 inch thick stainless steel slip.The diameter of the residing circle 400 of all scintillator crystals among Figure 15 (drawing the hatched part in cross section) preferably is about 0.800 inch.

The preferably about 0.10cm of the length of scintillator in the detector array 110, preceding input surface is 0.135cm * 0.135cm preferably.Scintillator crystals is YSO, LSO or BGO preferably, but other materials also can use as mentioned above.For being the die-away time (approximately 50nm) that suitably reduces for the light output of scintillator in this application, BGO need be heated to 100 ℃.So a stratie need be provided.

Figure 17 represents detector assembly 402 according to a preferred embodiment of the present invention.X-light passes X-optical window 404 and enters lead shield 406 from the top.X-optical window 404 is preferably circular, and its diameter is approximately 1.91cm (0.75in) so that the X-light of launching from the hole of alignment grating 90 bump in the scattered light decay enters detector array 100.A light shield 408 is used to make the light of detector shielding from the opposite way round.Aluminium foil or beryllium paper tinsel that it can select to approach are made with attenuate light unattenuated substantially X-light.The thickness of paper tinsel is 0.0125 centimetre.

Detector array 110 is provided with near combine the heating element of selecting for use 410 with the BGO scintillator.Heating element 410 can be a resistive heating element, and it can make detector array 110 remain on about 100 ℃ working temperature.Article one, imaging fiber conical conductor 412 will import the photo-multiplier (PMT) 416 of 96 passages from the photon of bottom 414 outgoing of detector array 110.Detector assembly 402 is sealed in the close shell 418 of light, produces noise to prevent veiling glare.Have three shoulder screws 420 and three central screws 422 to be used for plane and linear collimation location, this is well-known to those skilled in the art.Positioning of rotating is by realizing shell 418 with respect to 426 rotations of PMT erecting frame.Imaging fiber cone-shaped body 412 can be available from Collimated Holes of Campbell, CA, and it has circular input hole and the circular delivery outlet that diameter is 3.38 centimetres (1.33 inches) that a diameter is 2.03 centimetres (0.8 inches).Cone-shaped body 412 is 0.06 inch with the spacing of each scintillator crystals coupling, and the spacing between the PMT416 is 0.10 inch, has promptly amplified 1.667 times.On two surfaces of cone-shaped body, used the suitable high viscosity optically-coupled liquid (200 type) of producing by DowCorning of the refractive index of refractive index and glass as a kind of optical coupling medium, so that from scintillator crystals 160 to cone-shaped body 412 and from the light transmission efficiency maximum on cone-shaped body 412 to PMT input surfaces 424.

Photo-multiplier 416 is pipes (passage is corresponding to a scintillator crystals 160) of one 96 passage, can adopt the XP1724A type of being produced by Philips company.It has a fibre faceplate, thereby can carry out space orientation accurately scintillator arrays is aimed at the PMT photocathode of another the lip-deep PMT that is arranged in panel.The X-light photon that hits a scintillator 160 produces a light pulse of being coupled to the PMT photocathode.This has produced a corresponding electric pulse in photocathode, this pulse is amplified to 1000000 times in a passage of PMT multiplier tube structure.

This PMT output pulse is connected to the input end of a 30MHz bandwidth amplifier, and its output pulse is in 0.5 to 5.0 volt scope, and width was 30 nanoseconds.Amplifier is the AC coupling, to eliminate drifting problem.This AC coupling low-frequency cut-off frequency is higher, for example be 30MHz, thereby this pulse is distinguished.Thereby when pulse rate changes, no longer need the DC restoring circuit to keep the baseline reference voltage constant.

A comparer is sent in the output of amplifier, no matter and the size of its input value of comparer all provides the output pulse of a steady state value.The reference voltage of comparer is provided with one than the high slightly value of amplifier noise output, thereby it can not triggered by noise.The amplifier loop repeats work 96 times, is a scintillator crystals job in the detector array at every turn.Comparer output pulse provides raw data for data acquisition and image reconstruction system.The test shows forme-producing system of taking pictures can be counted with the X-photon that occurs at random up to 10MHz speed.

Though represented and described a plurality of embodiment of the present invention and application for a person skilled in the art, under the prerequisite that does not break away from notion of the present invention, can also have than above-mentioned more improved procedure.So the present invention except that the restriction that is subjected to appending claims, is not subjected to other any restrictions.

Claims

1. X-photo-detector, it comprises:

An array, it lacks and partly is made up of three discrete X-photodetections, and each X-photodetection partly has a detector input surface, and described detector input surface is not to arrange linearly mutually;

It is characterized in that described each photo-detector partly comprises a scintillator part and a photo-detector part; The bright body portion of described valve is made with X-optical flare body material.

2. X-photo-detector as claimed in claim 1 is characterized in that: described X-optical flare body material is selected from the one group of material that is made of YSO, LOS, BGO and plastic scintillant.

3. X-photo-detector as claimed in claim 1, it is characterized in that, described scintillator partly comprises the scintillator output face of the scintillator input face and a second area of a first area, described scintillator output face is parallel to described scintillator input face and separated by a distance with it, wherein said detector input surface and the common expansion in described scintillator input surface.