Title of Invention

High Quality, Low Power, Low X-ray Dose, 2D and 3D Imaging for NDT and

Medical Uses

TECHNICAL FIELD

This invention generally refers to a new generation of transmission x-ray tubes which can produce high quality medical images with the x-ray tube considerably closer to patients consequently reducing space required to generate images and electrical power to produce an image; with a high melting temperature/ high heat conductivity tungsten based target material allowing for higher tube currents, smaller spot sizes and elimination of rotating anodes; with both low and high energy filtering of generated x-rays not needed to make an image reducing unneeded x-ray flux by producing a quasi-monochromatic x-ray spectrum centered around the k-lines of tungsten and other target materials generating substantially improved image quality; and with a highly efficient way to produce excellent image quality 2D and 3D images utilizing wide cone angle tomosynthesis. Priority Date: There are 2 provisional patents relating to this filing, 61/573,490 dated September 8, 2011 and Titled: Thick Copper/tungsten or Copper/molybdenum Composite X-ray Targets, and 61/626,957 dated October 6, 2011 Titled: X-ray tube with wide angle cone beam for digital tomosynthesis.

BACKGROUND OF THE INVENTION

The majority of currentX-ray tube technology was developed by Coolidge in the early 1900's. It consists of using reflection type x-ray tubes to produce very low quantum efficient x-rays which have cone angle outputs of generally less than 50 arc degrees. The x-ray spectrum is not symmetrical about the center of the output cone beam influenced by the heel effect. The small cone angle requires the patient in medical imaging to be placed far from the x-ray source further exacerbating the amount of energy and resultant x-ray tube heat generated as a result. Large amounts of filtering are used to decrease low energy, useless dose to patients but does so at the loss of useful x-rays and additional heat generation as well.

Tungsten and tungsten rhenium alloys are well-known materials used for producing x-ray tube anode targets for medical and non-destructive testing applications of x-rays, with rhenium making the tungsten more ductile and resistant to wear. Tungsten has a relatively high heat conductivity of 170 W/mK and the highest melting temperature of any metal at 3422° C.

A recent breakthrough in thick, self-filtering targets for use in transmission x-ray tubes is detailed in US Patent application 20120051496 dated 1 March 2012, which is incorporated herein in its entirety.

In most reflection tubes, the anode is suspended on ball bearings with silver powder lubrication which provide almost negligible cooling by conduction and rotated at high speeds to spread the heat across a wide surface area while the anode rotates. The focal spot temperature can reach 2,500 °C (4,530 °F) during an x-ray exposure, and the anode assembly can reach 1,000 °C (1,830 °F) following a series of exposures. In some low power medical imaging applications solid anode Coolidge x-ray tubes are used. In medical imaging tungsten is the standard target material for most imaging applications.

Transmission x-ray tubes are well known for producing useful wide x-ray cone angles of 175 degrees or more used in non-destructive testing of electronic circuits, especially ball grid circuits, circuit boards for inspection of through holes, integrated circuits, and many other non-medical applications. Target heat problems have limited the use of transmission s-ray tubes in medical imaging. Recently developed thick target transmission tubes promise wider use in medical applications, but heat dissipation issues continue to constrain their use. Even though transmission tubes promise reduced patient dose even lower dose levels are needed.

It is well known that as the tube voltage increases in a transmission tube, the amount of flux produced increases dramatically compared to similar increases in voltage for reflection tubes. However, it does so at the expense of introducing higher energy x-rays which are detrimental to the quality of the x-ray images. Thus there is a maximum tube voltage of about 100 to 120 kVp above which x- ray quality decreases with increasing tube voltage.

Recently tomosynthesis has been introduced to produce 3D images of various parts of the human body using about the same amount of total dose that a single 2d image with conventional medical imaging x-ray tubes produces. The quality of such images produced by tomosynthesis is strongly influenced by the cone angle of the x-ray exiting the Coolidge x-ray tubes. SUMMARY OF THE INVENTION The present inventions includes target materials for both transmission and reflection tubes made from the composite metal copper/tungsten. It includes applications of the x-ray tubes with such target materials in medical imaging and non-destructive testing (NDT) normal 2D projection x-ray imaging as well as 3D images utilizing tomosynthesis imaging. The x-ray tubes provide higher quality images at considerably reduced x-ray power requirements.

In another preferred embodiment transmission x-ray tubes with wide cone angles up to about 175 arc degrees are used with target materials and target thicknesses chosen to provide higher quality images and lower electrical power required when used in producing 2D and 3D tomosynthesis imaging for both medical and NDT applications.

In another preferred embodiment unique filter materials which have a k-edge between the k-alpha and k-beta energies of the primary x-ray target material as well as filter materials whose atomic number are higher than the atomic numbers of the primary target material useful primarily in medical and non-destructive test imaging. BRIEF DESCRIPTION OF DRAWINGS Figure 1 is a schematic, elevational, cross sectional representation of a transmission x-ray tube of the current invention.

Figure 2 is a schematic, elevational, cross sectional representation of a prior art reflection type x-ray tube.

Figure 3 is a schematic, elevational, cross sectional representation of a prior art rotating anode reflection x-ray tube.

Figures 4A and 4B are schematic, cross sectional representations of parts comprising a target assembly of prior art and current invention for a reflection type x-ray tube.

Figure 5A and 5B are schematic, cross sectional representations of two target assemblies for a transmission x-ray tube.

Figure 6 is a graphical representation of the spectrum from a transmission x-ray tube with a 2 micron thick tantalum target at various angles from centerline.

Figures 7A and 7B are pictorial representations of two x-ray images taken at 0 degrees and 80 degrees from centerline.

Figure 8 is a graphical representation of the spectrum from a transmission type x- ray tube with a tantalum target 25 microns thick at different angles from centerline. Figure 9 is a graphical representation of the absorption coefficients of tantalum Figure 10 is a graphical representation of the spectrum from three transmission type x-ray tubes with 50, 75 and 100 micron thick tantalum target materials.

Figure 11 is a graphical representation of the spectrum from a transmission type x-ray tube with a target of tantalum 45 microns thick.

Figure 12 is a graphical representation of the spectrum of a transmission type tube with a 100 microns tantalum target filtered with a filter of 80 microns of ytterbium. Figure 13 is a graphical representation of the spectrum of a transmission type tube with a 50 micron thulium target filtered with an Over Z filter of 80 microns of ytterbium.

Figure 14 is a graphical representation of a reflection type tube with tungsten target and with an Over Z filter of 60 microns of Rhenium.

Figure 15 is a phase diagram for copper/tungsten composite.

Figures 16A, 16B, 16C, and 16D are pictorial representations of methods to acquire tomosynthesis data with different x-ray tube distances from the object. Figure 17 is a diagram of a hemispherical shaped filter to be used with a transmission x-ray tube of the current invention.

DETAILED DESCRIPTION OF THE INVENTION

Open transmission x-ray tubes are typically used for imaging of electronic circuits as well as other high-resolution applications and may alternatively be used as the x-ray source when high multiplication factors are required of the object's image. Closed tubes are sealed with a vacuum whereas open or "pumped down" tubes have a vacuum pump continuously attached drawing a vacuum as the tube is used usually to allow for frequent replacement of tube parts which tend to fail in operation. For purposes of this invention transmission x-ray tubes include both open and closed transmission type tubes except as otherwise stated. All x-ray tubes used in x-ray imaging require a vacuum between through which electron beams travel from the cathode to the anode. Unless otherwise specified x-ray tube spectral data was taken with an Amptek Model XR-100 with a CdTe sensor 1 mm thick and 10 mils of Be filter. The sensor was placed at a distance of 1 meter from the x-ray tube and a tungsten collimator with a collimator hole of 100 μιη diameter placed in front of the sensor. Emission of electrons from the cathode of the current invention can be done by electrical or photon stimulated emission or any other method known to those skilled in the art.

For purposes of this invention electron accelerating voltages are expressed in kVp or thousands of volts and range from 10 kVp to 800 kVp. No attempt has been made to include electron accelerating voltages in excess of 1 MVp. Additionally the energy of x-ray photons is expressed in kev, kilo-electron volts.

For purposes of this invention a K-beta Filter is a filter material whose primary filtering material's k-edge energy lies between the characteristic k-alpha and k- beta energies of the primary x-ray generating material in the target. K-beta filters decrease unwanted high energy above the primary characteristic k-alpha radiation of the primary x-ray generating material in the target and low energy x-rays which do not add appreciably to the formation of an image from an x-ray target used for x-ray imaging.

For purposes of this invention an Over Z filter is a filter whose primary filtering material is selected from an element whose atomic number is higher than the atomic number of the primary x-ray generating material. Over Z filters decrease unwanted high energy x-rays above the characteristic k-alpha and k-beta of the primary x-ray generating material in the target and low energy x-rays which do not add appreciably to the formation of an image from an x-ray target used for x- ray imaging. Over Z filters tend to produce x-rays of overall higher energy than K-beta filters. The transmission tube of the current invention 6 of Fig. 1 is comprised of an evacuated housing 10, and end- window with an anode disposed at the end of the housing exposed to atmosphere. An x-ray target foil 2 is deposited onto the end- window also called the substrate 1. In some embodiments of the current invention the end-window and the x-ray target are made of the same material, eliminating the need for a separate end-window material through which x-rays pass. A cathode 3 emits electrons, which are accelerated along the electron beam path 4 and strike the anode target 2 producing x-rays 8. A power supply 7 is connected between the cathode and anode to generate electrons and provide the accelerating energy for the electron beam. X-rays 8 produced exit the x-ray tube through the end- window in a cone angle of output x-rays. The thickness of the target material 2 may be tailored to specific applications. An optional focusing mechanism 5 typically electrically biased, focuses the electron beam above, below or onto a spot on the target. Although a cup like focusing mechanism 5 is shown for simplicity there are many sophisticated focusing mechanisms well known to those skilled in the art used to focus the electron beam to very small diameter beams as small as fractions of a micron. The largest dimension of the spot on the surface of the target is referred to as the focal spot size or spot size. The output x-rays contain bremsstrahlung and characteristic line radiation unique to the target material. Transmission type x-ray tubes in strict contrast to reflection x-ray tubes the focusing cup 5 and the electron beam 4 of the transmission type x-ray tube are symmetrical with respect to the center of axis of the x-ray tube, causing the electron beam to impinge on the end window in a round spot with a very even distribution of electrons over the surface of the focal spot, producing superior quality images especially desirable for medical imaging. The x-rays produced have a very symmetrical distribution around the centerline of the x-ray tube as a result. Fig. 2 schematically represents a reflection type x-ray tube comprised of an evacuated housing in which the cathode 3 and anode 11 are located. The anode 11 is comprised of an x-ray target deposited onto a substrate which substrate removes heat generated when x-rays impinge the anode. A power supply 7 is connected between the cathode 3 and the anode 11 to provide an electric field which accelerates the electrons from the cathode along an electron beam path 4 and strikes the anode 11 in a spot generating a beam of x-rays Item 8 which then exit the tube through a side window 12. The "heel effect" is one of the main reasons that the cone beam angle of the reflection tube is constrained below about 50 arc degrees. The tube "heel effect" well known to those skilled in the art results in as much as 20% more x-ray photons at the cathode side of the x-ray beam and 25% fewer photons on the anode side of the x-ray tube. In addition the shape of the spectrum of the photons at the cathode side is significantly different than the shape at the anode side. Heel effect is eliminated with the use of transmission type x-ray tubes. Two basic construction principles exist for reflection type x-ray tubes: stationary anodes and rotating anodes both using tungsten or tungsten rhenium targets in most applications. Stationary anodes used in medical imaging are found mostly in dental equipment, mobile C-arm units for fluoroscopy and low-load radiography which applications do not require high tube currents. Fig. 3 is a graphical representation of a reflection type of x-ray tube with a rotating anode. The anode 11 is suspended on ball bearing with lubrication and rotated to spread the heat evenly over the surface of the anode. Because the cone beam output 8 is usually less than 50 arc degrees about centerline, images of the human chest or abdomen require x-ray tube to be places far from the patient significantly increasing power 7 needed to operate the tube. A stream of electrons is emitted from the cathode 3 forming an electron beam 4 which impinges the anode 11 generating x-rays 8 that exit the tube through a side window 12 substantially transparent to x-rays.

Copper/tungsten composites are pseudo-alloys of copper and tungsten. As copper and tungsten are not mutually soluble, the composite material is composed of distinct particles of one metal dispersed in a matrix of the other. The microstructure is therefore rather a metal matrix composite than a true alloy. Copper tungsten composites are readily available as standard products from a number of suppliers in the market to include Sumitomo Electric of Japan, and Eagle Alloys of Talbot, Tn. as well as many others.

In one embodiment of the current invention the dual material stationary reflection type anode Fig. 4B made of tungsten 11 mounted onto a copper block 21 is replaced with a single member made from copper/tungsten composite Fig. 4A. Alternatively only the tungsten portion of the anode is replaced and the copper block remains. The phase diagram of copper/tungsten composite is shown Fig. 15. The melting point is in excess of 5300° C compared to only 3422° C for tungsten and 1084° C for copper. The horizontal axis represents the percent by weight of copper and tungsten and the vertical axis is in degrees C. Its heat conductivity is considerably higher than tungsten alone and depends on the percent of copper (See Table 1 below). In one preferred embodiment the percentage of copper is above 10% and less than 70% by weight. Typically between 40 and 50% provides the highest heat conductivity. Coefficie nt of Thermal

Ultimate Thermal Expansi

Composition Tensile Conducti on (@

Composition % by Strength Density vity 20C) % by Weight Volume PSI g/cm3 W/mK ppm/C

50 W; 50 Cu 32 W; 68 Cu 12.02 -275 12.2

55 W; 45 Cu 36 W; 64 Cu 63,000 251 12.5 -260 11.7

68 W; 32 Cu 50 W; 50 Cu 75,000 219 13.93 -240 9.7

70 W; 30 Cu 52 W; 48 Cu 85,000 214 14.18 -235 9.4

75 W; 25 Cu 58 W; 42 Cu 90,000 201 14.84 220 8.5

80 W; 20 Cu 65 W; 35 Cu 15.56 195 7.6

87 W; 13 Cu 75 W; 25 Cu 16.7 -170 6.4

90 W; 10 Cu 80 W; 20 Cu 17.23 160 6.1

Table 1

The disc of a rotating anode x-ray tube usually has a tungsten rhenium target area as rhenium prevents crazing of the anode surface where the electrons hit the target. The tungsten/rhenium is faced onto a molybdenum disc as molybdenum has twice the heat capacity of tungsten. In another embodiment of the current invention the rotating anode material is changed from tungsten/rhenium to copper/tungsten. In some applications where the heat capacity of copper/tungsten is sufficient to prevent melting of the anode material, the molybdenum portion of the rotating anode can be replaced as well.

Example 1:

Amsys CFX heat simulation software was used to compare the heat dissipation of a target of the current invention of Fig. 5A with the current state of the art of attaching a target of tungsten to a beryllium window of a transmission tube. A target thickness of 100 microns of tungsten was attached to 1 mm of beryllium substrate 2. This was compared to a target wherein the beryllium/tungsten assembly was replaced by a 200 micron thick copper/tungsten disc of the same dimensions with a 50% copper and 50% tungsten composition by weight using heat conductivity parameters from Table 1. It was assumed that the electron beam would impinge both targets in a spot size of 100 microns with beams of two different wattages (40 and 80 Watts). The amount of tungsten in each target was the same, producing substantially equal amounts of bremsstrahlung radiation and similar spectrum, except that the output of the copper/tungsten composite target would filter the output through an equivalent 100 micron copper filter, further reducing dose in medical imaging applications and providing some amount of characteristic k-lines of copper. Beryllium is essentially transparent to x-rays generated in the tungsten target.

Table 2

Doubling the wattage from 40 to 80 watts increases the simulated beryllium temperature by 54%, the tungsten temperature by 76%, and the copper/tungsten composite temperature by 79%. Using straight line interpolation the beryllium would melt at 63 watts of power, the tungsten at 71 watts of power and the copper/tungsten at 183 watts substantially 3 times compared to the beryllium/tungsten combined target. If the configuration of Fig. 5B were used, there would be significantly higher amounts of tube current allowed even compared to the configuration in Fig. 5A.

There are two main configurations of permanently installed medical imaging systems. One class commonly utilizes a radiolucent patient examination table with an under-table mounted tube and an imaging system mounted over the table, while the other is commonly referred to as a C-arm system used where greater flexibility in the examination process is needed such as neurological or cardiac imaging.

The non-C-arm based systems are used in most X-ray departments as 'screening rooms' typically for barium studies, endoscopy studies, chest x-rays and fertility studies. The C-arm systems are commonly used for studies requiring the maximum positional flexibility such as angiography studies, therapeutic, cardiac studies, and orthopedic procedures (ORIF, DHS, MUA, spinal work). For head imaging panoramic and CT type systems are used to image the head for dental implants and restorations, oral and maxillofacial surgery, TME and sinuses, and orthodontics. The high heat dissipation and high melting temperature of the copper/tungsten target increases the allowable heat density to increase on the target allowing smaller spot sizes and/or higher tube current. These are both major benefits to using the copper/tungsten target material to replace current anode materials for both fixed and rotating anode reflection type x-ray tubes in medical applications such as to C-arm based x-ray systems, to x-ray equipment used in screening rooms and to x-ray equipment used for imaging the head for dental implants and restorations, oral and maxillofacial surgery, TME and sinuses, and orthodontics.

Fig. 5A is a schematic drawing for one embodiment of a target assembly for a transmission x-ray tube of the current invention. The target 2 is typically made by sputtering or diffusion bonding a target material which can be as thin as sub- micron and as thick as many hundred microns typically onto substrate material. The end- window substrate is typically made of low Z material chosen from one of beryllium, aluminum, copper, lithium, boron and alloys, composites, compounds or intermetallic compounds thereof, but there are alternative end-window substrate materials well known to those skilled in the art. Tube failure is often caused by melting beryllium which has a melting temperature of 1287° C. When copper/tungsten target material is used the material is typically rolled to the desired target thickness, cut and brazed directly to the bracket 13 which can be any material commonly used in high vacuum applications. The bracket assembly is in turn typically welded to a ceramic or glass tube which can have a metal bonded to it, typically kovar or a kovar alloy used for making hermetic seals with the harder Pyrex glasses and ceramic materials. Also a very thin disc of copper/tungsten composite can be attached to a low Z substrate similarly to other target materials. This process allows for thin targets from less than 10 microns to about 200 microns. As was shown in Example 1, the heat load that can impinge the surface of the copper/tungsten target is about 3 times that for a similar tungsten target attached to a beryllium substrate.

In another embodiment of the current invention Fig. 5B the target assembly 2 of Fig. 5A is replaced by a single piece of copper/tungsten composite 9 wherein the thickness of the target at the spot where the electron beam 4 impinges the anode is machined into copper/tungsten material 9 which is in turn brazed directly to the end- window bracket 13. This configuration is less expensive and the brazing joint to the end- window bracket 13 can be further from point where electrons impinge the target increasing heat dissipation and lowering the temperature at the brazing joint. Thickness of the machined copper/tungsten target 9 where electrons impinge the target can be on the order of about 150 to 700 microns. Other metal removal processes well known to those skilled in the art can be used to make the smaller dimensions down to a thickness of 50 microns thick. A circulating fluid heat removal system can be easily added directly to the copper/tungsten airside surface or to the airside surface of the substrate material, usually beryllium. The heat removal fluid includes but is not limited to water, propylene glycol, ethylene glycol, and mixtures thereof. Direct turbulent impingement cooling recently developed for cooling integrated circuits is also well suited to cooling the end- window. There are many potential heat removal designs well known by anyone skilled in the art.

One of the major uses of the transmission x-ray tube of the current invention utilizing copper/tungsten composite as the target material generating the x-rays is in non-destructive imaging of ball grid array circuits, circuit boards, discrete electronic components, integrated circuits, micro-electro-mechanical systems (MEMS) devices, small animals, organic samples, semiconductor chip packing, baggage inspection and geological samples. The high heat dissipation and melting temperature of the copper/tungsten composite target material allows for higher tube currents with the same focal spot size or smaller spot sizes for the same tube currents, increasing the useful applications of such transmission x-ray tubes. The wide cone angle and the thin end window allows the object to be brought very close to the place where the electron beam impinges on the target, allowing for highly magnified images within a very small space.

Another major use for x-rays produced from a target of copper/tungsten composite by a transmission x-ray tube is in medical imaging. Any unwanted characteristic radiation of copper from the copper/tungsten target is at a very low energy of about 8 kev and can be easily filtered out by any of a number of possible filter materials well known by those skilled in the art without seriously decreasing the higher k-alpha energies of tungsten needed for medical imaging. The wide cone angle of the reflection type tube allows the x-ray tube to be brought closer to the patients reducing tube wattage. The significantly higher quality of the x-ray beam provides considerably higher image quality when compared to state of the art reflection x-ray tubes. The thick target of the transmission x-ray tube acts as a very effective filter of low energy x-rays which do not produce useful x-rays but do contribute to the dose a patient receives. As the voltage of the transmission type x-ray tube is increased there is a major increase in the x-ray flux produced, whereas the increase of flux for the same increase in voltage for a reflection type tube is minimal Thick targets copper/tungsten are especially useful in medical imaging of the head, breast, chest, abdomen, GI series, joints and extremities. Example 2:

A test was conducted comparing reflection type fixed anode tungsten target (non- rotating) to a transmission x-ray tube using a 25 micron thick tantalum target, wherein both tubes were filtered by an aluminum equivalent 9mm filter. All other parameters were identical. The amount of useful flux for imaging purposes was arbitrarily set at the total flux between 40 kev and 65 kev produced by each x-ray tube. The reflection tube used 3mA tube current to produce x-ray flux at each voltage and the transmission tube required only 1.91 mA at 80k Vp, 1.69mA at 90kVp, 1.52mA at lOOkVp, and 1.30mA at l lOkVp, 1.12mA at 120kVp compared to 3 mA for the reflection type tube, demonstrating an advantage of the transmission tube . The higher the tube voltage the more efficient the x-ray generation for a transmission type x-ray tube.. Table 3 below compares the amount of unwanted flux from the same test arbitrarily set at above 90 kev and below 40 kev for both the reflection type tube and the 25 micron tantalulm target transmission tube. The data clearly shows as the tube voltage increases the transmission tube performs considerably better than the reflection tube. This was accomplished with a target thickness of only 25 microns. Thicker transmission targets produce even better results however with an increase in tube current to compensate for a small loss of x-rays in the useful x- ray range.

Table 3

Example 3:

Data taken using a transmission x-ray tube in one embodiment of the current invention is detailed in Fig. 6 illustrating the wide cone angle of x-rays from a transmission tube and how the strength of the x-ray signal changes from centerline to 80 arc degrees from centerline. An x-ray tube with a tantalum target foil 2 microns thick sputtered onto 1 mm of beryllium end- window was used to measure output spectrum with a tube voltage of 60 kVp and a focal spot size of less than 50 microns useful in non-destructive testing applications. Tube current was fixed at 50 microamperes. Measurements were made by a spectrometer from 0 to 80 degrees from centerline. Superimposed spectrum data from 0 to 50 degrees 17 from centerline, 60 degrees 18, 70 degrees 19, and 80 degrees 20 from centerline respectively are shown. Table 4 below summarizes the results for useful x-rays of energies between 45 and 55 kev every increasing 10 arc degrees from centerline.

Table 4

The visual effects of the 25% decrease in energy at 80 arc degrees are demonstrated in images of a ball grid array circuit in Fig. 7A 0 degrees and Fig. 7B 80 degrees. Images were taken at equal distances from the x-ray tube. This shows clearly that the transmission x-ray tube of the current invention has a useful cone angle of more than 160 arc degrees and even higher than 170 arc degrees, compared to a reflection tube of less than 50 arc degrees. In addition the spectrum from the transmission type tube is uniform in all directions as there is no preferred direction from the symmetrical design of the end window providing superior images with no heel effect. Example 4

In one embodiment of the current invention a transmission x-ray tube suitable for use in obtaining medical images with a target foil of tantalum 25 microns thick was deposited on a beryllium substrate 1 mm thick which also acted as a low energy x-ray filter as well as an end-window to pass generated x-rays outside of the tube. Fig. 8 is the graph of the output spectrum with a vertical axis representing the number of photons produced at each specific x-ray energy represented by the horizontal axis in kev. Similar to Example 3 above spectral data was taken every 10 degrees from centerline. The spectra 50 from 0 arc degrees to 40 arc degrees are virtually the same. The spectrum 49 at 50 degrees from centerline shows some minimum decrease in flux as well as the spectrum 48 at 60 degrees from centerline. At lower kev energies there is significant difference in the spectrum, but for medical imaging useful energies are generally as high as or higher than 55 kev. It can be seen that the amount of x-ray flux at the k-alpha of tantalum is only minimally decreased, allowing the tube to be used for medical imaging up to 60 arc degrees from centerline and even larger angles up to as much as 70 or 80 arc degrees from centerline.

With a reflection type tube the "heel effect" well known to those skilled in the art results in typically 20% more x-ray photons at the cathode side of the x-ray beam and 25% fewer photons on the anode side of the x-ray tube at centerline or a 45% difference. With the transmission tube the maximum difference in k-oc between centerline and 60 arc degrees 48 from center is 33% and for k-β is 32%. This demonstrates that a transmission x-ray tube of the current invention with a 25 micron thick target has an effective cone angle of at least 120 degrees. Thinner target materials will result in even wider cone angles of as much as 175 degrees. When thicker target material is used the cone angle for the transmission x-ray tube can be as small as about 50 arc degrees. It is possible to construct any filters so that they form a hemisphere centered at the output of the transmission x-ray tube whereby the thickness of the filter at each angle from centerline can be controlled.

Table 5

Examples 7 and 8 below demonstrate the effect of K-beta and Over Z filters with a transmission type x-ray tube. Wide angle imaging of the transmission x-ray tube of the current invention allows the x-ray source to be brought closer to the object to be imaged. In one embodiment chest x-rays which are normally taken with the x-ray source to imaging sensor of 180 cm can be taken at a distance of 50 cm or less from the imaging sensor, decreasing the amount of x-radiation needed by a factor of about 13. The approximate 3 fold increase in the number of useful x-rays produced by the transmission x-ray tube compared to a reflection x-ray tube at 120kVp tube voltage reduces the power requirements even further by a factor of a little less than 3. When copper/tungsten composite target material is used there is another 3 fold increase in the amount of heat the transmission tube can withstand. Finally when transmission tube applied voltages are increased from 90 or 140 kVp, the characteristic k line radiation output of the transmission tube increases some 6 fold with an electrical power (heat) increase of 55%. Some of the increase in k-line radiation can be sacrificed to provide either lower dosage to the patient, higher image quality by filtering the output x-rays from the copper/tungsten composite target with filter materials chosen from lutetium, ytterbium, thulium and hafnium or Over Z filtering materials or both. Over Z filtering materials include all materials with a Z number of 75 (rhenium) or greater. Doubling of the target thickness of a transmission x-ray tube demonstrates a decrease in the amount of x-radiation below 40 kev (considered to be dose with no imaging content) delivered to a patient by a factor of 55% which can be further decreased by lutetium, ytterbium* thulium and hafnium or Over Z filtering to a level close to 10% of the current below 40kev dose level from corresponding reflection type x-ray tubes using low Z filter materials such as copper or aluminum.

Fig. 9 is a plot of the absorption coefficient ρ/μ in cm ² /gm (vertical axis) against the photon energy in Mev (horizontal axis) of the target element tantalum. A circle is drawn around the tantalum target characteristics 22 where k-line emissions and the k-edge of tantalum occur. The k-edge is 67.4 lkev is the energy at which the absorption coefficient jumps from 2.315 cm /gm to 11.80 cm /gm. It is an objective of the current invention to decrease higher energy x-rays above k- edge while not appreciably decreasing k-alpha radiation. Each target material will have a similar plot of absorption coefficient where in the k-edge increases as the atomic number of the filter element increases.

Unlike prior-generation imaging systems for use in non-destructive testing and medical imaging which generate two-dimensional images, tomosynthesis produces three-dimensional images which are intended to reveal the inner architecture of the imaged object or body part, free from the distortion typically caused by shadowing or density differences inside the object or body part. Tomosynthesis improves upon conventional geometric tomography in that it allows an arbitrary number of in-focus planes to be generated retrospectively from a sequence of projection radiographs that are acquired during a single motion of the x-ray tube. By shifting and adding these projection radiographs, specific planes may be reconstructed. It has been shown that by using tomosyntheis imaging in mammography that the radiation dose to produce a single 2D image is about the same as the total radiation dose to produce a 3D tomosynthesis image. For mammography typically 11 to 20 single exposures are summed to provide tomosynthesis imaging. For the purposes describing the scope of this invention tomosynthesis images include 3D images, 2 D images and any other combination images which include information from both 3D and 2D images in the same image. This definition refers to all images obtained using tomosynthesis to include but not limited to medical and non-destructive testing images.

Recently there are many image sensor technologies being developed, well known by those skilled in the art, with particular application to tomosynthesis. CMOS sensor-based imaging allows for high-resolution, high dynamic range and low noise. Amorphous silicon technologies with high-performance thin film transistors are capable of amplifying the sensor value with a user controllable gain over a wide input range for high resolution, low noise x-ray tomosynthesis applications. Active pixel sensor architectures, in contrast to the industry standard passive pixel sensor architecture, have the capability to provide high gain and low noise imaging that can potentially reduce patient delivered dose for tomosynthesis.

Figs. 16A, 16B, 16C, and 16D illustrate how a transmission x-ray tube of the current invention can be used in tomosynthesis imaging. The x-ray tube of the current invention 39 is moved in either a linear or curved path over the patient or object 40 placed in close proximity to a digital imager 38. When the x-ray tube is brought closer to the patient there is a need to use a somewhat wider digital sensor to collect all of the images. Alternatively the sensor can be moved at the same time as the x-ray tube. There is also a distortion of distance between two different pixels at the center of the image compared to the edge. However, that distortion is a geometrical function of the angle of incidence of the x-rays to the pixel and can be easily removed in software processing. Although only 5 tube locations 41-45, called steps, are shown in Fig. 16 A, typically for a single tomosynthesis image as few as about 5 and as many as 20 or even 40 steps may be required to allow for high quality images to be reconstructed. Figs. 16B, 16C, and 16D illustrate the way the cone angle of the x-ray tube must be increased as the x-ray tube is brought closer to the patient or object to be imaged. The target materials and target thickness and any subsequent filtering of the transmission x-ray tube are chosen to provide the widest possible cone angle for the particular object to be imaged. For illustration purposes the number of pixels at which data is collected in Figs. 16 B, 16C and 16D is only nine for ease of illustration. There would be actually thousands of such pixels depending on the pixel density of the imager plate. In another preferred embodiment of the current invention the focusing cup 5 of a transmission x-ray tube of Fig. 1 of the current invention can be used as a high speed switch to turn on and off the electron beam 4 very rapidly by switching the bias of the cup from normal operation to highly negative, stopping the flow of electrons, and then to normal again. Alternatively an additional grid may be added between the focusing cup 5 and the anode target 2 and that grid switched from highly negative to zero to turn-on and off the flow of electrons in very rapid fashion. In tomosynthesis applications the process of obtaining as many as 40 images or more over the path of movement of the x-ray tube must be done in a very short time on the order of seconds. Typically each step requires between 200 microseconds to as much as 1.5 seconds to obtain the x-ray image for medical images and somewhat longer for NDE images.

Researchers from the University of Wisconsin-Madison have developed a method of image reconstruction applicable to image processing in tomosynthesis imaging using total variation minimization (Limited View Angle Tomographic Image Reconstruction Via Total Variation Minimization, Julia Velikina, Shuai Leng, and Guang-Hong Chen, Department of Medical Physics,University of Wisconsin- Madison, Madison, WI 53704; Department of Radiology, University of Wisconsin-Madison, Madison, WI 53792). For a given limited range of projection angles, the quality of the reconstructed images changes with the number of projections. Images produced with such small projection angles of 30 degrees are of marginal quality and there is no significant image quality change when the number of projections increases from 20 to 40, 60, and 100. When the view angle range is larger for example 150 degrees the imaged object is reconstructed exactly using only 10 or 20 projections. When the view angle range is 120 degrees, high contrast can be reconstructed from 30 projections. Using reflection type x-ray tubes, the view angle range of acquired projections is usually less than about 50 degrees. In another embodiment the transmission type of x-ray tube of the current invention provides useful high quality x-rays for use in both medical and NDT tomosynthesis imaging. See Examples 3 and 4 above for an explanation of cone angles for transmission x-ray tubes. Useful emitted x-ray cone angles are greater than 50 degrees and as high as 175 degrees. In tomosynthesis imaging the primary x-ray target materials will be typically chosen from pure elements, alloys, composites, or intermetallic compounds of at least one of Vanadium, Scandium, Titanium, Chromium, Iron, Cobalt, Nickel, Copper, Zinc, Yttrium, Zirconium, Niobium, Molybdenum, Ruthenium, Rhodium, Palladium, Silver, Tin, Antimony, Tellurium, Lanthanum, Terbium, Gadolinium, Dysprosium, Holmium, Erbium, Thulium, Ytterbium Lutetium, Hafnium, Tantalum, Tungsten, Rhenium, Iridium, Gold or Uranium wherein one of the composite materials includes copper/tungsten with a ratio of the percentage of copper by weight chosen to be greater than about 10% and less than 70%. The composition of copper is chosen to be preferably between 40 and 50% by weight however lower or higher percentages will work also. When copper/tungsten composite is chosen, target thickness of can be as thin as 10 microns and as thick as 700 microns. At those thicknesses high quality medical x-ray images can be produced when applied electron accelerating voltages are as low as 25 kVp and as high as 150 kVp and high quality NDT images can be produced when applied electron accelerating voltages are as low as 20 kVp and as high as 200 kVp chosen for each imaging application

Tomosynthesis medical imaging with transmission x-ray tubes of the current invention are particularly useful for applications such as breast, chest, joints and extremeties, head, abdomen, GI series, and to guide high energy radiation therapy to the precise location in the patient where such therapy is to be applied. Head imaging is used for dental implants and restorations, oral or maxillofacial surgery, TME and sinuses. Typically medical imaging requires spot sizes in the range of 100 microns to 2 mm.

In another embodiment of the current invention the transmission x-ray tube is a "closed" tube used to obtain tomosynthesis medical images of the head, chest, abdomen, GI series, joints or extremities or non-destructive testing (NDT) images of circuit boards, discrete electronic components, micro-electro-mechanical systems (MEMS) devices, small animals, organic, geological samples and many other inanimate objects used in various industries. The target material is a copper/tungsten composite with a target thickness of greater than 10 microns and as thick as 300 microns wherein the percent of copper in the composite is between 10% and 70%. Thinner targets provide a wider useful x-ray cone beam angle. The applied electron accelerating voltages can range from as low as about 20 kVp to as high as 200 kVp providing useful cone angle of emitted x-rays from 50 to 150 degrees. Generally the thinner the target the wider the useful x-ray cone angle. A target thickness of copper/tungsten composite target with 50% copper by weight is an equivalent thickness of about 5 microns of tungsten. Such a thin target can provide useful cone angles of 150 degrees or greater depending on the applied tube accelerating voltage. Spot sizes vary typically from about 0.5 microns or less up to 1.0 mm or even 2.0 mm depending on specific application. When a thin copper/tungsten target is attached to a substrate substantially transparent to x-rays such as a beryllium substrate of 0.5 mm to3.0 mm in thickness, the object to be imaged can be brought very close to the place on the target where the spot is generated particularly useful in NDT imaging. The speed of imaging, the allowable spot size, and the magnification of the image can all be improved.

There are many mathematical algorithms used to solve the problem of reconstructing meaningful images from the raw data obtained in tomosynthesis image acquisition, including but not limited to back-projection algorithms, shift and add algorithms, iterative Maximum Likelihood algorithms, and many more well known by those skilled in the art.

In another embodiment of the current invention for use in mammography a transmission x-ray tube has a target material comprised at least partially of molybdenum. The thickness of the target material can be between 25 microns and 200 microns. Single or multiple filament Over Z materials whose k-edge lies between 29 kev and 80 kev can be used to take advantage of the increase in x-ray flux from high beam energies and still produce imaging x-rays from the k-alpha lines of molybdenum. The increased x-ray flux output from a transmission x-ray tube with a thick molybdenum target when the applied tube voltage increases from 29 kVp and 80 kVp is remarkably high eliminating the need for rotating anodes in mammography applications. The variable applied voltage and the filtering system are adjusted until maximum image quality is obtained. In one embodiment of a transmission x-ray tube of the current invention Fig. 17 represents one way to make any filter system wherein the filter 48 is shaped in a hemisphere centered at the spot where the electron beam 4 impinges the anode 11. The filter can be of a single filter material and thickness, or of a number of different filters and thicknesses. The first filter material 48 can be shaped by any metal forming technology into the filter thickness desired. Additional filters can be added by sputtering directly to the surface of the first filter material either on the inside or the outside of the hemisphere. Sputtering thickness depends on the desired thickness of each filter material used in add-on filters. In this way the amount of filtering the x-ray beam sees at centerline of the tube is about the same mount at any angle from centerline in the output beam of the x-ray tube. This insures that the effect of the filter thickness will vary appreciably over all angles of the output beam of x-rays. Although a curved surface for the filter is preferred any filter geometry is acceptable.

When x-rays are required of infants or adolescent patients, because the size of their body is considerably smaller than adults, a transmission x-ray tube made from target chosen from materials with low Z numbers can be chosen to produce soft quasi-monochromatic x-rays ideal to producing high quality images standard medical images as well as tomosynthesis images with minimal dose to the patient. K-beta and/or Over Z filtering can be used to further decrease low end photons below about 20 to 25 kev while filtering out unwanted high energy photons as well. . Example 5

Fig. 10 is a comparison of the spectrum from 3 transmission type x-ray tubes of the current invention with tantalum target foils of 50 microns thick 25, 75 microns thick 24 and 100 microns thick 23 attached to a 1mm beryllium substrate and used at a tube voltage of 120 kVp. The thicker the target the more monochromatic are the spectrum. The vertical axis represents the number of photons produced at the x-ray energy in kev shown on the horizontal axis. It is clear that as the target thickness is increased there is significant reduction in the unwanted dose of x-ray energies between 0 and 40 kev. However there is very little reduction in the k- alpha energy. The decrease in k-alpha from a 50 micron thick target 25 to 100 microns 23 showed a small loss indicating that there was additional 23% k-alpha generated by absorption of the 100 Ta target at energies above k-edge of tantalum.

Example 6

The horizontal axis of Fig. 11 represents measured photon energies in kev and the vertical axis the number of photons measured at each energy at centerline. A transmission x-ray tube with a 45 micron thick tantalum target attached to a copper substrate 300 microns thick was tested at gradually increasing voltages from 90kVp 26, lOOkVp 27, HOkVp 28, 120kVp 29, 130kVp 30 and 140kVp 31. Table 3 shows increases in total x-rays produced of 7.5 fold and produced with energies between 55 and 60kev, characteristic k-line emissions from tantalum, of 6 fold. The increase in x-ray output is remarkably different from that of reflection x-ray tubes as is well understood by those skilled in the art.

Table 6

Example 7

Fig. 12 illustrates the effect of an ytterbium K-beta filter 80 microns thick on the output of a transmission x-ray tube with a 100 micron thick tantalum target. The vertical axis represents the number of photons output at each photon energy band in kev of the horizontal axis. A tube voltage of 120 kVp was used to produce the spectrum 32 without filtering. The output spectrum after filtering with the k-beta ytterbium filter 33 is compared to the tube output with no filtering 32. Table 7 illustrates changes in counts in energy bands <40 kev, 55-60 kev, and greater than 70 kev with an without the filter. Any reduction of useful k-alpha radiation can be compensated by increasing tube current. It is obvious that there is a considerable decrease in low energy photons <40kev and high energy photons >70 kev when the k-beta filter is applied.

Table 7

Example 8

An Over Z ytterbium filter 80 microns thick was used with a 50 micron thick thulium target of a transmission x-ray tube of this invention to calculate the effect of the filter at 140 kVp of tube voltage. Fig. 13 compares the two spectrum before 34 and after 35 Over Z filtering. The horizontal axis represents 5 kev energy bands of photos in kev and the vertical axis the number of photons in each band. The results are compiled in Table 8 below. There was a decrease in number of photons in the k-line emission range of 50-55 kev of 21.4% however a larger decrease in energy above 70kev of 30.6% and a sizable decrease in dose below 40 kev of 74.2% demonstrating the effectiveness of Over Z filtering. It should be noted that in actual operation a significant portion of the energy absorbed by the filter above the k-edge of ytterbium (61.3 kev) will generate additional k-alpha and k-beta energies of ytterbium between 50 and 55 kev. Thus the actual reduction of counts between 50 and 55 kev will be considerably less than the 21.4% calculated. Target Thickness Counts <40 kev Counts 50-55 kev Counts >70 kev

50 Th no Filter 6219 4815 5595

50 Th 80μιη Yb Filter 1609 (-74.2%) 3788 (-21.4%) 3883 (-30.6%)

Table 8

Example 9

Fig. 14 shows the difference between the spectra of a typical reflection type x-ray tube 36 with a low Z filter of 9 millimeters of aluminum equivalent used in medical imaging with an x-ray target whose material is predominantly tungsten, Z=74 operated at a tube voltage of 120 kVp. Output spectra 37 from the same tube is filtered through an Over Z filter material of 60 micron thick rhenium, Z=75 on lieu of the 9 mm aluminum equivalent filter. Clearly the output of the Over Z filter provides a significant reduction in dose of unused low energy x-ray photons of energies below about 35 kev and a similar decrease in the output from about 85 kev to 110 kev which reduces x-ray hardening and loss of image quality. The overall effect is a more quasi-monochromatic x-ray spectrum for improved image quality. When the high energy photons from 85 to 110 kev are absorbed by the filter a portion as high as about 25% depending on the filter thickness then transforms the absorbed photons into k-alpha radiation of rhenium which is 59.7 kev compared to the k-alpha of the tungsten target which is 57.9 kev and adds to the overall characteristic k-line flux important to obtain not only medical images but NDT images as well.