CN109698105B

CN109698105B - High dose delivery, transmission and reflection target X-ray system and method of use

Info

Publication number: CN109698105B
Application number: CN201811234524.1A
Authority: CN
Inventors: 菲利浦·肯特·奥斯本
Original assignee: Rad Source Technologies Inc
Current assignee: Rad Source Technologies Inc
Priority date: 2017-10-23
Filing date: 2018-10-23
Publication date: 2021-08-10
Anticipated expiration: 2038-10-23
Also published as: CN109698105A

Abstract

A high dose delivery, transmissive and reflective targeted X-ray tube and method of use generally includes an X-ray tube for accelerating electrons at high voltage potentials having: an evacuated high pressure housing; a hemispherical transmissive and reflective target anode disposed in the housing; a cathode structure for deflecting electrons towards a hemispherical anode disposed in the housing; a filament located at the geometric center of an anode hemisphere disposed in the housing; a power supply connected to the cathode to provide an accelerating voltage to the electrons.

Description

High dose delivery, transmission and reflection target X-ray system and method of use

Cross Reference to Related Applications

To the full extent allowed by law, the present U.S. non-provisional patent application is the continuation of the U.S. non-provisional patent application entitled "High Dose Output, Through Transmission Target X-Ray System and Methods of Use" filed on 31/12/2014 and assigned serial number 14/587,634, hereby claiming priority and full benefit.

Technical Field

The present disclosure relates generally to X-ray tube technology and more particularly to X-ray tubes having particular anode, cathode, filament configurations and material selection to produce high dose X-ray output.

Background

In many typical states of the prior art X-ray tubes, the cathode assembly and the anode assembly are vacuum sealed in a glass or metal enclosure (envelope). Electrons are generated by at least one cathode filament in the cathode assembly. These electrons are accelerated towards the anode assembly by a high voltage electric field. The energetic electrons generate X-rays upon impact with the anode assembly. A by-product of this process is the generation of a large amount of heat.

Conventional X-ray tube configurations are known in the art, for example, the Coolidge type X-ray tube. In the Coolidge tube, the X-ray photons, shown as spot output radiation patterns, are generated by impinging an electron beam emitted from a filament onto the surface of a target anode. The Coolidge tube can be single-ended in which the cathode is at a negative potential and the anode is at ground, or double-ended in which the cathode is at a negative potential and the anode is at a positive potential. In either configuration, the energy of the acceleration is the difference between the electrode potentials. In the Coolidge X-ray tube, the target anode is made of a heavy metal such as tungsten, tantalum or iridium, and this material is chosen due to its density and high melting point. The material of the target anode is most often mounted on a thermally conductive material (such as copper) and externally cooled by water or dielectric oil.

The target anode is placed in line with the electron beam and the radiation is emitted at right angles to the electron beam. The spectrum of the output radiation is predominantly bremsstrahlung and is varied by varying the acceleration energy of the electron beam. Tubes of this nature are used for industrial imaging, medical imaging, analysis, and radiation applications. The main limitations of this type of tube are the watt density loading of the target anode before melting occurs, the limited utilization of the generated X-ray photons and the symmetry of the resulting radiation field. Since the resolution of the imaging device (electronics or thin film) is a function of the size of the electron beam projected onto the target anode. For optimal image resolution a small focus is required, but for optimal image contrast a large number of X-ray photons is required. These two requirements are opposite and cannot be addressed in conventional tube designs. In addition, the reflective properties of the emitted radiation are asymmetric about the beam centerline and are very inefficient for X-ray radiation applications.

Recently, some low power transmission (through transmission) X-ray tubes have been marketed. These pipes use a single element (element) as a combined target and output window. The most commonly used element is tungsten because of its higher melting point, but at the expense of reduced radiation output.

It is therefore apparent that there is a recognizable unmet need for a high dose output, a transmission transport target X-ray system and method of use with a large surface area anode target to dissipate heat and thus achieve a higher atomic number target material, with improved radiation output, lower melting point and higher evaporation pressure, and a low electrode potential required to produce higher output radiation.

Disclosure of Invention

Briefly described, in an exemplary embodiment, the present apparatus overcomes the above-mentioned disadvantages and meets the recognized need for high dose delivery, a transmission transfer target X-ray tube and method of use generally including an X-ray tube for accelerating electrons at high voltage potentials, the X-ray tube comprising: an evacuated enclosure that is sealed; a transmissive transmission target anode structure disposed on the housing, the anode structure configured as a hemisphere having a geometric center; a cathode structure disposed in the housing, the cathode configured to deflect electrons toward the hemispherical anode; a filament disposed in the housing, the filament positioned proximate the geometric center of the hemisphere and between the anode and the cathode; an evacuated envelope configured to vacuum enclose therein the anode, the cathode and the filament, and therefore such an X-ray tube serves as an anode target providing a large surface area for heat dissipation and enables the use of different z-materials to take advantage of characteristic X-rays with improved radiation output, lower melting points and lower electrode potentials required to produce higher output radiation.

In accordance with its principal aspects and broadly stated, a high dose delivery, transmission transfer target X-ray tube and method of use generally includes an X-ray tube for accelerating electrons at high voltage potentials having: evacuating the high pressure enclosure; a hemispherical transmissive transmission target anode disposed in the housing; a cathode structure deflecting electrons towards a hemispherical anode disposed in the housing; a filament located at the geometric center of an anode hemisphere disposed in the housing; a power supply connected to the cathode to provide an accelerating voltage to the electrons.

In an exemplary embodiment of a transmissive transport target X-ray tube and method of use, the X-ray tube comprises: an evacuated enclosure that is sealed; a transmissive transmission target anode structure disposed on the housing, the anode structure configured as a hemisphere having a geometric center; a cathode structure disposed in the housing, the cathode configured to deflect electrons toward the anode structure; a filament disposed in the housing, the filament positioned approximately at a geometric center of the hemisphere and between the anode and the cathode, wherein the evacuated housing is configured to vacuum seal the anode structure, the cathode structure, and the filament therein.

In another exemplary embodiment of a transmissive and reflective target X-ray tube for accelerating electrons at high voltage potentials, comprising: a housing; a transmissive and reflective transmissive target anode structure disposed on the housing, the anode structure configured as a hemisphere having a center of a circle created by the 2D substrate; a cathode structure disposed in the housing, the cathode structure configured to deflect electrons toward the anode structure; a filament disposed in the housing, the filament positioned proximate a center of a circle created by the hemispherical 2D substrate and between the anode structure and the cathode structure, wherein the evacuated housing is configured to vacuum seal the anode structure, the cathode structure, and the filament therein.

In an exemplary embodiment of a transmission target X-ray tube, a method for producing a monochromatic output X-ray spectrum includes the steps of providing an X-ray tube for accelerating electrons at a high voltage potential, the X-ray tube having: an evacuated enclosure that is sealed; a transmissive and reflective transmissive target anode structure disposed on the housing, the anode structure configured as a hemisphere having a geometric center; a cathode structure disposed in the housing, the cathode structure configured to deflect electrons toward the anode structure; a filament disposed in a housing, the filament positioned proximate a center of a circle created by the hemispherical 2D substrate and between the anode structure and the cathode structure, wherein the circle created by the hemispherical 2D substrate is in direct contact with the cathode structure, and wherein the evacuated housing is configured to vacuum seal the anode structure, the cathode structure, and the filament therein to a K below only the at least one target element_αThe output X-ray spectrum of the energy is filtered and K is adjusted to be higher than at least one target element_αThe cathode voltage of the energy.

Thus, a feature of the high dose delivery, transmission delivery target X-ray tube and method of use is its ability to generate a symmetric X-ray field.

Another feature of the high dose delivery, transmissive transmission target X-ray tube and method of use is its ability to provide a large surface area anode target to dissipate heat.

Yet another feature of the high dose delivery transmission target X-ray tube and method of use is its ability to use different z-materials to take advantage of the characteristic X-rays that would increase radiation output.

Yet another feature of the high dose delivery, transmissive delivery targeted X-ray tube and method of use is its ability to use targeted materials with lower melting points for special applications (e.g., generation of monochromatic X-rays) and therapeutic applications.

Yet another feature of the high dose delivery, transmission delivery target X-ray tube and method of use is its ability to utilize lower electrode potentials to produce higher output radiation.

Yet another feature of the high dose delivery, transmissive transmission target X-ray tube and method of use is its ability to provide a new anode configuration, which allows the use of alternative target materials with different characteristic X-rays.

Yet another feature of the high dose delivery, transmission delivery target X-ray tube and method of use is its ability to provide an X-ray tube that requires no or limited heat dissipation in the form of air cooling or liquid cooling. Furthermore, forced air cooling is more efficient as the surface area of the new anode configuration is increased.

Yet another feature of the high dose delivery, transmissive transmission target X-ray tube and method of use is its ability to provide an X-ray tube with increased lifetime due to the heat sinking capability of the large surface area anode target.

Yet another feature of the high dose delivery, transmissive transmission target X-ray tube and method of use is its ability to provide new structures and geometries for the anode to increase the surface area of the anode.

Yet another feature of the high dose delivery, transmissive transmission target X-ray tube and method of use is its ability to provide an anode configuration with better heat transfer characteristics, which will enable the anode to operate at lower temperatures and thus enable the selection of lower melting point materials and increased radiation output, and extend the useful life of the X-ray tube.

Yet another feature of the high dose delivery, transmissive transmission target X-ray tube and method of use is its ability to provide new structures and geometries for the cathode, which deflect and/or accelerate electrons towards the new structures and geometries of the anode.

Yet another feature of the high dose delivery, transmissive, transport target X-ray tube and method of use is its ability to provide new structures and geometries for the filament that releases electrons evenly distributed toward the new structure and geometry anode.

Yet another feature of the high dose delivery transmission target X-ray tube and method of use is its ability to provide a minimum anode target to the radiation sample distance, thereby creating an X-ray source that can be placed closer to the object.

Yet another feature of the high dose delivery, transmission target X-ray tube and method of use is its ability to generate X-rays for biological or organic material radiation, radiation therapy, treatment of certain diseases by killing or altering human cells, imaging (e.g., medical, industrial and dual energies, non-destructive evaluation of objects, X-ray defects, X-ray diffraction patterns, therapeutic X-rays, analytical X-rays, and X-ray microscopes).

Yet another feature of the high dose output, transmissive and reflective target X-ray tube and method of use is its ability to generate a useful amount of reflected photons opposite the forward electron travel of the target anode. This phenomenon can be observed from a Z-axis plot of the X-ray field.

Yet another feature of the high dose delivery, transmissive and reflective targeted X-ray tube and method of use, due to the hemispherical shape of the anode and corresponding cathode structure, is the ability to passively manipulate the electron field by varying the distance between the two structures.

Yet another feature of the high dose delivery, transmissive and reflective targeted X-ray tube and method of use is the ability to vary the filament size and shape to alter the electron emission characteristics.

These and other features of the high dose output, transmissive and reflective target X-ray tube and method of use will become more readily apparent to those skilled in the art from the following detailed description of the exemplary embodiments and the claims when read in light of the accompanying drawings.

Drawings

A high dose output transmission target X-ray tube and method of use will be better understood by reading the detailed description of the exemplary embodiments with reference to the drawings, in which like numerals represent like structures and refer to like elements throughout, and in which:

fig. 1 is a schematic cross-sectional representation of a prior art X-ray tube of the Coolidge type;

FIG. 2 is a schematic cross-sectional representation of an exemplary embodiment of a transmissive transmission target X-ray tube showing a cross-section of an electron trace being emitted from a cathode filament and showing a cross-section of output radiation being emitted from an anode target;

FIG. 3 is a graphical representation of X-ray energy keV versus gold target dose;

FIG. 3.1 is a graphical representation of the dose of X-ray energy keV versus a combination of materials such as tungsten, iridium and gold as targets;

FIG. 4 is a graphical representation of target anode thickness versus dose for different kV;

FIG. 5 is a pictorial representation of an exemplary application for irradiating biological material using the transmissive transmission target X-ray tube of FIG. 2;

FIG. 6 is a schematic elevational cross-sectional representation of an exemplary embodiment of the transmissive transmission target X-ray tube of FIG. 2 in combination with a monochromatic filter;

FIG. 7 is a graphical representation of X-ray energy versus dose for a transmission target X-ray tube combined with the monochromatic filter of FIG. 6;

FIG. 8 is a schematic elevational cross-sectional representation of an alternative exemplary embodiment of a transmissive transmission target X-ray tube showing a cross-section of an electron trace being emitted from a cathode filament and showing a cross-section of output radiation being emitted from an anode target;

FIG. 9 is a flow chart of a method of generating a symmetrically shaped X-ray field;

FIG. 10 is a schematic cross-sectional representation of an alternative exemplary embodiment of a transmissive and reflective target X-ray tube showing a cross-section of an electron trace being emitted from a cathode filament and showing a cross-section of output radiation being emitted from an anode target and reflected off the anode target;

FIG. 11A is a cross-sectional representation of an alternative exemplary embodiment of the transmissive and reflective target X-ray tube of FIG. 10, shown with the tube sidewalls extended;

FIG. 11B is a cross-sectional representation of an alternative exemplary embodiment of the transmissive and reflective target X-ray tube of FIG. 10, shown with shortened tube sidewalls;

FIG. 12 is a cross-sectional representation of an alternative exemplary embodiment of the transmission and reflection target X-ray tube of FIG. 10 and a graphical representation of photon intensity of photons or radiation associated with the tube centerline; and

FIG. 13 is a pictorial representation of an exemplary application for irradiating biological material using the transmissive and reflective target X-ray tube of FIG. 10.

It is noted that the drawings presented are for illustrative purposes only and, thus, they are neither intended nor intended to limit the disclosure to any or all of the precise details of construction shown, except insofar as they may be considered essential to the claimed invention.

Detailed Description

In describing exemplary embodiments of the present disclosure, specific terminology is employed for the sake of clarity as illustrated in FIGS. 1-3, 3.1, 4-10, 11A, 11B, 12-13. However, the disclosure is not intended to be limited to the specific terminology so selected, and it is to be understood that each specific element includes all technical equivalents that operate in a similar manner to accomplish a similar function. The examples set forth herein are non-limiting examples and are merely examples of other possible examples.

Referring now to fig. 1, there is shown a schematic cross-sectional representation of a Coolidge-type X-ray tube 12, as shown in fig. 1, comprising: an X-ray tube housing 1, which may be glass or metal; a high-voltage insulator 2; and a vacuum dielectric 3 contained in the X-ray tube housing 1. In the Coolidge tube, X-ray photons, shown as fan-shaped output radiation patterns 7, are generated by impinging an electron beam, shown as electron trajectory 6, emitted from a filament 5 onto the surface of a target anode 9, shown as X-ray target 9. The Coolidge tube can be single-ended in operation, with the cathode (shown as cathode assembly 4) at a negative potential and the target anode 9 at ground, or double-ended in operation, with the cathode 4 at a negative potential and the target anode 9 at a positive potential. In either configuration, the energy of the acceleration is the difference between the electrode potentials. In the Coolidge X-ray tube, the target anode 9 is made of a heavy metal such as tungsten, tantalum or iridium, and this material is chosen for its density (tungsten-19.35, tantalum-16.65 or iridium-22.4 gr/cm3 (grams per cubic centimeter)) and high melting point (tungsten-3410, tantalum-2996 or iridium-2410 degrees Celsius (C)). The material of the target anode 9 is most often mounted on a thermally conductive material such as copper (shown as the anode thermal conductor 8). Furthermore, in the Coolidge-type X-ray tube design 12, the amount of electrical energy between the electrode potentials for generating a given resulting radiation 7 is very high, resulting in that the heating of the target anode 9 material requires specific target cooling considerations, such as rotating the target anode 9, air-cooling or liquid-cooling the target anode 9, e.g. by water or dielectric oil flowing through the cooling line 10. The purpose of cooling the anode assembly is to enable higher power operation of the X-ray tube.

Furthermore, the target anode 9 is placed in line with the electron beam 6, and the resulting radiation 7 is emitted through an output radiation window 11 forming a beam of output radiation 7 at right angles to the electron beam 6.

Referring now to fig. 2, a schematic cross-sectional representation of an exemplary embodiment of a high dose output, transmissive transmission target X-ray tube 20 is shown by way of example and not limitation. Preferably, the transmission transport target X-ray tube 20 includes an evacuated sealed chamber or enclosure, such as a housing 21, which may be glass, an alloy, or a metal, that creates an evacuated space 25. One end, the first end 21.1, of the housing 21 is preferably connected to a first connector 31 of a high voltage power supply 33. The main elements contained in the envelope 21 are preferably an anode structure 22, a cathode structure 23, a first filament lead 27, a second filament lead 28 and a filament 24. Furthermore, the anode structure 22 preferably comprises an anode which is a transmissive transmission target 43 as part of the anode structure 22, wherein the target 43 is preferably deposited on the inner surface 42 of the first end 21.1 of the housing 21. The cathode structure 23 may be connected to a second connector 32 of a ground or high voltage power supply 33. The filament 24 is preferably connected to a first filament leg 27 of the heating current power supply 44 and a second filament leg 28 of the heating current power supply 44. Preferably, the target 43 is an electron interactive material deposited on the inner surface 42 of the first end 21.1 of the housing 21 and includes the anode structure 22 in conjunction with an arc (arc) or circular cross-section, dome or hemispherical first end 21.1 of the housing 21. Also contained within the housing 21 is preferably a high voltage insulator 26 which partially encloses the housing 21 at the second end 21.2 of the housing 21.

Bremsstrahlung and characteristic radiation 30 is preferably emitted from the transmissive transmission target X-ray tube 20 in an arc or semi-circular cross-sectional, dome or hemispherical radiation pattern. When high-energy electrons are emitted from the heating filament 24 and accelerated by the high-voltage power supply 33, characteristic radiation is generated, and electrical energy (shown as electron trajectory 35) between the electrode potentials of the anode structure 22 and the cathode structure 23 on the surface of the target anode structure 22 knocks electrons off their trajectories from the target element (target 43). When this occurs, the electrons in the next higher energy orbit will fall into the lower energy orbit and emit a radiation burst equal to the energy difference between the two electron orbits. Because each element or material of the target 43 has a different atomic structure, the energy level of the emitted radiation is the only and characteristic of that element. The atomic levels are designated as K, L, M, n. For example, if there is a vacancy (vacancy) in the K track of the element (target 43) and an electron descends from the L track to fill the vacancy, the energy emitted is equal to E_X-ray＝E_kα-E_L. The main and most useful characteristic radiation is K of various elements (target 43)_αEnergy level and between the electrode potentials of the anode structure 22 and the cathode structure 23Occurs at an energy of less than 100 kilovolts (kV) of the high voltage power supply 33. It is recognized herein that the transmission target X-ray tube 20 may preferably use the K of the target 43 or a composite target 43 composed of various elements_αCharacteristic radiation to fill the X-ray spectrum below 100 kilovolts (kV) and using bremsstrahlung radiation from higher Z elements to produce X-ray tubes of excellent performance characteristics.

Table I target 43 materials or combinations

Anode material (43)	Bremsstrahlung	Characteristic radiation	Total radiation
				Tungsten	700	300	1000
Tungsten + gold	700	600	1300
				Tungsten + gold + iridium	700	900	1600

All readings normalized for the tungsten anode and operation at 200 kV.

Novel shell and target anode structure shape

To address the shortcomings of conventional X-ray tubes and current transmissive transmission tubes, transmissive transmission target X-ray tubes 20 having selective anode structures 22 have been designed. Preferably, the transmissive transport target X-ray tube 20 utilizes a large diameter 52, hemispherical structure of the anode structure 22 of the housing 21 formed of a low Z material (e.g., aluminum or beryllium, carbon, ceramic, stainless steel, or alloys thereof) for a substrate on which various target 43 elements or materials can be deposited to form the anode structure 22 (one target element is formed on the anode structure via one of electrochemical flat plating, mechanical bonding, or vapor deposition using evaporation or sputtering techniques). The hemispherical anode structure 22 is preferably used because it has twice the surface area of a disk-shaped substrate of the same diameter. A hemispherical area of 2 π r²And the area of the disk is pi r². This increase in surface area allows for increased power dissipation, improved electron symmetry on the target 43, increased surface area to dissipate heat, and thus increased cooling efficiency. Further, the anode substrate 22 can be coated with various elements, combinations of elements, or alloys thereof as targets 43 to form the anode structure 22 and produce the desired characteristic radiation 30 for a particular purpose or with high-Z elements to produce increased output with a combination of bremsstrahlung and characteristic radiation 30. This is done at a reduced cathode potential, high voltage power supply 33, as shown in fig. 1, for the same radiation 30 output as compared to the Coolidge-type X-ray tube.

It is recognized herein that a hemispherical anode structure 22 configured with a large surface area results in a self-cooled or cooler or lower temperature anode due to its ability to dissipate heat over a larger surface area, and therefore does not require any internal cooling system (e.g., a rotating anode or a cooling fluid with internal channels) to dissipate the heat generated in the anode structure 22 during operation.

It is also recognized herein that the hemispherical anode structure 22 configured with a large surface area symmetrically provides a uniform distribution of electrons across the anode structure 22 and thus generates a uniform distribution of radiation 30.

Target anode substrate for specific coating

Preferably, the transmission target X-ray tube 20 utilizes gold as the target 43 coated with the hemispherical anode structure 22, which is deposited on the inner surface 42 of the first end 21.1 of the housing 21 to form the anode structure 22. Gold, which is the target 43 element of anode structure 22, has a K at about 68.8keV_αPeak sum K at about 77keV_βPeak, when operating at 150 to 160 kilovolts (kV), the high voltage power supply 33, bremsstrahlung and characteristic radiation spectrum, radiation 30, as shown in fig. 3, is ideally suited for one or more high output radiation applications and far superior to conventional X-ray tubes, such as the Coolidge-type X-ray tube shown in fig. 1, because of the following advantages. Preferably, the gold target 43 element of the anode structure 22 provides the following advantage, since the efficiency of the generation of radiation 30 is proportional to the atomic number of the gold target 43 of the anode structure 22 multiplied by the kV of the high voltage supply 33. Here, for the Coolidge type X-ray tube, the atomic number of gold (Au) is 79 and the atomic number of tungsten (W) is 74, as shown in fig. 1. The percentage difference between the two atomic numbers was 6.75%. Formula 1- ((kV x Z)/(kV x Z)) x100 ═ efficiency kV x Z10 "6 based on the calculated radiation 30; the gold used as the target anode structure 22 is 6.75%, which is more efficient in generating bremsstrahlung, radiation 30 at the same kilovoltage (kV) level as the tungsten (W) target 43 for the Coolidge-type X-ray tube, as shown in fig. 1. The conventional X-ray tube cannot utilize the use of gold as the target 43 material for high power irradiation tubes because the melting temperature of gold (1064 degrees celsius) is lower compared to tungsten (3422 degrees celsius) due to the small surface area design of the anode thermal conductor 8 of the Coolidge type X-ray tube, as shown in fig. 1. However, for larger target 43 areas provided by the hemispherical structure of the anode structure 22 of the transmissive transmission target X-ray tube 20, the anode surface area 22 is increased, allowing for increased power dissipation; and therefore, a lower melting point element (e.g., gold) may be used for target 43. For example, the area of the anode structure 22 may be as large as 25 square inches, which may be at 1 megaW/cm, as compared to 1 square inch for the target anode 9 of FIG. 1²Is carried out byAnd (5) operating. It is contemplated herein that other lower melting point elements for target 43 (e.g., elements having atomic numbers between 74 and 82, and more particularly, lead (Pb) and uranium (U), etc.) may be used as the target 43 material for anode structure 22. It is also contemplated herein that the target 43 material for the anode structure 22 is preferably 4-40 microns thick, and the thickness of the target 43 material is selected depending on the material selected for the target 43, the desired type of radiation 30 emission, and the accelerating voltage of the high voltage power supply 33. These characteristics enable the monochromatic light beam to have an increased radiation at the lower kV of the high voltage power supply 33.

Target 43 is preferably formed of a suitable material, such as gold (Au) or lead (Pb), including other elements having atomic numbers between 74 and 82, and in addition copper (Cu), silver (Ag), and uranium (U) may be used for target anode structure 22. Preferably, these materials include other suitable properties, such as high K_αEnergy levels, high electron to X-ray conversion rates, or other beneficial characteristics as understood by those skilled in the art.

Furthermore, the filament 24 of the transmissive transmission target X-ray tube 20 is preferably configured in an arc or circular cross-sectional or hemispherical configuration, positioned within the cathode structure 23, and this configuration electrostatically focuses the electron beam 29 along the electron trajectory 35 towards the anode structure 22, or more particularly, in a 180 degree (180) or hemispherical pattern on the target 43 of the anode structure 22, with the electron beam 29 evenly distributed over the target 43, the inner surface 42 of the first end 21.1 of the anode structure 22 of the housing 21. Further, the filament 24 is preferably coated with an oxide material about 40-50 microns thick and the filament 24 is heated using indirect heating with a nicon wire connected to the first filament leg 27 of the heating current power supply 44 and the second filament leg 28 of the heating current power supply 44 to provide thermal vibrational energy for releasing electrons from the filament 24. As described above, this distribution of electron beam 29 across target 43 and anode structure 22 reduces or lessens the watt density (watts/area, W/cm) of target 43 and anode structure 22²) Loading and thus preventing hot spots due to uniform heating of the target 43 and the anode structure 22.

It is recognized herein that the hemispherical anode structure 22 and the arc or hemispherical filament 24 in combination symmetrically provide an even distribution of electrons across the target 43 and the anode structure 22.

It is also recognized herein that the hemispherical anode structure 22 and the arc or hemispherical filament 24 in combination provide a collimated electron trajectory 35 across the target 43 and the anode structure 22.

It is also recognized herein that the hemispherical anode structure 22 and arc or hemispherical filament 24 combination provide an equalized electron travel distance 58, the distance the electron beam 29 travels from the filament 24 to the target anode structure 22.

Still further, the cathode structure 23 of the transmission transfer target X-ray tube 20 is preferably configured in a 'V' or cut-out 'V' shaped cross-section, or a bowl or horn (flared) configuration, or the like, and such configuration electrostatically directs the electron beam 29 (effectively and uniformly distributed along the electron trajectory 35 towards the target 43 and the anode structure 22 or more particularly in a 180 degree (180 °) pattern onto the target 43 and the anode structure 22) such that the electron beam 29 is uniformly distributed over the hemispherical target 43 and the anode structure 22, the inner surface 42 of the first end 21.1 of the housing 21.

It is recognized herein that the hemispherical anode structure 22, the arc or hemispherical filament 24 and the horn cathode structure 23 in combination provide maximum generation of directional X-rays that are approximately symmetric about the centerline CL.

It is recognized herein that the transmission target X-ray tube 20 may include an anode structure with a specific coating target 43, for example by using parameters of the target 43 material.

It is recognized herein that the transmissive transmission target X-ray tube 20 may include a specific coating target 43 of the anode structure 22 for X-ray deflection, such as by using a low Z material for the target 43.

It is recognized herein that the transmissive transmission target X-ray tube 20 may include a specific coating target 43 of the anode structure 22 for X-ray deflection, for example by using a high-Z material for the target 43.

It is recognized herein that the transmissive transmission target X-ray tube 20 may include a specific coating target 43 of the anode structure 22, for example by using parameters of the target 43 material, for low power requirements of the high voltage power supply 33 or for high dose radiation 30 applications.

It is recognized herein that the transmissive transmission target X-ray tube 20 may include a specific coating target 43 of the anode structure 22 for medical imaging, for example by using molybdenum as the target 43 material.

It is recognized herein that the transmission target X-ray tube 20 may include a specifically coated target 43 of the anode structure 22 for industrial imaging, for example by using gold as the target 43 material to increase the number of X-ray photons, which in turn improves image contrast.

Referring now to fig. 3, by way of example and not limitation, a graphical representation of X-ray energy keV (λ) versus output radiation dose for a transmission target X-ray tube 20 is shown with target 43 gold material. In this graph of characteristic radiation R1 of the target 43 gold material, the Y-axis represents the dose (in photons) of a given amount of output radiation (radiation 30), e.g., the number of photons, and the X-axis represents kilovolts (wavelength) of a given amount of X-ray energy, and as the kilovolts (wavelength) varies, the number of photon doses represented by the graph for gold also varies. As shown in the figure, a radiation spike occurs and is designated as K_αAnd K_βA dose peak, which is a characteristic radiation peak resulting from the use of the target 43-gold material. Use of the target 43-gold material results in an increase in the radiation dose of the target 43-material without requiring an increase in the input power kV (λ), since the radiation spike is shown to approximately correspond to K_α68.7kV (. lamda.) of peak value and corresponding K_βPeak 77 kV (λ).

Referring now to fig. 3.1, a graphical representation of X-ray energy kV (λ) versus output radiation dose for a transmission target X-ray tube 20 is shown, by way of example and not limitation, showing a target 43 material having a configuration based on a combination of materials of the target 43. Preferably, target 43 is preferably formed of a suitable material, such as gold (Au), lead (Pb), including other elements having atomic numbers between 74 and 82, and additionally copper (Cu), silver (Ag), and uranium (U). In this diagram of the characteristic radiation R2 of the combined target 43 of gold tungsten and iridium, the Y-axis represents a given amount of radiationThe dose (in photons) of radiation, for example, the number of photons, and the X-axis represents the kilovolts (wavelength) of a given amount of X-ray energy, and as the kilovolts (wavelength) varies, the photon dose number also varies. As shown in the figure, the radiation spike occurs and is designated as K for both gold tungsten and iridium_αAnd K_βA peak, which is a characteristic radiation peak caused by the material combination using the target 43. The use of target 43 materials of gold tungsten and iridium results in an increase in the output radiation (radiation 30) dose that the combined material, target 43, undergoes without requiring an increase in the input power kV (λ) high voltage power supply 33, as shown in table II.

TABLE II

Element(s)	K_αl	K_α2	K_β1	K _β2
					Gold (Au)	68.804	66.990	77.985	80.182
Tungsten	59.318	57.982	67.244	69.1
					Iridium (III)	64.896	63.287	73.560	75.620

It is recognized herein that for each material selected from the above list of elements to constitute a combined material, for target 43, as shown, characteristic radiation R2 of target 43 will have additional and different K_αAnd K_βThe peak value of each material selected and added to target 43. It is also recognized herein that the addition of each material selected from the above list results in the generation of additional and different K's based on the multiple materials and the target 43 material for each material_αAnd K_βPeak, and therefore increasing output radiation dose, is directed to the radiation 30 that is incident on the target 43 of the combined material, without requiring an increase in input power kV (λ), high voltage power supply 33, as indicated by the increase in area of the plot of characteristic radiation R2. By adding a plurality of the combined materials listed above to the target 43, the improved output radiation (radiation 30) dose occurring for the target 43 of combined materials will be greatly increased. If a power of 1000 watts generates a dose of 100 Gray (Gray), the combined target 43 may generate more than 50% of the dose.

It is also recognized that based on the above material list, an increase in the output radiation (radiation 30) dose to the target 43 material without requiring an increase in the input power (kV (λ) × mA), reduces the cooling requirements.

It is also recognized, based on the above material list, that an increase in the output radiation (radiation 30) dose to the target 43 material without requiring an increase in the input power (kV (λ) × mA), high voltage power supply 33, enables radiation and illumination applications, such as medical applications, at lower input powers (kV (λ) × mA).

Referring now to fig. 4, by way of example and not limitation, a graphical representation of empirically determined thickness versus dose output radiation (radiation 30) for a target 43 material of a transmission target X-ray tube 20 is shown. In this plot of bremsstrahlung versus material thickness for target 43, the Y-axis represents the dose (in photons) of a given amount of output radiation (radiation 30), e.g., the number of photons, and the X-axis represents the target 43 material thickness (in microns), and as the target 43 material thickness changes, the number of photons represented by the graph also changes. Representative curves exhibit electrical energy between the electrode potentials of anode structure 22 and cathode structure 23, e.g., 50kV Ra, 100kV Rb, and 200kV Rc, for varying high voltage power supply 33. In each curve, the dose is ramped up, smoothed (plateaus) and tapered down based on increasing target 43 material thickness. It is recognized herein in fig. 4 that the thickness of the target 43 material of the target anode structure 22 is preferably about 4-40 microns, and more preferably the thickness of the target 43 material of the target anode structure 22 is about 4-18 microns, with the thickness of the target 43 material being selected depending on the material selected, the desired type of emission of radiation 30, and the acceleration voltage of the high voltage power supply 33.

In fig. 4, it is also recognized herein that a higher acceleration voltage, high voltage power supply 33, more efficient transmission transfer target X-ray tube 20 is in position to convert electrons emitted by the filament 24 into increased dose output radiation 30 to take advantage of the characteristic radiation peak.

In fig. 4, it is also recognized herein that no sharp points appear in the characteristic radiation R-curves 50kV Ra, 100kV Rb, and 200kV Rc, and more particularly, that the representative radiation R-curve 50kV Ra has a plateau from the target 43 of about 3-5 microns thickness, the representative radiation R-curve 100kV Rb has a plateau from the target 43 of about 7-10 microns thickness, the representative radiation R-curve 200kV Rc has a plateau from the target 43 of about 14-18 microns thickness, and the collective representative radiation R-curves 50kV Ra, 100kV Rb, and 200kV Rc have plateaus from the target 43 of preferably 4-18 microns thickness.

Design variables of the transmission target X-ray tube 20, such as the material to be selected for the target 43 (material with Z from 73 to 79 heavy elements), the selected target 43 material thickness (in microns), and the selected voltage of the high voltage power supply 33, change the dose output radiation 30, such as an increased dose output radiation at a lower high voltage power supply 33 power.

In fig. 4, it is again recognized herein that varying the selected target 43 material and/or the selected target 43 material thickness (in microns) varies the dose output radiation 30.

In fig. 4, it is still further recognized herein that the dual energy of a transmissive transmission target X-ray tube 20 having a high voltage power supply 33 operating at two voltages, e.g., 50kV and 100kV, may preferably select the material thickness of the target 43 to accommodate both energies, e.g., a target 43 material thickness of between 3-10 microns may be selected, where a smooth representative radiation R-curve 50kV Ra having a thickness of about 3-5 microns from the target 43(50kV) and a smooth representative radiation R-curve 100kV Rb having a thickness of about 7-10 microns from the target 43(100kV) overlap.

Referring now to fig. 5, an exemplary application of irradiating biological material with a transmissive transport target X-ray tube 20 is shown by way of example and not limitation, as shown and described in fig. 2. In use, characteristic radiation of the transmission target X-ray tube 20 and bremsstrahlung radiation 30 are preferably emitted from the transmission target X-ray tube 20 in an arc or semi-circular cross-sectional, dome or hemispherical radiation 30 mode. Preferably, the transmission target X-ray tube 20 produces characteristic radiation 30, the characteristic radiation 30 being configured such that a large area of intense radiation 30 can improve throughput radiation while radiating a greater or increased number of samples S. Furthermore, the sample S may be positioned close to or adjacent to the anode structure 22 of the housing 21 of the transmission target X-ray tube 20, either stationary or by moving mechanical structures, depending on the application by utilizing the geometry of the radiation pattern to increase the uniform exposure level of the sample S to the characteristic radiation 30. Furthermore, the transmission target X-ray tube 20 preferably generates a symmetric radiation field (radiation 30) around the first end 21.1 of the transmission target X-ray tube 20 to provide a uniform dose of radiation 30 to all areas of the sample S.

It is recognized herein that the radiation 30 output of the transmission delivery target X-ray tube 20 is increased compared to the Coolidge-type prior art X-ray tube, shown in fig. 1. For example, if the transmission target X-ray tube 20 outputs twice the radiation 30, then the sample S requires half the run time required to transmit the transmission target X-ray tube 20 and additionally uses a higher dose of radiation 30 with lower power requirements for the high voltage power supply 33, a lower thermal load in the BTU with lower power demand air conditioning load savings, the high voltage power supply 33, all resulting in lower operating costs for the transmission target X-ray tube 20. Furthermore, the hemispherical anode structure 22 configured with a large surface area results in a cooled or cooler or lower temperature anode due to its heat sinking capability, and therefore does not require any internal or external cooling system, such as a rotating anode or a cooling fluid with internal channels, to dissipate the heat generated in the anode structure 22 during operation and thus reduce the operating cost of the transmission target X-ray tube 20.

Referring now to fig. 6, a schematic cross-sectional representation of the high dose output of a transmissive transmission target X-ray tube 20 (alternatively 50) in combination with a monochromatic filter 60 is shown by way of example and not limitation. Preferably, a monochromatic filter 60 may be positioned near or adjacent the first end 21.1 of the housing 21 in the path of the radiation 30 or between the anode structure 22 of the housing 21 and the sample S in the path of the radiation 30 to attenuate or filter selected radiation from the radiation 30. Referring again to FIG. 3, by way of example and not limitation, the monochromatic filter 60 may be configured to filter or attenuate designated radiation 30, e.g., all radiation less than 54kV (λ), to produce K_αAnd K_βA peak dose of the prescribed radiation 30, which is a characteristic radiation peak induced using the target 43-gold material.

Fig. 7 is a graphical representation of X-ray energy versus dose of the transmission target X-ray tube 20 combined with the monochromatic filter of fig. 6. In this graph of characteristic radiation R1 of the target 43 gold material, the Y-axis represents the dose (in photons) of a given amount of output radiation 30, e.g., the number of photons, and the X-axis represents the kilovolts (wavelength) of a given amount of X-ray energy, between 75kV (λ) and 85kV (λ), and varies with kilovolts (wavelength)The number of photons represented by the dose radiation 30 for gold also varies. As shown in the figure, a radiation spike occurs and is designated as K_αAnd K_βA dose peak, which is a characteristic radiation peak resulting from the use of the target 43-gold material. In use, the selected target 43 material and its K_αAnd K_βThe dose peaks and selected monochromatic filters 60 preferably achieve a desired radiation distribution of the radiation 30, and thus can be specified to achieve various specific radiation 30 distributions for the transmission target X-ray tube 20 (alternatively 50) for a specified image and treatment example or use.

Referring now to fig. 8, a schematic cross-sectional representation of an alternative exemplary embodiment of a high dose output of a transmission target X-ray tube 50 is shown by way of example and not limitation. Preferably, the transmission target X-ray tube 50 comprises an evacuated sealed chamber or enclosure, such as the housing 21, which may be glass, an alloy, or a metal. One end, the first end 21.1, of the housing 21 is preferably connected to a first connector 31 of a high voltage power supply 33. The main elements contained within the envelope 21 are preferably an anode structure 22, a cathode structure 23.1, a first filament leg 27, a second filament leg 28, and a filament 24.1. Furthermore, the anode structure 22 preferably comprises an anode which is transmissive to the transmission target 43 as part of the anode structure 22, wherein the target 43 is preferably deposited on the inner surface 42 of the first end 21.1 of the housing 21. The cathode structure 23.1 may be connected to a second connector 32 of a ground or high voltage power supply 33. The filament 24.1 is preferably connected to a first filament leg 27 of the heating current supply 44 and a second filament leg 28 of the heating current supply 44. The target 43 is an electronically interactive material deposited on the inner surface 42 of the first end 21.1 of the housing 21 (i.e., one target element is formed on the anode structure via one of electrochemical plating, mechanical bonding, or vapor deposition using evaporation or sputtering techniques) and includes the anode structure 22 with the arc or circular cross-section, dome or hemispherical first end 21.1 of the housing 21. Also contained within the housing 21 is preferably a high voltage insulator 26 which partially encloses the housing 21 at the second end 21.2 of the housing 21.

The filament 24.1 of the transmissive transmission target X-ray tube 50 is preferably configured in a straight or slightly curved cross-sectional or planar or disk-shaped configuration within the cathode structure 23.1 and this configuration electrostatically focuses the electron beam 29.1 along the electron trajectory 35.1 towards the target 43 and the anode structure 22 or, more particularly, in a focal spot configuration pattern on the target 43 and the anode structure 22 such that the electron beam 29.1 near the center line CL is focused on the inner surface 42 of the first end 21.1 of the housing 21.

Still further, the cathode structure 23.1 of the transmission transfer target X-ray tube 50 is preferably configured in a 'U' shaped cross-section, or a cylindrical configuration or other focusing configuration, and this configuration electrostatically accelerates the electron beam 29.1 along the electron narrow trajectory 35.1 towards the target 43 and the anode structure 22, or more particularly onto the target 43 and the anode structure 22 in a focusing mode, so that the electron beam 29.1 is narrowly distributed over the hemispherical target 43 and the anode structure 22, the inner surface 42 of the first end 21.1 of the housing 21. This concentration of the electron beam 29.1 enables high dose output in a narrow spot configuration with an anode diameter 52, and the transmission target X-ray tube 50 can be used, such as for applications to produce focused X-rays for radiotherapy, imaging (e.g., medical, industrial and dual energy, non-destructive evaluation of objects).

As shown in fig. 2, it is contemplated herein that the spot diameter 52 may be scaled up/down or increased or decreased in size based on design factors such as the opening or gap, e.g., the inner diameter 56 of the cathode structure 23.1, the electron travel distance 58 of the electron beam 29.1 (the distance the electrons travel from the filament 24.1 to the target 43 and the anode structure 22), and/or the diameter 52 of the hemispherical anode structure 22 of the housing 21. For example, in use, the inner diameter 56 of the cathode structure 23.1, the electron travel distance 58 of the electrons of the electron beam 29.1 traveling from the filament 24.1 to the target 43 and the anode structure 22, and/or the diameter 52 of the hemispherical anode structure 22 of the housing 21 may be specified to achieve a spot diameter 52 proportional to the tumor size at the depth of radiation treatment subject to X-ray penetration.

It is recognized herein that the hemispherical anode structure 22, the filament 24.1 and the narrowed cathode structure 23.1 in combination provide a focused, radially linearly symmetric X-ray field.

It is recognized herein that the hemispherical anode structure 22, filament 24.1 and horn cathode structure 23.1 in combination generate directed X-rays close to the centerline CL for therapeutic X-ray treatment of melanoma and other cancer cells.

It is contemplated herein that the monochromatic filter 60 may be used with a transmission target X-ray tube 50 similar to that shown and disclosed in fig. 6 and 7.

Referring now to FIG. 9, by way of example and not limitation, a flow chart 900 of a method of generating a symmetric hemispherical X-ray field is shown. In block or step 910, a high dose delivery, transmissive transmission target X-ray tube 20/50 is provided having an evacuated sealed envelope 21, a hemispherical anode structure 22, a cathode structure 23, a target 43, and a filament 24 as described herein. In block or step 915, a material or combination of materials, z-materials, is selected for the target 43. In block or step 920, an acceleration voltage is selected for the high voltage power supply 33. In block or step 925, the high dose output transmission target X-ray tube 20 is caused to generate X-rays for the biological material radiation. In block or step 930, the high dose output transmission target X-ray tube 20 is caused to generate X-rays for non-destructive evaluation of the object. In block or step 935, the high dose output transmission target X-ray tube 20 is caused to generate X-rays for destructive treatment of the biological sample. Other treatments include imaging, such as medical, industrial, and dual energy, non-destructive evaluation of objects.

Referring now to fig. 10, a schematic cross-sectional representation of an alternative exemplary embodiment of a high dose output, transmissive and reflective target X-ray tube 50 is shown by way of example and not limitation. Preferably, the transmission and reflection target X-ray tube 50 includes an evacuated sealed chamber or enclosure, such as the housing 21, which may be glass, an alloy, or a metal. One end, the first end 21.1, of the housing 21 is preferably connected to a first connector 31 of a high voltage power supply 33. The main elements contained within the envelope 21 are preferably an anode structure 22, a cathode structure 23.1, a first filament leg 27, a second filament leg 28, and a filament 24.1. Furthermore, the anode structure 22 preferably comprises an anode which is transmissive to the transmission and reflection target 43 as part of the anode structure 22, wherein the target 43 is preferably deposited on the inner surface 42 of the first end 21.1 of the housing 21. The cathode structure 23.1 may be connected to a second connector 32 of a ground or high voltage power supply 33. The filament 24.1 is preferably connected to a first filament leg 27 of the heating current supply 44 and a second filament leg 28 of the heating current supply 44. The target 43 is an electronically interactive material deposited on the inner surface 42 of the first end 21.1 of the housing 21 and includes the anode structure 22 with the arc or circular cross-section 2-D circular or base or hemispherical first end 21.1 of the housing 21. Also contained within the housing 21 is preferably a high voltage insulator 26 which partially encloses the housing 21 at the second end 21.2 of the housing 21.

The filament 24.1 of the transmissive and reflective target X-ray tube 50 is preferably configured in a straight or slightly curved cross-sectional or planar or disk-shaped configuration within the cathode structure 23.1 and this configuration electrostatically shapes the electron beam 29.1 along the electron trajectory 35.1 towards the target 43 and the anode structure 22 or, more particularly, in a large area electron pattern on the target 43 and the anode structure 22 so that the electron beam 29.1 near the center line CL is evenly distributed over the inner surface 42 of the first end 21.1 of the housing 21. Furthermore, the filament 24.1 and the cathode structure 23.1 may be positioned between the centre line CL of the 2-D circular or base or hemispherical first end 21.1 (the centre of the circle created by the 2D base of the hemispherical first end 21.1) close to the arc or circular cross section of the envelope 21 and the cathode structure 23.1.

Still further, the cathode structure 23.1 of the transmission and reflection target X-ray tube 50 preferably or may be configured in a 'U' shaped cross-section, or a cylindrical configuration or other defocused configuration, and such configuration electrostatically accelerates the electron beam 29.1 along the electron trajectory 35.1 towards the target 43 and the anode structure 22, or more particularly, onto the target 43 and the anode structure 22 in a defocused mode, so as to evenly distribute the electron beam 29.1 over the hemispherical target 43 and the anode structure 22, the inner surface 42 of the first end 21.1 of the housing 21. This distribution of the electron beam 29.1 enables a high dose output in a forward photon direction 52 (forward being defined as the same direction in which electrons of the electron beam 29.1 travel in a fan-shaped pattern from the anode structure 22) and using a reflected photon direction 52.1 opposite to the anode 22 (reflecting the opposite direction in which electrons defined as the electron beam 29.1 travel in a reflected fan-shaped pattern from the anode structure 22). In use, an alternative exemplary embodiment of a high dose output transmissive and reflective target X-ray tube 50 may be used for batch irradiation of the sample S via the transmissive and reflective transmissive target X-ray tube 50, for example a symmetric radiation field (radiation 30) around the first end 21.1 of the transmissive target X-ray tube 50 to provide a uniform dose of radiation 30 to all areas of the sample S positioned proximate the first end 21.1 of the housing 21, as shown in fig. 5; and to transmit the alternative radiation field (radiation 30) around the second end 21.2 of the transmission target X-ray tube 50 to provide a consistent dose of radiation 30 to all areas of the sample S positioned parallel to the centerline CL and proximate to the housing 21 (and more particularly proximate to the second end 21.2 of the housing 21), as shown in fig. 13.

Referring now to fig. 11A, a cross-sectional representation of an alternative exemplary embodiment of a high dose output transmissive and reflective target X-ray tube 50A showing a reduced anode-to-cathode spacing is shown by way of example and not limitation. Preferably, alternative exemplary embodiments of the high dose output transmissive and reflective target X-ray tube 50A may be configured with the first end 21.1 of the housing 21 or the anode structure 22 positioned at a linear distance 110A from the second end 21.2 of the housing 21, or the cathode structure 23.1; wherein the electron beam 29.1 achieves a high dose output in a direction that passes forward photons 52A (forward being defined as the same direction in which electrons of the electron beam 29.1 travel in a wide fan pattern from the anode structure 22) and in a direction 52.1A that uses reflected photons opposite to the anode 22 (reflecting the opposite direction in which electrons defined as the electron beam 29.1 travel in a reflected wide fan pattern from the anode structure 22).

Referring now to fig. 11B, a cross-sectional representation of an alternative exemplary embodiment of a high dose output transmissive and reflective target X-ray tube 50B is shown, by way of example and not limitation, illustrating an increased anode-to-cathode spacing. Preferably, alternative exemplary embodiments of the high dose output transmissive and reflective target X-ray tube 50B may be configured with the first end 21.1 of the housing 21, or the anode structure 22 or the cathode structure 23.1 positioned at a linear distance 110B from the second end 21.2 of the housing 21; wherein the electron beam 29.1 achieves a high dose output in a forward photon direction 52B (forward being defined as the same direction in which electrons of the electron beam 29.1 travel in a narrow fan pattern from the anode structure 22) and in a reflected photon direction 52.1B opposite the anode 22 (reflecting the opposite direction in which electrons of the electron beam 29.1 travel in a reflected narrow fan pattern from the anode structure 22).

As shown in fig. 2, it is contemplated herein that spot diameter 52 may scale up/down or increase or decrease the size of radiation 30 based on design factors such as the opening or gap, e.g., cathode inner diameter 56 of cathode structure 23.1, electron travel distance 58 of electron beam 29.1 (the distance electrons travel from filament 24.1 to target 43, anode structure 22), and/or diameter 52 of hemispherical anode structure 22 of housing 21.

Referring now to fig. 11A, 11B and 12, by way of example and not limitation, there is shown a graphical representation of photon intensity passing through radiation 30 (e.g., photon P (transmitted X-ray field or spectrum)) and reflected radiation 30 (e.g., reflected photon Pr (transmitted X-ray field or spectrum)) with respect to a tube centerline C1 of a transmission-target X-ray tube 50. In this plot of the normalized amount of radiation 30, the photons P/Pr that are transmitted and reflected through the target X-ray tube 50 are, for example, a distance in millimeters from the baseline insulator 26 that partially encloses the housing 21 at the second end 21.2 of the housing 21. Further, referring again to fig. 12, the Y-axis represents a normalized photon number P/Pr (in photons), e.g., number of photons, for a given amount of photons P and reflected photons Pr with respect to the tube centerline C1 through which the targeted X-ray tube 50 or output radiation (radiation 30) is transmitted and reflected, and the X-axis represents the distance in millimeters from the baseline insulator 26 that partially encloses the housing 21 at the second end 21.2 of the housing 21. In this exemplary illustration of fig. 12, the first end 21.1 of the housing 21 or the anode structure 22 is positioned at a linear distance 110A from the second end 21.2 of the housing 21 or the cathode structure 23.1 may be about two hundred seventy five millimeters (275 millimeters), and the normalized photon count P/Pr is ramped up, smoothed, and gradually decreased based on the linear distance 110A. It is recognized herein in fig. 12 that the normalized photon count P/Pr slopes upward at about seventy-five millimeters (75 millimeters), levels off from about seventy-five millimeters (75 millimeters) to about two hundred and fifty millimeters (250 millimeters), and gradually decreases thereafter at two hundred and fifty millimeters (250 millimeters).

It is contemplated herein that the design factors listed above may be used to change or adjust the location, position, amount, and number of transmitted photons P and/or reflected photons Pr generated by the transmission and reflection target X-ray tube 50.

As shown in fig. 11A, 11B and 12, it is contemplated herein that design factors such a linear distance 110 (e.g., the first end 21.1 of the housing 21 or the anode structure 22 positioned at the linear distance 110 from the second end 21.2 of the housing 21 or the cathode structure 23.1), and such an adjusted, designed or predetermined linear distance 110, change the number and location of reflected photons Pr (reflectively transmitted X-ray field) generated by the reflectively transmitted target X-ray tube 50.

In fig. 12, it is also recognized herein that the spot diameter 52 may scale up/down or increase or decrease the size of the reflected photons Pr or side output radiation 30 based on design factors (e.g., the linear distance 110A of the first end 21.1 or anode structure 22 of the housing 21 from the second end 21.2 or cathode structure 23.1 of the housing, as shown in fig. 2, and the opening or gap, such as the cathode inner diameter 56 of the cathode structure 23.1, the electron travel distance 58 of the electron beam 29.1 (the distance the electron travels from the filament 24.1 to the target 43, the anode structure 22), and/or the diameter 52 of the hemispherical anode structure 22 of the housing 21).

Referring now to fig. 13, by way of example and not limitation, an exemplary application of irradiating biological material with a transmissive transport target X-ray tube 50 is shown, as shown and described in fig. 10. In use, the transmission target X-ray tube 50 characteristic radiation and bremsstrahlung radiation 30, P is preferably emitted from the transmission target X-ray tube 50 in an arc or semi-circular cross-section, dome or hemispherical radiation 30 mode (symmetrical hemispherical reflective transmission X-ray field) and reflected radiation 30, Pr is preferably reflected from the anode structure 22 of the housing 21, as shown in fig. 12. Preferably, the transmission and reflection transmission target X-ray tube 50 generates characteristic radiation 30, the characteristic radiation 30 being configured such that a large area of intense radiation 30 can improve throughput radiation and reflected radiation while radiating a greater or increased number of samples S. Furthermore, the sample S may be positioned close to, near or adjacent to the anode structure 22 of the housing 21 of the transmission target X-ray tube 50, either stationary or by moving mechanical structures, depending on the application by using the geometry of the radiation pattern to increase the uniform exposure level of the sample S to the characteristic radiation 30 (passing through photons P and/or reflecting photons Pr), as described in fig. 12. Furthermore, the transmission target X-ray tube 50 preferably creates a symmetric radiation field (radiation 30) (passing photons P and/or reflected photons Pr) around the first and second ends 21.1, 21.2 of the housing 21 of the transmission target X-ray tube 50 to provide a consistent dose of radiation 30 to all areas of the sample S.

It is recognized herein that the radiation 30 output of the X-ray tube 50 of the transmission target through and reflected is increased compared to the Coolidge-type prior art X-ray tube, as shown in fig. 1. For example, if the transmission target X-ray tube 50 outputs twice the radiation 30, then the sample S requires half the required run time of the transmission target X-ray tube 50 and additionally uses a higher dose of radiation 30 with lower power requirements for the high voltage power supply 33, a lower thermal load in the BTU with air conditioning load savings of lower power requirements, the high voltage power supply 33, all resulting in lower operating costs for the transmission target X-ray tube 20. Furthermore, the hemispherical anode structure 22 configured with a large surface area results in a cooled or cooler or lower temperature anode due to its heat sinking capability, and therefore does not require any internal or external cooling system, such as a rotating anode or a cooling fluid with internal channels, to dissipate the heat generated in the anode structure 22 during operation and thus reduce the operating cost of the transmissive transmission target X-ray tube 50.

It is still further contemplated herein that the transmission transmitting and reflecting target X-ray tube 50 radiation 30 may be used as a biological cell irradiator, a virus inactivation irradiator, an insect irradiator, a blood irradiator, a food irradiator.

The foregoing description and drawings comprise illustrative embodiments of the disclosure. Having thus described exemplary embodiments, it should be noted by those of ordinary skill in the art that the within disclosure is exemplary only and that various other substitutions, adaptations, and modifications may be made within the scope of the present invention. Listing or numbering the steps of a method in only a certain order does not constitute any limitation on the order of the steps of the method. Many modifications and other embodiments of the invention will come to mind to one skilled in the art to which this invention pertains having the benefit of the teachings presented in the foregoing descriptions and the associated drawings. Although specific terms may be employed herein, they are used in a generic and descriptive sense only and not for purposes of limitation. Furthermore, the present invention has been described in detail; it should be understood that various changes, substitutions and alterations can be made hereto without departing from the spirit and scope of the invention as defined by the appended claims. Accordingly, the present invention is not limited to the specific embodiments shown herein, but only by the following claims.

Claims

1. An X-ray tube for accelerating electrons at a high voltage potential, the X-ray tube comprising:

a housing;

a transmissive and reflective target anode structure disposed on the housing, the anode structure configured as a hemisphere having a center of a circle created by a 2D substrate, wherein the anode structure disposed on the housing is formed of a low Z material, and wherein the target anode structure has a target element coated thereon, the target element having a thickness between 2 and 50 microns;

a cathode structure disposed in the housing, the cathode structure configured to deflect the electrons toward the anode structure;

a filament disposed in the housing, the filament positioned proximate a center of the circle created by the hemispherical 2D substrate and between the anode structure and the cathode structure;

wherein the housing is configured to vacuum seal the anode structure, the cathode structure, and the filament therein, and

wherein the low-Z material of the anode structure, the thickness of the target element, and the hemisphere shape of the anode structure facilitate generation of forward photons traveling in the same direction as the electrons and reflected photons traveling in a different direction than the electrons based on the electrons in contact with the target element, and facilitate irradiation with the forward photons and the reflected photons, a sample included around a housing of the X-ray tube being located in the same direction as the electrons and in the different direction.

2. The X-ray tube of claim 1, wherein the anode structure is coated with at least one target element to generate bremsstrahlung X-rays from a plurality of accelerated electrons originating from the filament.

3. The X-ray tube of claim 1, wherein the anode structure is formed of a material that is substantially X-ray transparent.

4. The X-ray tube of claim 2, wherein the at least one target element is formed on the anode structure via one of electrochemical flat plating, mechanical bonding, or vapor deposition using evaporation or sputtering techniques.

5. The X-ray tube of claim 3, wherein the material consists of one or more of the group consisting of: beryllium, carbon, aluminum, ceramic, stainless steel, or alloys thereof.

6. The X-ray tube of claim 2, wherein the cathode structure generates an electrostatic field that evenly distributes the plurality of accelerated electrons originating from the filament onto the at least one target element formed on the anode structure.

7. The X-ray tube of claim 6, wherein the X-ray tube produces a symmetric hemispherical transmission X-ray field.

8. The X-ray tube of claim 7, wherein the X-ray tube generates a reflected transmitted X-ray field.

9. The X-ray tube of claim 8, wherein a linear distance between the anode structure and the cathode structure is adjusted to vary the reflected transmitted X-ray field.

10. The X-ray tube of claim 9, wherein the linear distance is increased to produce a broad mode of the reflected transmitted X-ray field.

11. The X-ray tube of claim 9, wherein the linear distance is reduced to produce a narrow mode of the reflected transmitted X-ray field.

12. The X-ray tube of claim 4, wherein the at least one target element has a thickness between 2 and 50 microns.

13. The X-ray tube of claim 4, wherein the X-ray tube produces an output X-ray spectrum determined by the at least one target element and a cathode voltage.

14. The X-ray tube of claim 4, wherein the X-ray tube produces k from the at least one target element_αAn output X-ray spectrum determined by the energy line and the cathode voltage.

15. The X-ray tube of claim 12, wherein at least one target element thickness is determined by a cathode voltage and a conversion efficiency of the at least one target element.

16. A method of producing a monochromatic output X-ray spectrum, the method comprising the steps of:

providing an X-ray tube for accelerating electrons at a high voltage potential, the X-ray tube comprising:

an evacuated enclosure that is sealed; a transmissive and reflective transmissive target anode structure disposed on the housing, the anode structure configured as a hemisphere having a geometric center; a cathode structure disposed in the housing, the cathode structure configured to deflect the electrons toward the anode structure; a filament disposed in the enclosure, the filament positioned proximate a center of a circle created by the hemispherical 2D substrate and between the anode structure and the cathode structure, wherein the circle created by hemispherical 2D substrate is in direct contact with the cathode structure, and wherein the evacuated enclosure is configured to: wherein the anode structure, the cathode structure, and the filament are vacuum sealed; for K lower than at least one target element only_αFiltering said output X-ray spectrum of energy; and is

For said K higher than only said at least one target element_αThe cathode voltage of the energy is adjusted,

adjusting a linear distance between the anode structure and the cathode structure to change a reflected transmitted X-ray field, wherein the linear distance is increased to produce a broad mode of the reflected transmitted X-ray field or decreased to produce a narrow mode of the reflected transmitted X-ray field.

17. The method of claim 16, further comprising the step of generating an output X-ray spectral level having a known spectrum for analysis of the X-ray application.

18. The method of claim 16, further comprising the step of generating a dual output X-ray spectrum for image analysis.

19. The method of claim 17, further comprising the step of generating a transmission and reflection target output X-ray spectrum.

20. The method of claim 17, further comprising the step of generating a reflected target output X-ray spectrum.

21. The method of claim 17, further comprising selecting k of the at least one target element by utilizing_αA line, a step of generating a high level of output X-ray spectrum at a reduced cathode voltage.

22. The method of claim 16, further comprising the step of generating a target output X-ray spectrum for the biological cell irradiator.

23. A method according to claim 16, further comprising the step of generating a target output X-ray spectrum for the virus-inactivation irradiator.

24. The method of claim 16, further comprising the step of generating a target output X-ray spectrum for the insect irradiator.

25. The method of claim 16, further comprising the step of generating a target output X-ray spectrum for the food irradiator.