WO2014172932A1 - X-ray source for medical detection, and movable ct scanner - Google Patents

X-ray source for medical detection, and movable ct scanner Download PDF

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Publication number
WO2014172932A1
WO2014172932A1 PCT/CN2013/076000 CN2013076000W WO2014172932A1 WO 2014172932 A1 WO2014172932 A1 WO 2014172932A1 CN 2013076000 W CN2013076000 W CN 2013076000W WO 2014172932 A1 WO2014172932 A1 WO 2014172932A1
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Prior art keywords
anode
cold cathode
ray source
medical detection
lab6
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PCT/CN2013/076000
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French (fr)
Chinese (zh)
Inventor
徐如祥
代秋声
高枫
张涛
Original Assignee
中国人民解放军北京军区总医院
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    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
    • A61B6/00—Apparatus for radiation diagnosis, e.g. combined with radiation therapy equipment
    • A61B6/44—Constructional features of apparatus for radiation diagnosis
    • A61B6/4405—Constructional features of apparatus for radiation diagnosis the apparatus being movable or portable, e.g. handheld or mounted on a trolley
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
    • A61B6/00—Apparatus for radiation diagnosis, e.g. combined with radiation therapy equipment
    • A61B6/02—Devices for diagnosis sequentially in different planes; Stereoscopic radiation diagnosis
    • A61B6/03—Computerised tomographs
    • A61B6/032—Transmission computed tomography [CT]
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
    • A61B6/00—Apparatus for radiation diagnosis, e.g. combined with radiation therapy equipment
    • A61B6/40—Apparatus for radiation diagnosis, e.g. combined with radiation therapy equipment with arrangements for generating radiation specially adapted for radiation diagnosis
    • H—ELECTRICITY
    • H01—ELECTRIC ELEMENTS
    • H01J—ELECTRIC DISCHARGE TUBES OR DISCHARGE LAMPS
    • H01J35/00—X-ray tubes
    • H01J35/02—Details
    • H01J35/04—Electrodes ; Mutual position thereof; Constructional adaptations therefor
    • H01J35/06—Cathodes
    • H01J35/065—Field emission, photo emission or secondary emission cathodes

Definitions

  • the present invention claims the priority of a Chinese patent application filed on April 27, 2013, to the Chinese National Intellectual Property No. 201310151634.2, entitled "X-ray source for medical testing and mobile CT scanner”.
  • the present invention relates to the field of medical devices, and in particular to an X-ray source for medical detection and a mobile CT scanner.
  • the X-ray tube is a key component of a small medical CT device.
  • the cathode is the core component of the X-ray tube, which directly determines the performance of the X-ray tube, the quality of the image such as resolution and contrast, and the efficiency of the whole machine.
  • the X-ray tube is usually an X-ray tube based on tungsten (W) wire thermal emission, that is, a cathode made of tungsten (W) wire to make an X-ray tube, and the working principle is that the tungsten (W) wire is heated to its working temperature. When electrons are emitted, the electrons that are emitted by the heat bombard the anode, thereby generating X-rays.
  • W tungsten
  • W tungsten
  • the prior art X-ray tube based on tungsten (W) wire thermal emission has at least the following disadvantages:
  • the cathode in the existing X-ray tube uses tungsten with high electron emission work ((
  • ) W 4.52 eV), emission The current density is small.
  • the pure tungsten material has a thermal emission current density of only 0.3 A/cm 2 at 2200 ° C. If a larger total emission current is to be obtained, the cathode temperature is usually increased, but the cathode temperature is increased to cause the cathode material.
  • the evaporation rate increases, the evaporation of the cathode material causes the tungsten filament to become thinner, and the thinned tungsten cathode increases the cathode temperature and the cathode evaporation is intensified, thereby forming a vicious cycle; in addition, the evaporated tungsten cathode material is deposited on the On the shell, a continuous or intermittent tungsten conductive film is formed, which destroys the insulation strength of the X-ray tube, reduces the tube pressure, and the tube is scrapped, thereby reducing the life of the X-ray tube; at the same time, the tungsten conductive film blocks the output.
  • the X-ray intensity of the window reduces the imaging sensitivity.
  • the invention provides an X-ray source for medical detection and a mobile CT scanner for improving the overall performance of the X-ray source, and can meet the application requirements of medical testing and the like.
  • the invention provides an X-ray source for medical detection, comprising:
  • An X-ray tube comprising an anode, a cold cathode, a grid and a casing; the casing for supporting the anode, the cold cathode and the grid, and the vacuum working environment of the anode, the cold cathode and the gate to the outside Insulation, the anode is grounded;
  • a high voltage generator for providing a first electric field between the cold cathode and the gate to cause electrons to be emitted by the cold cathode field, and a second electric field between the gate and the anode to accelerate The cold-exposed electrons are bombarded with the anode to produce X-rays.
  • the invention also provides a mobile CT scanner comprising an X-ray source for medical testing as described above.
  • the X-ray source adopts a cold cathode, and a gate is arranged between the anode and the cold cathode. Due to the protection of the grid, most of the air ions cannot directly collide with the cold cathode, thereby reducing the cold cathode being irradiated. The probability of damage;
  • the on or off control of the cold cathode emission can also be realized, the pulse emission of the electron beam can be easily realized, the response speed is fast, and the service life is long.
  • the X-ray tube cooperates with the electric field applied by the high voltage generator and grounds the anode, so that the X-ray source can easily realize the high-frequency pulse emission of the electron beam, the response speed is fast, the service life is long, and the safety is high, that is, the overcoming
  • the inherent shortcomings of the existing hot filament X-ray source improve the overall performance of the X-ray source, and can better meet the practical application requirements such as medical testing.
  • FIGS. 1A-1B are schematic diagrams showing an optional structure of an X-ray source for medical detection according to an embodiment of the present invention
  • 2A-2C are SEM photographs and field emission characteristics of an optional diode LaB6 cone-cone field emission array according to an embodiment of the present invention.
  • FIG. 3A-3C are SEM photographs and field emission characteristics of an optional triode LaB6 cone-cone field emission array according to an embodiment of the present invention
  • FIG. 4 is an example of an X-ray tube anode model according to an embodiment of the present invention
  • 5 is an example of a curve of a maximum withstand current of an anode according to a thickness of a tungsten alloy sheet according to an embodiment of the present invention
  • FIG. 6 is an example of a relationship between an incident angle of an electron beam (or a target tilt angle) and a photon yield according to an embodiment of the present invention
  • FIG. 7 is a schematic diagram of an imaging principle of an X-ray tube in a medical examination such as a CT scan of a head according to an embodiment of the present invention
  • FIG. 8 is a different angle from a target surface when the target surface angle is 5 degrees according to an embodiment of the present invention;
  • An example of a distribution curve of the photon areal density;
  • FIG. 9 is an example of a distribution curve of the number of X-photons in an exit surface perpendicular to the incident direction of the electron beam at different target inclination angles according to an embodiment of the present invention;
  • FIG. 10 is an embodiment of the present invention.
  • FIG. 1A is a schematic diagram showing an optional structure of an X-ray source for medical detection according to an embodiment of the present invention.
  • the X-ray source for medical detection provided by the embodiment of the present invention includes: a X-ray tube and a high voltage power supply (HVPS).
  • HVPS high voltage power supply
  • the X-ray tube comprises: an anode 1, a cold cathode 2, a grid 3 and a bulb 4; the envelope 4 is for supporting the anode 1, the cold cathode 2 and the grid 3, and makes the anode 1, the cold cathode 2 and
  • the vacuum working environment of the gate 3 is insulated from the outside, the anode 1 is grounded, and the gate 3 is located between the anode 1 and the cold cathode 3.
  • a high voltage generator is used to provide a first electric field between the cold cathode 2 and the gate 3 to cause electrons to be emitted from the cold cathode 2 field, and a second electric field between the gate 3 and the anode 1 to accelerate electrons emitted by the cold cathode 2,
  • the anode 1 is bombarded to generate X-rays.
  • the X-ray source provided in this embodiment adopts a cold cathode, and a gate is arranged between the anode and the cold cathode. Due to the protection of the grid, most of the air ions cannot directly collide with the cold cathode, as shown in FIG. 1B, thereby being able to reduce The probability that the cold cathode is damaged by radiation; In addition, by controlling the voltage applied to the gate, the on or off control of the cold cathode emission can be realized, and the pulse emission of the electron beam can be easily realized, and the response speed is fast and the service life is long.
  • the gate may be a metal mesh gate made of a metal mesh.
  • the X-ray source provided by the embodiment provides a cold electric field to emit electrons through a first electric field provided between the cold cathode and the gate by the high voltage generator, and is provided between the gate and the anode through the high voltage generator.
  • the two electric fields accelerate the electrons emitted from the cold cathode field, causing them to bombard the anode to generate X-rays. Since the cold cathode X-ray tube with the gate is used, the electric field applied by the ⁇ combined high voltage generator acts, and the anode is grounded, so that the X-ray source can easily realize the high-frequency pulse emission of the electron beam, and the response speed is fast and the service life is long. It has high safety and overcomes the shortcomings inherent in the existing hot filament X-ray source, which can better meet the practical application requirements such as medical testing.
  • the relevant parameters of the first electric field applied between the cold cathode and the gate of the high voltage generator are, for example, a DC voltage of 500 ⁇ - 1000 ⁇ , a power higher than 50 W, and an operating frequency of 300 Hz - 3000 Hz, The pulse duty cycle is 20% -80%.
  • This scheme can effectively improve the beam intensity of the cold cathode field emission.
  • the parameters of the second electric field applied between the gate and the anode by the high voltage generator are, for example, a DC voltage higher than 140 kV, a tube current of 2 mA-16 mA, and a power higher than 2000 W.
  • the anode is grounded. The scheme can effectively accelerate the electrons emitted from the cold cathode field, causing it to bombard the anode to generate more X-photons and improve the X-ray intensity.
  • the cold cathode comprises: a substrate and a carbon nanotube emission array formed on the substrate.
  • a carbon nanotube cathode made of carbon nanotubes as a cathode material is a cold cathode compared to the hot filament cathode of the prior art.
  • the principle of X-ray generation by X-ray tube based on carbon nanotubes is as follows: The cathode of the carbon nanotube undergoes field emission under the action of the first electric field to generate electrons, and the electrons accelerate the bombardment of the anode under the second electric field, thereby generating X-rays.
  • Carbon nanotubes have a very low field emission on-field strength (1-3 V/ ⁇ ) and a high field emission current density ( ⁇ lA/cm 2 ), which can be stable for a long period of time under normal high vacuum ( ⁇ 10 ⁇ ).
  • the response time is on the order of nanoseconds, with continuous emission for 10,000 hours, the beam intensity is only reduced by 5%. Therefore, the scheme uses carbon nanotubes to develop an X-ray tube, which can easily realize high-frequency pulse emission of an electron beam, has a fast response speed and a long service life, thereby overcoming the disadvantages inherent in the existing hot filament X-ray source, and To meet the practical needs of medical testing and other practical applications.
  • the cold cathode comprises a substrate and a LaB6 cone field emission array formed on the substrate.
  • LaB6 nanomaterials have the best physical and chemical properties and electron emission properties.
  • a large number of experimental results show that the LaB6 nanomaterials have a work function of 2.4-2.8 eV, which is much lower than that of a pure tungsten cathode of 4.52 eV. It has the advantages of strong anti-poisoning ability, strong anti-ion bombardment ability, stable chemical property and long service life, which can meet the material selection requirements of field emission cathode.
  • LaB6 cone field emission array in the X-ray tube operates under vacuum, an absolute vacuum cannot be achieved in the X-ray tube, and a small amount of air molecules still exist. After being ionized by the high-energy electron beam, these air molecules will accelerate toward the cathode under the strong electric field in the tube, and may bombard the cathode, thereby causing radiation damage of the cathode. Because LaB6 nanomaterials have strong resistance to ion bombardment and high chemical stability, X-ray tubes based on LaB6 nanomaterial field emission have longer working life and stable and reliable performance compared with other X-ray tubes.
  • the solution uses LaB6 nanomaterials as the tip material of the X-ray Field Emission Arrays (FEAs) cathode, and the LaB6 cone-field emission array thus produced can generate a large amount of electrons generated by field emission under the electric field.
  • FEAs X-ray Field Emission Arrays
  • the X-rays generated by the electron bombardment anode are very stable, so that the X-rays generated by these electron bombardment anodes are consistent, which is beneficial to improve the definition and resolution of X-ray imaging, and reduce the radiation dose to the measured object.
  • the LaB6 cone field emission array comprises: a diode LaB6 cone field emission array
  • the diode LaB6 cone field emission array comprises: a silicon tip diode array and a LaB6 nano material film layer covering the surface of the silicon tip cone .
  • a Scanning Electron Microscope (SEM) photograph of an optional diode LaB6 tip cone field emission array is shown in Figure 2A and Figure 2B.
  • the field emission characteristics are shown in Figure 2C.
  • SEM Scanning Electron Microscope
  • the X-ray tube using the diode LaB6 cone-cone field emission array as the cathode has a lower threshold electric field, that is, the applied electric field required for achieving stable emission of X-rays is small, and can be used at ordinary high vacuum ( ⁇ 10 5 Pa).
  • Long-term stable operation it is easy to achieve high-frequency pulse emission of electron beam, fast response, long service life, and is conducive to reducing power consumption, reducing radiation dose to the measured object, environmental protection, health and other advantages, but more To meet the practical needs of medical testing and other practical applications.
  • the LaB6 conical field emission array comprises: a triode LaB6 conical field emission array
  • the triode LaB6 cone field emission array comprises: an array of cavities formed on the 3 ⁇ 4 ⁇ 4 base, distributed in each cavity
  • An SEM photograph of an optional triode LaB6 tip cone field emission array prepared by a conventional process such as the Spindt method is shown in Figure 3A.
  • An optional triode LaB6 tip prepared by mask oxidation technique (LOCOS method) is used.
  • the SEM image of the cone field emission array is shown in Fig. 3B.
  • the field emission characteristics are shown in Fig. 3C.
  • the emission current density of the X-ray tube is 0.6 A/cm 2 , which is equivalent to the average emission current of the single tip cone. 0.24 ⁇ . It can be seen that the X-ray tube using the triode LaB6 cone-cone field emission array as the cathode has a low field emission on-field strength and a high field emission current density, and can work stably for a long time under ordinary high vacuum ( ⁇ 10 5 Pa).
  • the high-frequency pulse emission of the electron beam can be easily realized, the response speed is fast, the service life is long, and the power consumption is reduced, the radiation dose to the measured object is reduced, and the environmental protection and health are better, which can better meet the medical detection. Such as the actual application needs.
  • the anode 1 includes an anode body 11 and a target surface 12 disposed on the anode body 11 .
  • the anode beam can be effectively increased by a reasonable selection of the anode material.
  • the anode body is a copper anode body
  • the target surface is a tungsten alloy target surface.
  • electrons emitted by the cold cathode are accelerated by an electric field and then impinge on the anode target to generate X-rays, wherein more than 99% of the energy of the electron beam is converted into heat deposited in the anode, and less than about 1% of the energy is converted into X-rays.
  • the X-ray tube can be designed by using a fixed anode scheme, that is, the anode in the X-ray tube is a fixed anode.
  • tungsten has a high melting point but poor thermal conductivity; copper has good thermal conductivity but low melting point. Although graphite has higher melting point and specific heat than tungsten and copper, its atomic number is low and X-ray generation efficiency is low. Therefore, copper can be used as the anode body to take advantage of its good thermal conductivity, and a tungsten alloy sheet is used as a target surface to utilize its high melting point performance. Due to the inconsistent properties of copper and tungsten, the thickness of the tungsten alloy sheet is a key parameter for anode design.
  • the thickness of the tungsten alloy sheet needs to be selected to an optimum value.
  • thermal analysis software can be used to simulate the temperature rise curve of the tungsten alloy sheet with different thicknesses under different electron beam pulse bombardment, tungsten alloy sheet and adjacent metal copper, and heat at the anode. In the transfer process, the relationship between material thickness, electron beam intensity and temperature is studied.
  • the physical model of the anode is shown in Figure 4:
  • the copper anode body has a geometry of 04Ox5Omm, the target surface material is tungsten, the tungsten alloy sheet has a diameter of 010mm, the focal diameter is 01mm, and the thickness of the tungsten alloy sheet ranges from 20 ⁇ to 2 ⁇ , X.
  • the tube voltage is 140kV and the current range is 2mA ⁇ 10mA.
  • the finite element model of the X-ray tube anode can be established using ANSYS 12 to perform thermal analysis calculations.
  • the temperature distribution on the anode can be calculated by changing the thickness and current intensity of the tungsten alloy sheet.
  • the electron beam is struck on the surface of tungsten with a focal diameter of 01 mm.
  • the average depth of electrons entering the surface of tungsten is 5 ⁇ m.
  • the electrons generate heat within this tiny volume.
  • E is the radiation force, and the unit is W/m 2 ;
  • s is the emissivity of the object;
  • c is the blackbody emissivity, 5.67 W / (m 2 .K 4 );
  • ⁇ is the surface temperature of the object. According to the electron beam focus temperature of 3300 degrees Celsius, the other surface temperature is 400. C estimates, then the radiant power of the anode is:
  • the input power of the anode is 1050W, then ⁇ ⁇ Oi ⁇ SS, the radiated power accounts for a small proportion of the input power, which can be ignored.
  • the following is a simulation result that ignores the radiation heat dissipation and the conduction heat dissipation of the insulating oil.
  • the maximum time for completing a CT scan is 30s, so the X-ray tube must be able to work continuously for 30s during scanning. Based on this, the optimal tungsten alloy sheet thickness and the maximum constant current that can be tolerated are calculated. value. As can be seen from Fig.
  • the maximum withstand current is 7.5 mA when the thickness of the tungsten alloy sheet is 400 to 500 ⁇ m.
  • the copper will melt first, and on the right, the tungsten alloy sheet will melt first.
  • the maximum pulse current that a tungsten alloy sheet of the same thickness can withstand at different duty cycles increases as the duty cycle decreases.
  • the embodiment of the present invention will select a tungsten alloy target surface having a thickness of 400-500 um, for example, preferably 0.5 mm, as a preferred thickness value of the tungsten alloy sheet.
  • the anode 1 of the X-ray tube includes an anode body 11 and a target surface 12.
  • the target surface 12 is formed with a predetermined target tilt angle ⁇ with respect to the reference direction, and the reference direction is perpendicular to the electron incident direction, as shown in FIG.
  • the target tilt angle ⁇ is a key parameter that directly affects the light yield, effective focus size, heat distribution and transfer of the X-ray tube.
  • Monte Carlo method can be used to simulate the calculation.
  • EGS software was used to simulate lxlO 7 140keV electrons bombarding tungsten targets with different dip angles, and the spatial distribution of light yield and photons was counted.
  • the relationship between target dip angle and photon yield is shown in Figure 6.
  • the smaller the target tilt angle the higher the X-ray photo yield.
  • the smaller the target angle the better, which requires careful analysis.
  • the X-ray photons in the fan beam which are approximately perpendicular to the incident direction of the electron beam are finally used. This part of the X-ray photo is actually contributing to the CT system (as shown in Fig. 7), so this angle The more X-rays in the range, the better.
  • the figure below shows the photon areal density at an angle different from the target surface at a target angle of 5 degrees.
  • the areal density of photons becomes smaller and smaller, and the number of X-photons that can be used for imaging becomes less and less. Therefore, although the total photon yield at a target angle of 5 degrees is high, the photon surface density at an angle of 85 degrees from the target surface is low.
  • the number of X-photons in the exit plane perpendicular to the incident direction of the electron beam at different target inclination angles is counted.
  • the statistical results are shown in Fig. 9. As can be seen from Fig.
  • the resolution of the tomographic image is the effective focus of the X-ray tube, not the actual focus. Assuming that the electron beams are incident in parallel, the relationship between the actual focus size L and the projected effective focus size d is as follows:
  • the size d of the effective focus can be controlled by reducing the target tilt angle ⁇ . If the density of the cross-sectional area of the incident electron beam cannot be increased, it can be seen from the following equation that increasing the electron beam width h by decreasing the target tilt angle ⁇ may increase the total number of imageable X-photons.
  • the target tilt angle is preferably 11 degrees, as shown in Fig. 1A.
  • the total length of the X-ray tube in the above embodiment is less than or equal to 120 mm, and/or the total weight of the X-ray source is less than 25 kg, so as to fully ensure the compact shape of the X-ray tube, which can be easily carried, and is convenient for the ship.
  • Special environments such as trucks, vehicles, and battlefield hospitals.
  • the maximum diameter in the above embodiment is less than or equal to 60 mm. Further preferably, the distance between the anode and the tip of the tip of the cathode in the above embodiment is less than or equal to 10 um. This ensures excellent performance of the X-ray tube.
  • the present invention also provides a mobile CT scanner comprising the X-ray source for medical detection provided by any of the above embodiments, wherein X-rays are generated by the X-ray source to perform body parts such as the brain. Medical testing.
  • the foregoing program may be stored in a computer readable storage medium, and the program is executed when executed.
  • the foregoing storage device includes the following steps:
  • the foregoing storage medium includes: a Read-Only Memory (ROM), a Random Access Memory (RAM), a magnetic disk, or an optical disk, and the like.
  • the medium of the program code includes: a Read-Only Memory (ROM), a Random Access Memory (RAM), a magnetic disk, or an optical disk, and the like.

Abstract

An X-ray source for a medical detection, and a movable CT scanner. The X-ray source for a medical detection comprises an X-ray tube and a high-voltage generator. The X-ray tube comprises an anode (1), a cold cathode (2), a gate (3), and a tube case (4). The tube case (4) is used for supporting the anode (1), the cold cathode (2) and the gate (3), and enabling a vacuum working environment of the anode (1), the cold cathode (2) and the gate (3) to be insulated from the outside. The anode (1) is grounded. The high-voltage generator is used for providing a first electric field between the cold cathode (2) and the gate (3) to enable the cold cathode (2) to emit electrons in a field mode, and is used for providing a second electric field between the gate (3) and the anode (1) to accelerate the electrons emitted by the cold cathode (2), thereby enabling the electrons to bombard the anode (1) to generate X-rays. The X-ray source can implement high-frequency pulse transmission of electron beams very easily, the response speed is fast, the service life is long, and security is high, that is, the overall performance of the X-ray source is improved, and practical application demands such as a medical detection can be better met.

Description

医学检测用 X射线源及移动 CT扫描仪  X-ray source for medical testing and mobile CT scanner
本发明要求 2013年 4月 27日向中国国家知识产^ ^提交的、申请号 为 201310151634.2、 名称为 "医学检测用 X射线源及移动 CT扫描仪" 的中国专利申请的优先权。 The present invention claims the priority of a Chinese patent application filed on April 27, 2013, to the Chinese National Intellectual Property No. 201310151634.2, entitled "X-ray source for medical testing and mobile CT scanner".
技术领域 Technical field
本发明涉及医疗器械领域，特别涉及一种医学检测用 X射线源及 移动 CT扫描仪。  The present invention relates to the field of medical devices, and in particular to an X-ray source for medical detection and a mobile CT scanner.
背景技术 Background technique
随着医学科学技术的发展，涌现出了各种各样的医用计算机断层扫描 仪（ Computer tomography; CT )设备。 其中 X射线管为一种小型医用 CT设备关键部件。 阴极是 X射线管的核心部件，直接决定着 X射线管的 性能、 成像的质量如分辨率和对比度， 以及整机的工作效率。  With the development of medical science and technology, a variety of medical computer tomography (CT) devices have emerged. The X-ray tube is a key component of a small medical CT device. The cathode is the core component of the X-ray tube, which directly determines the performance of the X-ray tube, the quality of the image such as resolution and contrast, and the efficiency of the whole machine.
现有技术中 X射线管通常是基于钨（W )丝热发射的 X射线管， 即 采用钨（W )丝制作 X射线管的阴极， 其工作原理是钨 ( W )丝加热至 其工作温度时发射电子， 热发射的电子轰击阳极， 从而产生 X射线。  In the prior art, the X-ray tube is usually an X-ray tube based on tungsten (W) wire thermal emission, that is, a cathode made of tungsten (W) wire to make an X-ray tube, and the working principle is that the tungsten (W) wire is heated to its working temperature. When electrons are emitted, the electrons that are emitted by the heat bombard the anode, thereby generating X-rays.
现有技术基于钨（W )丝热发射的 X射线管中至少存在如下缺点： 现 有的 X射线管中的阴极采用的钨的电子逸出功高（(|)W=4.52eV )，发射电流 密度小， 纯钨材料在 2200 "C时， 其热发射电流密度只有 0.3A/cm2。 如果 要想获得较大的总发射电流，通常采用提高阴极温度，但是提高阴极温度 会使阴极材料的蒸发率增加， 阴极材料蒸发会使钨丝变细， 变细后的钨丝 阴极又会使阴极温度升高， 阴极蒸发加剧， 从而形成恶性循环； 此外， 被 蒸发的钨阴极材料会沉积在管壳上，形成连续或断续的钨导电薄膜，破坏 了 X射线管的绝缘强度， 使管压降低、 管子报废， 降低了 X射线管的寿命; 同时， 这种钨导电薄膜还阻挡了输出窗口的 X射线强度， 降低了成像灵敏 度。 因此现有技术的基于钨（W )丝热发射的 X射线管的整体性能较差， 迫切需要研究一种新型的冷阴极 X射线管以代替现有基于热钨（W )丝的 X射线管。 发明内容 The prior art X-ray tube based on tungsten (W) wire thermal emission has at least the following disadvantages: The cathode in the existing X-ray tube uses tungsten with high electron emission work ((|) W = 4.52 eV), emission The current density is small. The pure tungsten material has a thermal emission current density of only 0.3 A/cm 2 at 2200 ° C. If a larger total emission current is to be obtained, the cathode temperature is usually increased, but the cathode temperature is increased to cause the cathode material. The evaporation rate increases, the evaporation of the cathode material causes the tungsten filament to become thinner, and the thinned tungsten cathode increases the cathode temperature and the cathode evaporation is intensified, thereby forming a vicious cycle; in addition, the evaporated tungsten cathode material is deposited on the On the shell, a continuous or intermittent tungsten conductive film is formed, which destroys the insulation strength of the X-ray tube, reduces the tube pressure, and the tube is scrapped, thereby reducing the life of the X-ray tube; at the same time, the tungsten conductive film blocks the output. The X-ray intensity of the window reduces the imaging sensitivity. Therefore, the overall performance of the prior art X-ray tube based on tungsten (W) wire thermal emission is poor, and it is urgent to study a new type of cold cathode X-ray tube. Instead of the existing X-ray tube based on hot tungsten (W) wire. Summary of the invention
在下文中给出关于本发明的简要概述，以便提供关于本发明的某些方 面的基本理解。应当理解， 这个概述并不是关于本发明的穷举性概述。 它 并不是意图确定本发明的关键或重要部分， 也不是意图限定本发明的范 围。其目的仅仅是以简化的形式给出某些概念， 以此作为稍后论述的更详 细描述的前序。  A brief summary of the invention is set forth below in order to provide a basic understanding of certain aspects of the invention. It should be understood that this summary is not an exhaustive overview of the invention. It is not intended to identify key or critical aspects of the invention, and is not intended to limit the scope of the invention. Its purpose is to present some concepts in a simplified form as a pre-
本发明提供一种医学检测用 X射线源及移动 CT扫描仪， 用以提 高 X射线源的整体性能， 可满足医学检测等应用需求。  The invention provides an X-ray source for medical detection and a mobile CT scanner for improving the overall performance of the X-ray source, and can meet the application requirements of medical testing and the like.
一方面， 本发明了提供一种医学检测用 X射线源， 包括： In one aspect, the invention provides an X-ray source for medical detection, comprising:
X射线管， 包括阳极、 冷阴极、 栅极和管壳； 所述管壳用于支撑所述 阳极、 冷阴极和栅极， 并使得所述阳极、冷阴极和栅极的真空工作环境与 外界绝缘， 所述阳极接地； An X-ray tube comprising an anode, a cold cathode, a grid and a casing; the casing for supporting the anode, the cold cathode and the grid, and the vacuum working environment of the anode, the cold cathode and the gate to the outside Insulation, the anode is grounded;
高压发生器，用于在所述冷阴极和所述栅极之间提供第一电场使所述 冷阴极场发射电子，以及在所述栅极和所述阳极之间提供第二电场以加速 所述冷阴 射的电子， 使之轰击所述阳极来产生 X射线。  a high voltage generator for providing a first electric field between the cold cathode and the gate to cause electrons to be emitted by the cold cathode field, and a second electric field between the gate and the anode to accelerate The cold-exposed electrons are bombarded with the anode to produce X-rays.
另一方面， 本发明还提供了一种移动 CT扫描仪， 包括如上所述的医 学检测用 X射线源。  In another aspect, the invention also provides a mobile CT scanner comprising an X-ray source for medical testing as described above.
本发明提供的技术方案中 X射线源采用冷阴极， 并在阳极和冷阴极 之间设有栅极， 由于栅极的保护， 大部分空气离子无法直接撞击冷阴极， 因此能够降低冷阴极被辐射损伤的概率； 此外，通过对栅极施加的电压控 制，还可实现冷阴极发射的导通或截止控制，可以很容易实现电子束的脉 冲发射， 响应速度快， 使用寿命长。 当采用脉冲曝光成像等方式工作时， 可以显著降低采样的投影角度数和辐射剂量， 并能有效抑制旋转伪影，进 而更好满足医学检测等实际应用需求。此外， 由于采用具有栅极的冷阴极 In the technical solution provided by the invention, the X-ray source adopts a cold cathode, and a gate is arranged between the anode and the cold cathode. Due to the protection of the grid, most of the air ions cannot directly collide with the cold cathode, thereby reducing the cold cathode being irradiated. The probability of damage; In addition, by the voltage control applied to the gate, the on or off control of the cold cathode emission can also be realized, the pulse emission of the electron beam can be easily realized, the response speed is fast, and the service life is long. When working by pulse exposure imaging, the number of projection angles and radiation dose can be significantly reduced, and the rotation artifacts can be effectively suppressed, thereby better meeting the practical application requirements such as medical detection. In addition, due to the use of a cold cathode with a gate
X射线管， 配合高压发生器施加的电场作用， 并将阳极接地， 使得 X射 线源可以很容易实现电子束的高频脉冲发射， 响应速度快， 使用寿命长， 安全性较高， 即克服了现有热灯丝 X射线源所固有的缺点， 改善了 X射 线源的整体性能， 可更好满足医学检测等实际应用需求。 The X-ray tube cooperates with the electric field applied by the high voltage generator and grounds the anode, so that the X-ray source can easily realize the high-frequency pulse emission of the electron beam, the response speed is fast, the service life is long, and the safety is high, that is, the overcoming The inherent shortcomings of the existing hot filament X-ray source improve the overall performance of the X-ray source, and can better meet the practical application requirements such as medical testing.
附图说明 为了更清楚地说明本发明实施例或现有技术中的技术方案，下面将对 实施例或现有技术描述中所需务 ί吏用的附图作简单地介绍， 显而易见地， 下面描述中的附图仅仅是本发明的一些实施例，对于本领域普通技术人员 来讲，在不付出创造性劳动的前提下，还可以根据这些附图获得其他的附 图。 DRAWINGS In order to more clearly illustrate the embodiments of the present invention or the technical solutions in the prior art, the drawings of the embodiments or the description of the prior art will be briefly described below, and obviously, in the following description The drawings are only some of the embodiments of the present invention, and other drawings may be obtained from those skilled in the art without departing from the drawings.
图 1A-图 1B为本发明实施例提供的医学检测用 X射线源的可选结构 示意图；  1A-1B are schematic diagrams showing an optional structure of an X-ray source for medical detection according to an embodiment of the present invention;
图 2Α-图 2C为本发明实施例提供的一种可选的二极管 LaB6尖锥场 发射阵列的 SEM照片、 场发射特性；  2A-2C are SEM photographs and field emission characteristics of an optional diode LaB6 cone-cone field emission array according to an embodiment of the present invention;
图 3A-图 3C为本发明实施例提供的一种可选的三极管 LaB6尖锥场 发射阵列的 SEM照片、 场发射特性； 图 4为本发明实施例提供的一种 X射线管阳极模型示例； 图 5 为本发明实施例提供的阳极最大耐受电流随钨合金片厚度变化 曲线示例； 图 6为本发明实施例提供的电子束入射角（或者靶面倾角）与光子产 额的关系曲线示例； 图 7为本发明实施例提供的 X射线管在如头部 CT扫描成像等医学检 测的成像原理示意图； 图 8为本发明实施例提供的靶面倾角 5度时， 与靶面不同夹角的光子 面密度的分布曲线示例； 图 9 为本发明实施例提供的不同靶面倾角下与电子束入射方向垂直 的出射面内 X光子的数量的分布曲线示例； 图 10为本发明实施例提供的靶面倾角与可用于成像的 X光子数的关 系曲线示例。 具体实施方式  3A-3C are SEM photographs and field emission characteristics of an optional triode LaB6 cone-cone field emission array according to an embodiment of the present invention; FIG. 4 is an example of an X-ray tube anode model according to an embodiment of the present invention; 5 is an example of a curve of a maximum withstand current of an anode according to a thickness of a tungsten alloy sheet according to an embodiment of the present invention; FIG. 6 is an example of a relationship between an incident angle of an electron beam (or a target tilt angle) and a photon yield according to an embodiment of the present invention; FIG. 7 is a schematic diagram of an imaging principle of an X-ray tube in a medical examination such as a CT scan of a head according to an embodiment of the present invention; FIG. 8 is a different angle from a target surface when the target surface angle is 5 degrees according to an embodiment of the present invention; An example of a distribution curve of the photon areal density; FIG. 9 is an example of a distribution curve of the number of X-photons in an exit surface perpendicular to the incident direction of the electron beam at different target inclination angles according to an embodiment of the present invention; FIG. 10 is an embodiment of the present invention. An example of the relationship between the target tilt angle and the number of X-photons available for imaging. detailed description
为使本发明实施例的目的、技术方案和优点更加清楚， 下面将结合本 发明实施例中的附图，对本发明实施例中的技术方案进行清楚、完整地描 述，显然，所描述的实施例是本发明一部分实施例，而不是全部的实施例。 在本发明的一个附图或一种实施方式中描述的元素和特征可以与一个或 更多个其它附图或实施方式中示出的元素和特征相结合。应当注意， 为了 清楚的目的， 附图和说明中省略了与本发明无关的、本领域普通技术人员 已知的部件和处理的表示和描述。基于本发明中的实施例，本领域普通技 术人员在没有付出创造性劳动的前提下所获得的所有其他实施例，都属于 本发明保护的范围。 The technical solutions in the embodiments of the present invention are clearly and completely described in conjunction with the drawings in the embodiments of the present invention. It is a partial embodiment of the invention, and not all of the embodiments. The elements and features described in one of the figures or an embodiment of the invention may be associated with one or The elements and features shown in more other figures or embodiments are combined. It should be noted that, for the sake of clarity, representations and descriptions of components and processes known to those of ordinary skill in the art that are not related to the present invention are omitted from the drawings and the description. All other embodiments obtained by a person of ordinary skill in the art based on the embodiments of the present invention without departing from the inventive scope are the scope of the present invention.
图 1A为本发明实施例提供的医学检测用 X射线源的可选结构示意 图。 如图 1A所示， 本发明实施例提供的医学检测用 X射线源包括： X射 线管和高压发生器（ High Voltage Power Supply, HVPS )。  FIG. 1A is a schematic diagram showing an optional structure of an X-ray source for medical detection according to an embodiment of the present invention. As shown in FIG. 1A, the X-ray source for medical detection provided by the embodiment of the present invention includes: a X-ray tube and a high voltage power supply (HVPS).
X射线管包括： 阳极 1、 冷阴极 2、 栅极 3和管壳 4; 所述管壳 4用 于支撑所述阳极 1、 冷阴极 2和栅极 3， 并使得阳极 1、 冷阴极 2和栅极 3 的真空工作环境与外界绝缘， 阳极 1接地， 栅极 3位于阳极 1和冷阴极 3 之间。  The X-ray tube comprises: an anode 1, a cold cathode 2, a grid 3 and a bulb 4; the envelope 4 is for supporting the anode 1, the cold cathode 2 and the grid 3, and makes the anode 1, the cold cathode 2 and The vacuum working environment of the gate 3 is insulated from the outside, the anode 1 is grounded, and the gate 3 is located between the anode 1 and the cold cathode 3.
高压发生器用于在冷阴极 2和栅极 3之间提供第一电场使冷阴极 2 场发射电子， 以及在栅极 3和阳极 1之间提供第二电场以加速冷阴极 2 发射的电子， 使之轰击阳极 1来产生 X射线。  A high voltage generator is used to provide a first electric field between the cold cathode 2 and the gate 3 to cause electrons to be emitted from the cold cathode 2 field, and a second electric field between the gate 3 and the anode 1 to accelerate electrons emitted by the cold cathode 2, The anode 1 is bombarded to generate X-rays.
本实施例提供的 X射线源采用冷阴极， 并在阳极和冷阴极之间设有 栅极， 由于栅极的保护， 大部分空气离子无法直接撞击冷阴极， 如图 1B 所示， 因此能够降低冷阴极被辐射损伤的概率； 此外， 通过对栅极施加的 电压控制，还可实现冷阴极发射的导通或截止控制，可以艮容易实现电子 束的脉冲发射， 响应速度快， 使用寿命长。 当采用脉冲曝光成像等方式工 作时，可以显著降低采样的投影角度数和辐射剂量， 并能有效抑制旋转伪 影，进而更好满足医学检测等实际应用需求。为了对阴极形成更好的保护， 可选的， 栅极可为采用金属网制成的金属网栅极。  The X-ray source provided in this embodiment adopts a cold cathode, and a gate is arranged between the anode and the cold cathode. Due to the protection of the grid, most of the air ions cannot directly collide with the cold cathode, as shown in FIG. 1B, thereby being able to reduce The probability that the cold cathode is damaged by radiation; In addition, by controlling the voltage applied to the gate, the on or off control of the cold cathode emission can be realized, and the pulse emission of the electron beam can be easily realized, and the response speed is fast and the service life is long. When working by pulse exposure imaging, the number of projection angles and radiation dose of the sample can be significantly reduced, and the rotation artifact can be effectively suppressed, thereby better meeting the practical application requirements such as medical detection. In order to form a better protection for the cathode, alternatively, the gate may be a metal mesh gate made of a metal mesh.
进一步的， 本实施例提供的 X射线源通过高压发生器在冷阴极和栅 极之间提供的第一电场使冷阴极场发射电子，并通过高压发生器在栅极和 阳极之间提供的第二电场，加速冷阴极场发射出的电子，使之轰击阳极来 产生 X射线。 由于采用具有栅极的冷阴极 X射线管， δ合高压发生器施 加的电场作用， 并将阳极接地， 使得 X射线源可以很容易实现电子束的 高频脉冲发射， 响应速度快， 使用寿命长， 安全性较高， 克服了现有热灯 丝 X射线源所固有的缺点， 可更好满足医学检测等实际应用需求。  Further, the X-ray source provided by the embodiment provides a cold electric field to emit electrons through a first electric field provided between the cold cathode and the gate by the high voltage generator, and is provided between the gate and the anode through the high voltage generator. The two electric fields accelerate the electrons emitted from the cold cathode field, causing them to bombard the anode to generate X-rays. Since the cold cathode X-ray tube with the gate is used, the electric field applied by the δ combined high voltage generator acts, and the anode is grounded, so that the X-ray source can easily realize the high-frequency pulse emission of the electron beam, and the response speed is fast and the service life is long. It has high safety and overcomes the shortcomings inherent in the existing hot filament X-ray source, which can better meet the practical application requirements such as medical testing.
可选的，高压发生器在冷阴极和栅极之间施加的第一电场的相关参数 例如： 500ν-1000ν的直流电压，高于 50w的功率，工作频率 300Hz-3000Hz， 脉冲占空比为 20% -80%。 该方案可有效提高冷阴极场发射的束流强度。 可选的，高压发生器在栅极和阳极之间施加的第二电场的相关参数例 如： 高于 140kv的直流电压， 2mA-16mA的管电流， 高于 2000W的功率。 阳极接地。该方案可有效加速冷阴极场发射出的电子，使之轰击阳极产生 更多的 X光子， 提高 X射线强度。 Optionally, the relevant parameters of the first electric field applied between the cold cathode and the gate of the high voltage generator are, for example, a DC voltage of 500 ν - 1000 ν, a power higher than 50 W, and an operating frequency of 300 Hz - 3000 Hz, The pulse duty cycle is 20% -80%. This scheme can effectively improve the beam intensity of the cold cathode field emission. Optionally, the parameters of the second electric field applied between the gate and the anode by the high voltage generator are, for example, a DC voltage higher than 140 kV, a tube current of 2 mA-16 mA, and a power higher than 2000 W. The anode is grounded. The scheme can effectively accelerate the electrons emitted from the cold cathode field, causing it to bombard the anode to generate more X-photons and improve the X-ray intensity.
可选的， 冷阴极包括： 基板以及形成于所述基板上的碳纳米管发射阵 列。将碳纳米管作为阴极材料制成的碳纳米管阴极，相对现有技术中的热 灯丝阴极而言是一种冷阴极。 基于碳纳米管的 X射线管产生 X射线的原理 是： 碳纳米管阴极在第一电场的作用下发生场致发射产生电子， 电子在第 二电场下加速轰击阳极， 从而产生 X射线。 碳纳米管具有很低的场发射 开启电场强度 (1-3 V/μιη)和很高的场发射电流密度（~ lA/cm2)， 可在普通 高真空度（~ 10 ^ )下长期稳定工作， 响应时间为纳秒量级， 连续发射 10000 小时， 束流强度只降低 5%。 因此， 该方案采用碳纳米管研制 X射 线管可以很容易实现电子束的高频脉冲发射， 响应速度快， 使用寿命长， 由此克服了现有热灯丝 X射线源所固有的缺点， 可更好满足医学检测等实 际应用需求。 Optionally, the cold cathode comprises: a substrate and a carbon nanotube emission array formed on the substrate. A carbon nanotube cathode made of carbon nanotubes as a cathode material is a cold cathode compared to the hot filament cathode of the prior art. The principle of X-ray generation by X-ray tube based on carbon nanotubes is as follows: The cathode of the carbon nanotube undergoes field emission under the action of the first electric field to generate electrons, and the electrons accelerate the bombardment of the anode under the second electric field, thereby generating X-rays. Carbon nanotubes have a very low field emission on-field strength (1-3 V/μιη) and a high field emission current density (~ lA/cm 2 ), which can be stable for a long period of time under normal high vacuum (~ 10 ^ ). Working, the response time is on the order of nanoseconds, with continuous emission for 10,000 hours, the beam intensity is only reduced by 5%. Therefore, the scheme uses carbon nanotubes to develop an X-ray tube, which can easily realize high-frequency pulse emission of an electron beam, has a fast response speed and a long service life, thereby overcoming the disadvantages inherent in the existing hot filament X-ray source, and To meet the practical needs of medical testing and other practical applications.
或者， 可选的， 冷阴极包括基板以及形成于所述基板上的 LaB6尖锥 场发射阵列。在所有的六硼化物中， LaB6纳米材料具有最优良的理化性 能和电子发射性能， 大量的实验结果表明， LaB6 纳米材料的逸出功为 2.4-2.8eV远低于纯钨阴极为 4.52 eV， 具有抗中毒能力强、 抗离子轰击能 力强、 化学性质稳定、 寿命长等优点， 可满足场发射阴极的选材要求。 此 外， 虽然 X射线管中 LaB6尖锥场发射阵列是在真空状态下工作，但是 X 射线管内无法实现绝对真空，依然存在少量空气分子。这些空气分子被高 能电子束电离后，在管内的强电场作用下会向阴极方向加速，有可能轰击 到阴极， 从而造成阴极的辐射损伤。 由于 LaB6纳米材料抗离子轰击的能 力强，化学稳定性高，故基于 LaB6纳米材料场发射的 X射线管相对其他 X射线管而言， 工作寿命较长， 性能也较为稳定和可靠。 因此， 该方案将 LaB6纳米材料作为 X射线管场发射 ( Field Emission Arrays, FEAs ) 阴 极的尖端材料， 由此制得的 LaB6尖锥场发射阵列在电场作用下可场致发 射产生的大量电子， 提高电子束流强度， 电子轰击阳极产生的 X射线非 常稳定，使得这些电子轰击阳极产生的 X射线具有一致性，有利于提高 X 射线成像的清晰度和分辨率， 降低对被测物的辐射剂量， 并便于实现 X 射线管的小型化，可满足如移动 CT扫描仪等便携式医学检测设备的设计 需求。 Alternatively, optionally, the cold cathode comprises a substrate and a LaB6 cone field emission array formed on the substrate. Among all hexaborides, LaB6 nanomaterials have the best physical and chemical properties and electron emission properties. A large number of experimental results show that the LaB6 nanomaterials have a work function of 2.4-2.8 eV, which is much lower than that of a pure tungsten cathode of 4.52 eV. It has the advantages of strong anti-poisoning ability, strong anti-ion bombardment ability, stable chemical property and long service life, which can meet the material selection requirements of field emission cathode. In addition, although the LaB6 cone field emission array in the X-ray tube operates under vacuum, an absolute vacuum cannot be achieved in the X-ray tube, and a small amount of air molecules still exist. After being ionized by the high-energy electron beam, these air molecules will accelerate toward the cathode under the strong electric field in the tube, and may bombard the cathode, thereby causing radiation damage of the cathode. Because LaB6 nanomaterials have strong resistance to ion bombardment and high chemical stability, X-ray tubes based on LaB6 nanomaterial field emission have longer working life and stable and reliable performance compared with other X-ray tubes. Therefore, the solution uses LaB6 nanomaterials as the tip material of the X-ray Field Emission Arrays (FEAs) cathode, and the LaB6 cone-field emission array thus produced can generate a large amount of electrons generated by field emission under the electric field. Increasing the intensity of the electron beam, the X-rays generated by the electron bombardment anode are very stable, so that the X-rays generated by these electron bombardment anodes are consistent, which is beneficial to improve the definition and resolution of X-ray imaging, and reduce the radiation dose to the measured object. And facilitate the miniaturization of X-ray tubes to meet the design of portable medical testing equipment such as mobile CT scanners. Demand.
例如：所述 LaB6尖锥场发射阵列包括:二极管 LaB6尖锥场发射阵列， 所述二极管 LaB6尖锥场发射阵列包括： 硅尖锥二极管阵列和覆盖在硅尖 锥表面上的 LaB6纳米材料薄膜层。 一种可选的二极管 LaB6 尖锥场发射 阵列的扫描电子显微镜 ( Scanning Electron Microscope, SEM )照片如 图 2A和图 2B所示， 其场发射特性如图 2C所示， 在阳极电压 1500V时， X 射线管的发射电流 32mA， 折合单尖锥的平均发射电流为 Ο. ΙμΑ, 阈值电 场为 8.0ν/μιη。 可见， 采用二极管 LaB6 尖锥场发射阵列作为阴极的 X射 线管具有较低的阈值电场， 即达到 X射线稳定发射时所需的外加电场较 小， 可在普通高真空度（~ 10 5Pa)下长期稳定工作， 可以很容易实现电子 束的高频脉冲发射， 响应速度快， 使用寿命长， 且有利于降低功耗， 减少 对被测物的辐射剂量， 具有环保、健康等优点， 可更好满足医学检测等实 际应用需求。 For example, the LaB6 cone field emission array comprises: a diode LaB6 cone field emission array, and the diode LaB6 cone field emission array comprises: a silicon tip diode array and a LaB6 nano material film layer covering the surface of the silicon tip cone . A Scanning Electron Microscope (SEM) photograph of an optional diode LaB6 tip cone field emission array is shown in Figure 2A and Figure 2B. The field emission characteristics are shown in Figure 2C. At an anode voltage of 1500V, X The emission current of the tube is 32 mA, and the average emission current of the single-cone is Ο. ΙμΑ, and the threshold electric field is 8.0 ν/μιη. It can be seen that the X-ray tube using the diode LaB6 cone-cone field emission array as the cathode has a lower threshold electric field, that is, the applied electric field required for achieving stable emission of X-rays is small, and can be used at ordinary high vacuum (~ 10 5 Pa). Long-term stable operation, it is easy to achieve high-frequency pulse emission of electron beam, fast response, long service life, and is conducive to reducing power consumption, reducing radiation dose to the measured object, environmental protection, health and other advantages, but more To meet the practical needs of medical testing and other practical applications.
又例如：所述 LaB6尖锥场发射阵列包括：三极管 LaB6尖锥场发射阵 列， 所述三极管 LaB6尖锥场发射阵列包括： 、 形成在所¾¾基上的 孔腔阵列、分布在各孔腔中的钼尖锥阵列、 以及覆盖在各钼尖锥表面上的 LaB6纳米材料薄膜层。 采用传统工艺 （如 Spindt法）制备的一种可选的 三极管 LaB6尖锥场发射阵列的 SEM照片如图 3A所示， 采用掩模氧化技 术（LOCOS法）制备的一种可选的三极管 LaB6 尖锥场发射阵列的 SEM 照片如图 3B所示， 其场发射特性如图 3C所示， 在阳极电压 1500V时， X 射线管的发射电流密度为 0.6A/cm2， 折合单尖锥平均发射电流 0.24μΑ。 可见， 采用三极管 LaB6 尖锥场发射阵列作为阴极的 X射线管具有很低的 场发射开启电场强度和很高的场发射电流密度， 可在普通高真空度（~ 10 5Pa)下长期稳定工作， 可以很容易实现电子束的高频脉冲发射， 响应速 度快， 使用寿命长， 且有利于降低功耗， 减少对被测物的辐射剂量， 具有 环保、 健康等优点， 可更好满足医学检测等实际应用需求。 For another example, the LaB6 conical field emission array comprises: a triode LaB6 conical field emission array, and the triode LaB6 cone field emission array comprises: an array of cavities formed on the 3⁄4⁄4 base, distributed in each cavity An array of molybdenum tip cones, and a layer of LaB6 nanomaterial film overlying the surface of each molybdenum tip. An SEM photograph of an optional triode LaB6 tip cone field emission array prepared by a conventional process such as the Spindt method is shown in Figure 3A. An optional triode LaB6 tip prepared by mask oxidation technique (LOCOS method) is used. The SEM image of the cone field emission array is shown in Fig. 3B. The field emission characteristics are shown in Fig. 3C. At an anode voltage of 1500 V, the emission current density of the X-ray tube is 0.6 A/cm 2 , which is equivalent to the average emission current of the single tip cone. 0.24 μΑ. It can be seen that the X-ray tube using the triode LaB6 cone-cone field emission array as the cathode has a low field emission on-field strength and a high field emission current density, and can work stably for a long time under ordinary high vacuum (~ 10 5 Pa). The high-frequency pulse emission of the electron beam can be easily realized, the response speed is fast, the service life is long, and the power consumption is reduced, the radiation dose to the measured object is reduced, and the environmental protection and health are better, which can better meet the medical detection. Such as the actual application needs.
可选的，所述阳极 1包括：阳极体 11以及设于阳极体 11上的靶面 12。 通过合理选择阳极材料， 可有效提高其承受的最大束流强度， 优选的， 所 述阳极体为铜阳极体， 所述靶面为钨合金靶面。 在 X射线管中，冷阴极发射的电子经电场加速后撞击到阳极靶上产生 X射线， 其中电子束 99%以上的能量转化成热量沉积在阳极内， 只有不 到 1%左右的能量转变成 X射线。如果电子在阳极靶上产生的大量热量得 不到及时有效的散失， 阳极靶表面的温升很快， 在很短的时间内， 阳极靶 的表面材料就会融化， 导致 X射线管损坏。 因此， 阳极靶的耐热和散热 性能直接影响了 X射线管的使用。 可选的， 可采用固定阳极方案设计 X射线管， 即 X射线管中阳极为固 定阳极。该方案的优点是有效降低 X射线源的重量和体积，并降低 X射线 管的制造和使用难度。 Optionally, the anode 1 includes an anode body 11 and a target surface 12 disposed on the anode body 11 . The anode beam can be effectively increased by a reasonable selection of the anode material. Preferably, the anode body is a copper anode body, and the target surface is a tungsten alloy target surface. In an X-ray tube, electrons emitted by the cold cathode are accelerated by an electric field and then impinge on the anode target to generate X-rays, wherein more than 99% of the energy of the electron beam is converted into heat deposited in the anode, and less than about 1% of the energy is converted into X-rays. If the large amount of heat generated by the electrons on the anode target is not lost in time, the temperature rise on the surface of the anode target is very fast, and in a short time, the anode target The surface material will melt and cause damage to the X-ray tube. Therefore, the heat resistance and heat dissipation performance of the anode target directly affect the use of the X-ray tube. Optionally, the X-ray tube can be designed by using a fixed anode scheme, that is, the anode in the X-ray tube is a fixed anode. The advantage of this solution is that it effectively reduces the weight and volume of the X-ray source and reduces the difficulty in manufacturing and using the X-ray tube.
X射线管的研制过程中一般涉及到以下几种材料： 表 1 : 材料特性参数 The following materials are generally involved in the development of X-ray tubes: Table 1: Material characteristics parameters
Figure imgf000009_0001
从材料的性能可知， 钨的熔点高， 但是导热性能差； 铜的导热性能好， 但是熔点低。 石墨虽然熔点和比热都比钨、 铜高， 但是其原子序数低， X 射线的产生效率低。 因此， 可以采用铜做阳极体， 以利用其良好的导热性 能， 采用钨合金片做靶面， 以利用其高熔点性能。 由于铜和钨的性能不一致， 钨合金片的厚度是阳极设计的一个关键参 数。 如果钨合金片太厚， 热量来不及传递， 则钨合金片可能先熔化； 如果 钨合金片太薄， 热量立刻传递给铜， 则铜可能先熔化。无论哪种情况出现， 都会影响到 X射线管的正常工作。因此，钨合金片的厚度需要选择最优值。 为了计算钨合金片的最优厚度值， 可使用热分析软件模拟不同厚度的 钨合金片在不同强度的电子束脉冲轰击下， 钨合金片与相邻金属铜的温度 上升曲线， 以及热量在阳极中的传递过程， 研究材料厚度、 电子束流强度 与温度之间的关系。 由于脉冲状态下电子束的热量生成比同强度下恒流状 态下的低， 为了给设计留有余量， 我们主要模拟恒流状态下的参数。 阳极的物理模型如下图 4所示： 铜阳极体的几何尺寸为 04Ox5Omm， 靶面材料为钨， 钨合金片的直径为 010mm， 焦点直径为 01mm， 钨合金 片的厚度范围为 20μιη~2ιηιη， X 射线管电压为 140kV， 电流范围为 2mA~10mA。 可使用 ANSYS 12建立 X射线管阳极有限元模型， 进行热分析计算， 通过更改钨合金片的厚度及电流强度来计算分析阳极上的温度分布。 电子束打在钨表面上， 其焦点直径为 01mm， 电子进入钨的表层平均 深度为 5μιη， 电子是在这段微小的体积内生热。 施加热载荷的方法有两 种： 一种是简化了的施加载荷方法， 将载荷施加在面上， 即在钨的中心 01的表面上施加热载荷， 根据电压和电流可以计算出施加在面上的热流 量大小； 另外一种方法是一局实际情况施加载荷， 将热载荷施加到体上，
Figure imgf000010_0001
Figure imgf000009_0001
From the properties of the material, tungsten has a high melting point but poor thermal conductivity; copper has good thermal conductivity but low melting point. Although graphite has higher melting point and specific heat than tungsten and copper, its atomic number is low and X-ray generation efficiency is low. Therefore, copper can be used as the anode body to take advantage of its good thermal conductivity, and a tungsten alloy sheet is used as a target surface to utilize its high melting point performance. Due to the inconsistent properties of copper and tungsten, the thickness of the tungsten alloy sheet is a key parameter for anode design. If the tungsten alloy sheet is too thick and the heat is too late to pass, the tungsten alloy sheet may be melted first; if the tungsten alloy sheet is too thin and heat is immediately transferred to the copper, the copper may be melted first. In either case, it will affect the normal operation of the X-ray tube. Therefore, the thickness of the tungsten alloy sheet needs to be selected to an optimum value. In order to calculate the optimum thickness value of the tungsten alloy sheet, thermal analysis software can be used to simulate the temperature rise curve of the tungsten alloy sheet with different thicknesses under different electron beam pulse bombardment, tungsten alloy sheet and adjacent metal copper, and heat at the anode. In the transfer process, the relationship between material thickness, electron beam intensity and temperature is studied. Since the heat generation of the electron beam in the pulse state is lower than that in the constant current state under the same intensity, in order to leave a margin for the design, we mainly simulate the parameters in the constant current state. The physical model of the anode is shown in Figure 4: The copper anode body has a geometry of 04Ox5Omm, the target surface material is tungsten, the tungsten alloy sheet has a diameter of 010mm, the focal diameter is 01mm, and the thickness of the tungsten alloy sheet ranges from 20μιη to 2ιηιη, X. The tube voltage is 140kV and the current range is 2mA~10mA. The finite element model of the X-ray tube anode can be established using ANSYS 12 to perform thermal analysis calculations. The temperature distribution on the anode can be calculated by changing the thickness and current intensity of the tungsten alloy sheet. The electron beam is struck on the surface of tungsten with a focal diameter of 01 mm. The average depth of electrons entering the surface of tungsten is 5 μm. The electrons generate heat within this tiny volume. There are two ways to apply a heating load: One is a simplified method of applying a load, applying a load on the surface, that is, applying a thermal load on the surface of the center 01 of tungsten, and calculating the applied surface according to the voltage and current. The amount of heat flow; another method is to apply a load to the actual situation, apply a thermal load to the body,
Figure imgf000010_0001
^ = ^ " ^ = 0.019 , 可以忽略。 为了建模求解方便， 在此使用面载荷的施 加方法， 计算公式如下：  ^ = ^ " ^ = 0.019 , can be ignored. For the convenience of modeling and solving, the application method of surface load is used here, and the calculation formula is as follows:
Q _ KA(Thot - Tcold ) Q _ KA(T hot - T cold )
t d 上式中： Q——时间 ί内的传热量或者热流量。  t d In the above formula: Q – the amount of heat transfer or heat flow in time ί.
Κ一一为热传导率。 One is the thermal conductivity.
Τ——温度。 Τ - temperature.
A—— 触面积。 d——两平面之间的 ii巨离。 在 X射线管工作中， 由于传导散热和辐射散热同时发生， 故可计算它 们对阳极温度上升的影响。 在实际使用过程中， 整个 X射线管都被放入油中绝缘、 冷却。 由于油 的导热系数很小， 因此在 X射线管工作的时候， 热量主要存储在阳极上。 扫描结束后， 经过一段时间才能冷却下来。 故在建模时， 可以先忽略油的 冷却效果。可通过热仿真来计算阳极上的温度分布，进而估算整个阳极的 辐射散热。 阳极温度分布中高温区域很小， 主要集中在电子束焦点， 绝大 部分表面的温度低于 468。C。 根据斯蒂芬-波尔兹曼定理:
Figure imgf000011_0001
A - touch area. d - the ii giant separation between the two planes. In the X-ray tube operation, since conduction heat dissipation and radiation heat dissipation occur simultaneously, their influence on the anode temperature rise can be calculated. In actual use, the entire X-ray tube is placed in the oil to insulate and cool. Since the thermal conductivity of the oil is small, heat is mainly stored on the anode when the X-ray tube is operated. After the scan is over, it will take a while to cool down. Therefore, when modeling, you can ignore the cooling effect of the oil. The temperature distribution on the anode can be calculated by thermal simulation to estimate the radiation dissipation of the entire anode. The high temperature region of the anode temperature distribution is small, mainly concentrated in the electron beam focus, and most of the surface temperature is lower than 468. C. According to Stephen Boltzmann's theorem:
Figure imgf000011_0001
E为辐射力, 单位为 W/m2; E is the radiation force, and the unit is W/m 2 ;
s为物体的辐射率； c为黑体辐射系数， 5.67W/(m2.K4); s is the emissivity of the object; c is the blackbody emissivity, 5.67 W / (m 2 .K 4 );
Γ为物体表面温度。 按照电子束焦点温度 3300摄氏度， 其他表面温度为 400。C进行估算， 则阳极的辐射功率为：
Figure imgf000011_0002
Γ is the surface temperature of the object. According to the electron beam focus temperature of 3300 degrees Celsius, the other surface temperature is 400. C estimates, then the radiant power of the anode is:
Figure imgf000011_0002
=(7T*r*r)* ε钨 *c*(7V100)4+(2*7T*r1*r1+2*7r*r1*/ * *c*(7V100)4 =(7T*r*r)* εTungsten*c*(7V100) 4 +(2*7T*r 1 *r 1 +2*7r*r 1 */ * *c*(7V100) 4
=92.17(W) 阳极的输入功率为 1050W , 那么 ^^ ^O.i^SS, 辐射的功率占输入功 率的比重艮小， 可以忽略掉。 下面是忽略辐射散热和绝缘油传导散热的仿真结果。 根据设计要求， 完成一次 CT扫描的最长时间为 30s， 故在扫描时， X射线管必须可以持 续工作 30s， 此为依据， 计算最优的钨合金片厚度以及可以耐受的最大恒 流电流值。 由图 5可见，在连续入射电子的情况下，当钨合金片厚度为 400~500μιη 的时候，最大耐受电流为 7.5mA。在图中曲线最高点的左边，铜将先熔化， 右边， 钨合金片将先熔化。 对于脉冲工作模式， 不同占空比下， 同一厚度的钨合金片所能够耐受 的最大脉冲电流随着占空比的减少而增加。 考虑阳极靶的使用寿命， 以及电子束的脉冲工作模式， 本发明实施例 将选用钨合金靶面的厚度为 400-500um，例如优选 0.5mm为钨合金片的优 选厚度值。 可选的， 如图 7所示， X射线管的阳极 1包括阳极体 11和靶面 12。 靶 面 12相对参考方向形成有预定的靶面倾角 α,参考方向与电子入射方向垂 直， 如图 7所示。 靶面倾角 α是一个关键参数， 它将直接影响到 X射线管的光产额、 有效 焦点尺寸、 热量分布与传递等。 为了研究靶面倾角的变化对 X光子的产额 和角度分布的影响， 可采用蒙特卡罗方法对其进行了模拟计算。 例如使用 EGS软件模拟了 lxlO7个 140keV的电子轰击不同倾角的钨靶， 统计了光产 额和光子的空间分布。 靶面倾角与光子产额的关系见图 6。 从图 6中可以 看出， 靶面倾角越小， X光子产额越高。 不过， 靶面倾角是不是越小越好， 这需要进行仔细的分析。在 CT扫描 过程中最终利用的是以电子束入射方向近似垂直的扇形束之内 X光子，这 部分 X光子才是真正为 CT成係教出贡献的 （如图 7所示）， 因此这个角 度范围内的 X光子越多越好。 下图为靶面倾角 5度时， 与靶面不同夹角的光子面密度。 从图 8中可 以看出， 随着与靶面夹角的增加， 光子的面密度越来越小， 即可用于成像 的 X光子数越来越少。 因此， 虽然靶面倾角 5度时的总光子产额很高， 但 是与靶面夹角 85度处的光子面密度却很低。 对不同靶面倾角下与电子束入射方向垂直的出射面内 X光子的数量进 行统计， 统计结果见图 9。 从图 9中可以看出， 随着靶面倾角的增加， 出 射面的光子数随之增加， 但是在 45度左右达到最大值， 然后便开始减小。 在 CT成像中， 影响断层图像分辨率的是 X射线管的有效焦点， 而不 是实际焦点。 假设电子束平行入射， 则实际焦点尺寸 L与投影后的有效焦 点尺寸 d之间的关系如下： =92.17(W) The input power of the anode is 1050W, then ^^ ^Oi^SS, the radiated power accounts for a small proportion of the input power, which can be ignored. The following is a simulation result that ignores the radiation heat dissipation and the conduction heat dissipation of the insulating oil. According to the design requirements, the maximum time for completing a CT scan is 30s, so the X-ray tube must be able to work continuously for 30s during scanning. Based on this, the optimal tungsten alloy sheet thickness and the maximum constant current that can be tolerated are calculated. value. As can be seen from Fig. 5, in the case of continuous incident electrons, the maximum withstand current is 7.5 mA when the thickness of the tungsten alloy sheet is 400 to 500 μm. On the left side of the highest point of the curve in the figure, the copper will melt first, and on the right, the tungsten alloy sheet will melt first. For the pulse mode of operation, the maximum pulse current that a tungsten alloy sheet of the same thickness can withstand at different duty cycles increases as the duty cycle decreases. Considering the service life of the anode target and the pulsed operation mode of the electron beam, the embodiment of the present invention will select a tungsten alloy target surface having a thickness of 400-500 um, for example, preferably 0.5 mm, as a preferred thickness value of the tungsten alloy sheet. Alternatively, as shown in FIG. 7, the anode 1 of the X-ray tube includes an anode body 11 and a target surface 12. The target surface 12 is formed with a predetermined target tilt angle α with respect to the reference direction, and the reference direction is perpendicular to the electron incident direction, as shown in FIG. The target tilt angle α is a key parameter that directly affects the light yield, effective focus size, heat distribution and transfer of the X-ray tube. In order to study the influence of the change of the target tilt angle on the yield and angular distribution of X-ray photons, Monte Carlo method can be used to simulate the calculation. For example, EGS software was used to simulate lxlO 7 140keV electrons bombarding tungsten targets with different dip angles, and the spatial distribution of light yield and photons was counted. The relationship between target dip angle and photon yield is shown in Figure 6. As can be seen from Figure 6, the smaller the target tilt angle, the higher the X-ray photo yield. However, the smaller the target angle, the better, which requires careful analysis. In the CT scanning process, the X-ray photons in the fan beam which are approximately perpendicular to the incident direction of the electron beam are finally used. This part of the X-ray photo is actually contributing to the CT system (as shown in Fig. 7), so this angle The more X-rays in the range, the better. The figure below shows the photon areal density at an angle different from the target surface at a target angle of 5 degrees. As can be seen from Fig. 8, as the angle with the target surface increases, the areal density of photons becomes smaller and smaller, and the number of X-photons that can be used for imaging becomes less and less. Therefore, although the total photon yield at a target angle of 5 degrees is high, the photon surface density at an angle of 85 degrees from the target surface is low. The number of X-photons in the exit plane perpendicular to the incident direction of the electron beam at different target inclination angles is counted. The statistical results are shown in Fig. 9. As can be seen from Fig. 9, as the target tilt angle increases, the number of photons on the exit surface increases, but reaches a maximum at about 45 degrees and then begins to decrease. In CT imaging, the resolution of the tomographic image is the effective focus of the X-ray tube, not the actual focus. Assuming that the electron beams are incident in parallel, the relationship between the actual focus size L and the projected effective focus size d is as follows:
从上式可以看出， 如果实际焦点的尺寸 L很难减小时， 可以通过减小 靶面倾角 α来控制有效焦点的尺寸 d。 如果入射的电子束单位横截面积的密度无法提高， 根据下式可知， 增 大电子束流宽度 h减小靶面倾角 α有可能提高可成像 X光子的总数。 As can be seen from the above equation, if the size L of the actual focus is difficult to reduce, the size d of the effective focus can be controlled by reducing the target tilt angle α. If the density of the cross-sectional area of the incident electron beam cannot be increased, it can be seen from the following equation that increasing the electron beam width h by decreasing the target tilt angle α may increase the total number of imageable X-photons.
if-― ¾ tgS£ 保持有效焦点尺寸和电子束单位横截面积的密度不变， 靶面倾角与可 用于成像的 X光子数之间的关系曲线见图 10。 从图 10中可以看出， 靶面倾角越小， 通过增加电子束流宽度可以有效 增加可用于成像的 X光子数量。 不过结合前图可知， 此时， 入射的电子束 流的总量显著增加， 进而增加了阳极所接受的热量， 这将给 X射线管的散 热提出了挑战。 因此， 阳极的靶面倾角的确定需要在可用于成像的 X光子 数量与入射电子的热量之间寻求一种平衡。 经过综合考虑， 靶面倾角优选 为 11度， 如图 1A所示。 If-― 3⁄4 tgS£ Keep the effective focus size and the density of the electron beam unit cross-sectional area constant. The relationship between the target tilt angle and the number of X-photons available for imaging is shown in Fig. 10. As can be seen from Fig. 10, the smaller the target tilt angle, the more the number of X-photons available for imaging can be effectively increased by increasing the beam width. However, as can be seen from the previous figure, at this time, the total amount of incident electron beam current is significantly increased, which in turn increases the heat received by the anode, which poses a challenge to the heat dissipation of the X-ray tube. Therefore, the determination of the target face tilt angle of the anode requires a balance between the amount of X-photons available for imaging and the heat of incident electrons. After comprehensive consideration, the target tilt angle is preferably 11 degrees, as shown in Fig. 1A.
可选的， 上述实施例中的 X射线管总长度小于或等于 120mm， 和 / 或， X射线源的总重量小于 25kg, 以充分保证 X射线管的小巧型， 可以 便于携带， 方便适用于舰载、 车载、 战地医院等特殊环境。  Optionally, the total length of the X-ray tube in the above embodiment is less than or equal to 120 mm, and/or the total weight of the X-ray source is less than 25 kg, so as to fully ensure the compact shape of the X-ray tube, which can be easily carried, and is convenient for the ship. Special environments such as trucks, vehicles, and battlefield hospitals.
可选的，上述实施例中的最大直径小于或等于 60mm。进一步优选地， 上述实施例中的阳极和阴极中尖锥顶部的距离小于或等于 10um。 这样可 以保证 X射线管的优良性能。  Alternatively, the maximum diameter in the above embodiment is less than or equal to 60 mm. Further preferably, the distance between the anode and the tip of the tip of the cathode in the above embodiment is less than or equal to 10 um. This ensures excellent performance of the X-ray tube.
此外， 本发明还提供了一种移动 CT扫描仪， 该移动 CT扫描仪包括 上述任一实施例提供的医学检测用 X射线源， 通过该 X射线源产生 X射 线以对脑部等人体部位进行医学检测。  In addition, the present invention also provides a mobile CT scanner comprising the X-ray source for medical detection provided by any of the above embodiments, wherein X-rays are generated by the X-ray source to perform body parts such as the brain. Medical testing.
在本发明上述各实施例中，实施例的序号或先后顺序仅仅为了便于描 述， 不代表实施例的优劣。对各个实施例的描述都各有侧重， 某个实施例 中没有详述的部分， 可以参见其他实施例的相关描述。  In the above embodiments of the present invention, the serial numbers or the sequence of the embodiments are merely for convenience of description, and do not represent the advantages and disadvantages of the embodiments. The descriptions of the various embodiments are all focused on, and the parts that are not detailed in an embodiment can be referred to the related description of other embodiments.
本领域普通技术人员可以理解：实现上述方法实施例的全部或部分步 骤可以通过程序指令相关的硬件来完成，前述的程序可以存储于一计算机 可读取存储介质中， 该程序在执行时， 执行包括上述方法实施例的步骤； 而前述的存储介质包括： 只读存储器（Read-Only Memory, 简称 ROM )、 随才 储器（Random Access Memory, 简称 RAM )、 磁碟或者光盘 等各种可以存储程序代码的介质。  A person skilled in the art can understand that all or part of the steps of implementing the above method embodiments may be completed by using hardware related to the program instructions. The foregoing program may be stored in a computer readable storage medium, and the program is executed when executed. The foregoing storage device includes the following steps: The foregoing storage medium includes: a Read-Only Memory (ROM), a Random Access Memory (RAM), a magnetic disk, or an optical disk, and the like. The medium of the program code.
在本发明的装置和方法等实施例中，显然，各部件或各步骤是可以分 解、 组合和 /或分解后重新组合的。 这些分解和 /或重新组合应视为本发明 的等效方案。 同时， 在上面对本发明具体实施例的描述中， 针对一种实施 方式描述和 /或示出的特征可以以相同或类似的方式在一个或更多个其它 实施方式中使用， 与其它实施方式中的特征相组合，或替代其它实施方式 中的特征。  In the embodiments of the apparatus and method of the present invention, it is apparent that the various components or steps may be recombined after being decomposed, combined, and/or disassembled. These decompositions and/or recombinations should be considered as equivalents to the invention. Also, in the above description of the specific embodiments of the present invention, features described and/or illustrated with respect to one embodiment may be used in the same or similar manner in one or more other embodiments, and in other embodiments. The features are combined or substituted for features in other embodiments.
应该强调， 术语 "包括 /包含" 在本文使用时指特征、 要素、 步骤或 组件的存在， 但并不排除一个或更多个其它特征、要素、 步骤或组件的存 在或附加。 最后应说明的是： 虽然以上已经详细说明了本发明及其优点，但 当理解在不超出由所附的权利要求所限定的本发明的精神和范围的情况 下可以进行各种改变、替代和变换。 而且， 本发明的范围不仅限于说明书 所描述的过程、 设备、 手段、 方法和步骤的具体实施例。 本领域内的普通 技术人员从本发明的公开内容将容易理解，根据本发明可以使用执行与在 此所述的相应实施例基本相同的功能或者获得与其基本相同的结果的、现 有和将来要被开发的过程、 设备、 手段、 方法或者步骤。 因此， 所附的权 利要求旨在在它们的范围内包括这样的过程、设备、手段、方法或者步骤。 It should be emphasized that the term "comprising" or "comprising" is used to mean the presence of features, elements, steps or components, but does not exclude the presence or addition of one or more other features, elements, steps or components. It should be noted that, although the invention and its advantages are described in detail, it is understood that various changes, substitutions and changes can be made without departing from the spirit and scope of the invention as defined by the appended claims. Transform. Further, the scope of the invention is not limited to the specific embodiments of the processes, devices, means, methods and steps described in the specification. It will be readily apparent to those skilled in the art from this disclosure that the present invention can be used in accordance with the present invention to perform substantially the same functions as the corresponding embodiments described herein or to obtain substantially the same results as the present and future The process, equipment, means, method or step being developed. Therefore, the appended claims are intended to cover such a process, apparatus, means, methods, or steps.

Claims

权利 要求 书 claims
1、 一种医学检测用 X射线源， 其特征在于， 包括： 1. An X-ray source for medical detection, which is characterized by including:
X射线管， 包括阳极、 冷阴极、 栅极和管壳； 所述管壳用于支撑所述 阳极、 冷阴极和栅极， 并使得所述阳极、冷阴极和栅极的真空工作环境与 外界绝缘， 所述阳极接地； The X-ray tube includes an anode, a cold cathode, a grid, and a tube shell; the tube shell is used to support the anode, cold cathode, and grid, and to separate the vacuum working environment of the anode, cold cathode, and grid from the outside world. Insulated, the anode is grounded;
高压发生器，用于在所述冷阴极和所述栅极之间提供第一电场使所述 冷阴极场发射电子，以及在所述栅极和所述阳极之间提供第二电场以加速 所述冷阴^ L射的电子， 使之轰击所述阳极来产生 X射线。 A high-voltage generator for providing a first electric field between the cold cathode and the grid to cause the cold cathode to field emit electrons, and a second electric field between the grid and the anode to accelerate the electrons. The electrons emitted by the cold cathode bombard the anode to generate X-rays.
2、 根据权利要求 1所述的医学检测用 X射线源， 其特征在于， 所述第一电场的相关参数包括： 500v-1000v 的直流电压， 高于 50w 的功率， 工作频率 300Hz-3000Hz， 脉冲占空比为 20% -80%； 和 /或， 所述第二电场的相关参数包括： 高于 140kv的直流电压， 2mA-16mA 的管电流， 高于 2000W的功率。 2. The X-ray source for medical detection according to claim 1, characterized in that the relevant parameters of the first electric field include: DC voltage of 500v-1000v, power higher than 50w, operating frequency 300Hz-3000Hz, pulse The duty cycle is 20% -80%; and/or, the relevant parameters of the second electric field include: a DC voltage higher than 140kv, a tube current of 2mA-16mA, and a power higher than 2000W.
3、 根据权利要求 1或 2所述的医学检测用 X射线源， 其特征在于， 所述冷阴极包括： 以及形成于所述^ 上的碳纳米管发射阵列； 或者， 3. The X-ray source for medical detection according to claim 1 or 2, wherein the cold cathode includes: and a carbon nanotube emission array formed on the cold cathode; or,
所述冷阴极包括： 基板以及形成于所述基板上的 LaB6尖锥场发射阵 列。 The cold cathode includes: a substrate and a LaB6 pointed cone field emission array formed on the substrate.
4、 根据权利要求 3所述的医学检测用 X射线源， 其特征在于， 所述 LaB6尖锥场发射阵列包括： 二极管 LaB6尖锥场发射阵列， 所 述二极管 LaB6尖锥场发射阵列包括：硅尖锥二极管阵列和覆盖在硅尖锥 表面上的 LaB6纳米材料薄膜层； 4. The X-ray source for medical detection according to claim 3, wherein the LaB6 cone field emission array includes: a diode LaB6 cone field emission array, and the diode LaB6 cone field emission array includes: silicon The pyramid diode array and the LaB6 nanomaterial film layer covering the surface of the silicon pyramid;
或者， or,
所述 LaB6尖锥场发射阵列包括： 三极管 LaB6尖锥场发射阵列， 所 述三极管 LaB6尖锥场发射阵列包括： 硅基、 形成在所述硅基上的孔腔阵 列、 分布在各孔腔中的钼尖锥阵列、 以及覆盖在各钼尖锥表面上的 LaB6 纳米材料薄膜层。 The LaB6 pointed cone field emission array includes: a triode LaB6 pointed cone field emission array; the triode LaB6 pointed cone field emission array includes: a silicon base, an array of holes formed on the silicon base, and distributed in each hole. An array of molybdenum cones, and a LaB6 nanomaterial film layer covering the surface of each molybdenum cone.
5、 根据权利要求 3所述的医学检测用 X射线源， 其特征在于， 所述 阳极为固定式阳极或者旋转式阳极。 5. The X-ray source for medical detection according to claim 3, characterized in that the anode is a fixed anode or a rotating anode.
6、 根据权利要求 5所述的医学检测用 X射线源， 其特征在于， 所述 固定式阳极包括：固定的铜阳极体以及固定于所述铜阳极体上的钨合金靶 面。 6. The X-ray source for medical detection according to claim 5, characterized in that the fixed anode includes: a fixed copper anode body and a tungsten alloy target surface fixed on the copper anode body.
7、 根据权利要求 6所述的医学检测用 X射线源， 其特征在于， 所述 钨合金靶面相对参考方向形成有预定的靶面倾角，所述参考方向与电子入 射方向垂直。 7. The X-ray source for medical detection according to claim 6, characterized in that the tungsten alloy target surface forms a predetermined target surface inclination angle relative to a reference direction, and the reference direction is perpendicular to the electron incident direction.
8、 根据权利要求 7所述的医学检测用 X射线源， 其特征在于， 所述 钨合金靶面的厚度为 400-500um， 和 /或， 所述靶面倾角为 11度。 8. The X-ray source for medical detection according to claim 7, characterized in that the thickness of the tungsten alloy target surface is 400-500um, and/or the inclination angle of the target surface is 11 degrees.
9、 根据权利要求 3所述的医学检测用 X射线源， 其特征在于， 所述 X射线管总长度小于或等于 120mm， 和 /或， 所述 X射线管的最大直径小 于或等于 60mm， 和 /或， 所述阳极和所述冷阴极顶部的距离小于或等于 lOum, 和 /或， 所述 X射线源的总重量小于 25kg。 9. The X-ray source for medical detection according to claim 3, characterized in that, the total length of the X-ray tube is less than or equal to 120 mm, and/or, the maximum diameter of the X-ray tube is less than or equal to 60 mm, and /or, the distance between the anode and the top of the cold cathode is less than or equal to 10um, and/or, the total weight of the X-ray source is less than 25kg.
10、 一种移动 CT扫描仪， 其特征在于， 包括如权利要求 1-9任一所 述的医学检测用 X射线源。 10. A mobile CT scanner, characterized by including the X-ray source for medical detection as described in any one of claims 1-9.
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