CN212214394U - Miniaturized flash radiotherapy device - Google Patents
Miniaturized flash radiotherapy device Download PDFInfo
- Publication number
- CN212214394U CN212214394U CN202021062058.6U CN202021062058U CN212214394U CN 212214394 U CN212214394 U CN 212214394U CN 202021062058 U CN202021062058 U CN 202021062058U CN 212214394 U CN212214394 U CN 212214394U
- Authority
- CN
- China
- Prior art keywords
- electron beam
- linear accelerator
- energy
- electron
- radio
- Prior art date
- Legal status (The legal status is an assumption and is not a legal conclusion. Google has not performed a legal analysis and makes no representation as to the accuracy of the status listed.)
- Active
Links
Images
Landscapes
- Radiation-Therapy Devices (AREA)
Abstract
The utility model discloses a miniaturized flash radiation therapy device, including grid-control electron gun, normal atmospheric temperature radio frequency linear accelerator, time domain synthesis radio frequency power source, X ray target and collimator, grid-control electron gun conducts the electron beam to normal atmospheric temperature radio frequency linear accelerator through first transmission line, and time domain synthesis radio frequency power source provides the microwave for normal atmospheric temperature radio frequency linear accelerator, and normal atmospheric temperature radio frequency linear accelerator passes through the second transmission line and conducts the electron beam to the X ray target, and the electron beam bombards and generates X ray on the X ray target, and X ray shines on the target via the collimator. According to the utility model discloses a miniaturized flash of light radiotherapy device can provide long macropulse high dose rate X ray, gives the very high dose of shining of target area in the short time, satisfies flash of light radiotherapy's requirement, reaches better radiotherapy effect.
Description
Technical Field
The utility model belongs to the field of radiation devices, in particular to a medical miniaturized flash radiotherapy device.
Background
At present, the incidence of cancer in China is higher and higher, and the cancer becomes one of the biggest killers harmful to the health of people in China, and common treatment means include operations, chemotherapy, radiotherapy and the like.
The principle of radiotherapy is that radiation with certain energy acts on cells to generate ionization effect, and releases large amount of energy locally to destroy the DNA chain of cancer cells, so as to inhibit and kill tumor cells. Under the condition of conventional dose rate irradiation, the irradiation range of normal tissues is reduced by improving the conformity degree of the radiation field of a radiotherapy machine to finally irradiate the tumor tissues as accurately as possible. However, due to the structural particularity of human organs, normal tissues in the radiotherapy process are necessarily irradiated by a certain dose of radiation, especially the normal tissues adjacent to tumor tissues. Serious toxic side effects of normal tissue after exposure to radiation are also one of the major causes of treatment-related death in patients. Thus, dose-limiting toxicity of normal tissues remains one of the major resistance to tumor radiotherapy development.
The radiation dose rate of the current medical accelerator is about 0.1Gy/s, the total time of the tumor patients completing all the radiation is about 7.5 hours, and the radiation is distributed in about 1.5 months. In addition to the long overall treatment time and high labor cost for the patient, the radiotoxicity of normal tissues gradually develops as the total dose of radiotherapy accumulates. The most common radioactive lung injury in toxic reactions, the radioactive cardiotoxicity, is even more of a cause of substantial death in patients who have had more or less cured tumors. On the other hand, the doctor in the tumor radiotherapy department is forced to sacrifice the treatment dosage of the target area of the tumor due to the fear of serious toxic and side effects of normal tissues, so that the tumor of a patient cannot be controlled for a long time, and the case of final relapse is very painful. How to control the toxic and side effects related to treatment while treating tumors becomes a bottleneck which is urgently needed to be broken through by the current tumor radiotherapy.
According to the literature report, by increasing the X-ray radiation dose rate to 106Gy/s to 108At Gy/s, the radiosensitivity of normal tissue can be reduced (radiation resistance occurs and toxic side effects are reduced), but tumor tissue is still sensitive to radiation, and this phenomenon is called "flash effect", and the condition for the "flash effect" is that high dose rate radiation is emitted in a very short time (usually in the order of nanoseconds to hundreds of milliseconds), and this is called "flash radiotherapy". "flash radiation therapy" is a radiation method that currently has the potential to break through the dose-limiting toxicity of normal tissues.
The existing radiotherapy device can not provide long-pulse high-dosage-rate X-rays.
SUMMERY OF THE UTILITY MODEL
In order to solve the above-mentioned problems, a compact flash radiotherapy device has been proposed. The utility model provides a following technical scheme:
a miniature flash radiotherapy device comprises a grid-control electron gun for generating a long pulse electron beam, a normal-temperature radio-frequency linear accelerator for energizing the electron beam, a time-domain synthesis radio-frequency power source for providing long pulse microwave power for the normal-temperature radio-frequency linear accelerator, an X-ray target for generating X-rays through electron beam bombardment and a collimator for adjusting an X-ray irradiation area, wherein an electron source conducts the electron beam to the normal-temperature radio-frequency linear accelerator through a first transmission line, the normal-temperature radio-frequency linear accelerator conducts the electron beam to the X-ray target through a second transmission line, the electron beam bombards the X-ray target to generate X-rays, and the X-rays irradiate a target needing radiotherapy through the collimator.
Furthermore, the pulse time length of the grid-control electron gun for generating the electron beam is adjustable, and the adjustment range is 10 microseconds to 10 milliseconds.
Furthermore, the grid-controlled electron gun is a microwave grid-controlled hot cathode electron gun or a high-voltage pulse grid-controlled hot cathode electron gun.
Furthermore, the microwave grid-control hot cathode electron gun generates an electron beam with high average current intensity of 1-100mA and long macropulse of 10 microseconds to 10 milliseconds, and the electron beam is injected into the normal-temperature radio-frequency linear accelerator through the first transmission line to be energized, so that the electron energy is increased to 9-15 MeV.
Furthermore, the electron beams comprise a low-energy electron beam and a high-energy electron beam, the low-energy electron beam is directly generated by the grid-control electron gun, and the high-energy electron beam is generated by the low-energy electron beam after being energized by the normal-temperature radio-frequency linear accelerator.
Furthermore, the electron beam emitted by the normal temperature radio frequency linear accelerator bombards the X-ray target through a second transmission line to generate a long macropulse 106Gy/s-108X-rays at Gy/s dose rate, single macroThe pulse dose is 10-80 Gy.
Furthermore, the time domain synthesis radio frequency power source is composed of a plurality of radio frequency power sources according to a time domain synthesis circuit.
Furthermore, the normal temperature radio frequency linear accelerator comprises a plurality of radio frequency resonant cavities used for increasing the energy of the electron beams, the radio frequency resonant cavities are distributed along the axis of the normal temperature radio frequency linear accelerator, and the time domain synthesis radio frequency power source is electrically connected with the radio frequency resonant cavities.
Further, the X-ray target is a high atomic number material ray target, and the electron beam interacts with the high atomic number material to generate X-rays.
Furthermore, a heat dissipation assembly is further arranged on the X-ray target, and the heat dissipation assembly is an automatic rotating disk or a cooling water pipe.
Has the advantages that:
according to the utility model discloses a miniaturized flash of light radiotherapy device can provide the X ray of the high dose rate of long macropulse, can give the very high dose that shines of target area in the short time to can adjust the energy of ray through the energy of adjusting the electron beam, adjust the time length of ray through the length of adjusting the electron beam pulse, adjust the dose rate through the current intensity of adjusting the electron beam, in order to reach the better radiotherapy effect of target. The device has simple structure and small scale, is suitable for the treatment method of 'flash radiotherapy' popularized in hospitals, reduces the purchase cost of equipment and further reduces the treatment cost of patients.
Drawings
FIG. 1 is a schematic view of the main structure of a miniaturized flash radiotherapy device according to an embodiment of the present invention;
FIG. 2 is a schematic diagram of the main structure of a grid-controlled electron gun according to an embodiment of the present invention;
FIG. 3 is a schematic diagram of an electron beam time structure according to an embodiment of the present invention;
fig. 4 is a schematic diagram of a microwave pulse structure for synthesizing a long pulse by a time domain synthesis rf power source according to an embodiment of the present invention;
fig. 5 is a schematic diagram of a structure of a radio frequency resonant cavity in a normal temperature radio frequency linear accelerator according to an embodiment of the present invention.
In the drawings: 1. a gated electron gun; 2. a first transmission line; 3. a normal temperature radio frequency linear accelerator; 4. time domain synthesis radio frequency power source; 5. a second transmission line; 6. an X-ray target; 7. a collimator; 8. a target; 9. a radio frequency resonant cavity; 10. a cathode; 11. a gate electrode; 12. an anode; 13. a radio frequency power source; 14. a time domain synthesis circuit; A. micro-pulsing; t, the macro-pulse time length.
Detailed Description
In order to make those skilled in the art better understand the technical solution of the present invention, the following description, together with the drawings of the present invention, clearly and completely describes the technical solution of the present invention, and based on the embodiments in the present application, other similar embodiments obtained by those skilled in the art without creative efforts shall all belong to the protection scope of the present application. In addition, directional terms such as "upper", "lower", "left", "right", etc. in the following embodiments are directions with reference to the drawings only, and thus, the directional terms are used for illustration and not for limitation of the present invention.
Example 1
As shown in fig. 1, a grid-controlled electron gun 1 is a miniaturized flash radiotherapy device using a microwave grid-controlled hot cathode electron gun, and includes a microwave grid-controlled hot cathode electron gun for generating an electron beam, a normal temperature radio frequency linear accelerator 3 for energizing the electron beam, a time domain synthesis radio frequency power source 4 for providing a long pulse microwave power to the normal temperature radio frequency linear accelerator 3, an X-ray target 6 for generating an X-ray by electron beam bombardment, and a collimator 7 for adjusting an X-ray irradiation area. The microwave grid-control hot cathode electron gun transmits electron beams to a normal-temperature radio-frequency linear accelerator 3 through a first transmission line 2, the normal-temperature radio-frequency linear accelerator 3 transmits the electron beams to an X-ray target 6 through a second transmission line 5, the electron beams bombard the X-ray target 6 to generate X-rays, and the X-rays irradiate a target 8 needing radiotherapy through a collimator 7. The microwave grid-control hot cathode electron gun and the normal-temperature radio-frequency linear accelerator 3 operate together to generate relativistic electron beams, the microwave grid-control hot cathode electron gun generates a low-energy electron beam EB1 with first energy E1, the low-energy electron beam EB1 enters the normal-temperature radio-frequency linear accelerator 3 through the first transmission line 2 and penetrates through the normal-temperature radio-frequency linear accelerator 3 to obtain energy increase delta E to become high-energy electron beam EB2, the high-energy electron beam EB2 bombards the X-ray target 6 along the second transmission line 5 to generate X rays, the X rays irradiate the target 8 through the collimator 7, and the collimator 7 adjusts the distribution of the X rays by filtering the X rays outside the target 8 area. In this embodiment, the target 8 is a human body requiring radiotherapy.
The microwave grid-controlled hot cathode electron gun comprises a cathode 10, a grid 11 and an anode 12, as shown in fig. 2. Electrons emitted from the cathode 10 are focused by a grid 11 having a certain shape, and are extracted by a high voltage between the anode 10 and the cathode 12. A microwave resonant cavity is formed between the cathode 10 and the grid 11, electrons are pulled out from the cathode 10 by utilizing the axial field of the radio frequency electromagnetic field in the resonant cavity, and the grid is led out in an energizing manner. The cathode electrons can be extracted only at a proper phase, namely, the electrons are selectively extracted, and an electron beam group with a short micropulse is obtained.
The temporal structure of the low-energy electron beam EB1 is shown in fig. 3, where the electron beam macropulse is composed of a plurality of sequential micropulses a. The high voltage pulse length between the anode and cathode determines the macro pulse length of the low energy electron beam EB 1. The microwave grid-controlled hot cathode electron gun can generate long-pulse low-energy electron beams, the time length T of the macro pulse can be adjusted, and in the embodiment, the pulse length is 10 microseconds to 10 milliseconds. In some embodiments, the pulse length may be made longer.
The time domain synthesis rf power source 4 comprises a plurality of rf power sources 13 and a time domain synthesis circuit 14, as shown in fig. 4. A single radio frequency power source generates short-pulse microwave pulses, and a time domain synthesis circuit synthesizes a plurality of short microwave pulses into a long microwave pulse. The microwave pulse generated by the single radio frequency power source 13 can independently drive the normal temperature radio frequency linear accelerator, and the time length of the microwave pulse is short and is in the order of several microseconds to sub-milliseconds. The microwave pulses generated by the plurality of rf power sources 13 are combined in the time domain by the time domain combining circuit 14, and a long pulse microwave pulse can be generated, as shown in fig. 4. The long-pulse microwave pulse drives the normal-temperature radio-frequency linear accelerator 3, and can energize long-pulse electron beams.
The electron beams comprise a low-energy electron beam and a high-energy electron beam, the low-energy electron beam is an electron beam EB1 directly generated by a microwave grid-control hot cathode electron gun, and the high-energy electron beam is an electron beam EB2 generated by the low-energy electron beam EB1 after being energized by the normal-temperature radio-frequency linear accelerator 3.
The normal temperature rf linear accelerator 3 is composed of a plurality of rf resonant cavities 9 for increasing the energy of the electron beam, as shown in fig. 5. The radio frequency resonant cavity 9 is distributed along the axis of the normal temperature radio frequency linear accelerator 3, the time domain synthesis radio frequency power source 4 is electrically connected with the radio frequency resonant cavity 13, the resonant frequency of the radio frequency resonant cavity 9 is the same as the frequency of the time domain synthesis radio frequency power source 4, the resonant frequency of the radio frequency resonant cavity 9 can be an L wave band, an S wave band, a C wave band or an X wave band, and the frequency of the corresponding time domain synthesis radio frequency power source 4 can be an L wave band, an S wave band, a C wave band or an X wave band. The low-energy electron beam EB1 penetrates through the normal-temperature radio-frequency linear accelerator 3 to obtain energy increase delta E, and the energy increase delta E becomes a high-energy electron beam EB2, the delta E is determined by the scale and the performance of the normal-temperature radio-frequency linear accelerator 3, the more the number of the radio-frequency resonant cavities 9 is, the larger the field gradient is, and the larger the energy increase delta E is. The energy gain Δ E in this example is 9-15 MeV. In this embodiment, the type and number of the rf resonant cavities 9 are determined, and the energy increase Δ E can be adjusted by adjusting the intensity of the rf field to meet the X-ray energy requirements for different radiotherapy treatments. The strength of the radio frequency field is positively correlated with the feed-in power of the power source, the power of the power source is adjusted, and the field gradient of the radio frequency field in the room-temperature radio frequency cavity is correspondingly changed.
The X-ray target 6 is a high atomic number material ray target, the electron beam and the high atomic number material interact to generate X-rays, and the X-ray target 6 is also provided with a heat dissipation assembly which is an automatic rotating disk or a cooling water pipe. The core component of the X-ray target 6 is a high atomic number material, i.e., a high Z material, which generates X-rays through the interaction of the electron beam and the high Z material. The X-ray target 6 generally uses tungsten or tantalum which is hard in texture, fast in heat transfer, and high in melting point. In this embodiment, the X-ray target 6 is a tungsten target. Part of the energy of the electron beam is converted into the energy of the X-ray, and the other part is deposited on the X-ray target 6 in the form of thermal energy, so that the X-ray target 6 has a cooling function to prevent the energy of the electron beam deposition from burning out the X-ray target 6. In this embodiment, a rotating target is used, i.e. the target body rotates continuously, so that the electron beam hits different parts and the heat is not too concentrated.
The long macro-pulse electron beam bombards the X-ray target to generate X-rays and generate long macro-pulse high-dosage-rate X-rays. The dosage rate can be up to 106Gy/s to 108Gy/s, the dose of a single macropulse can reach 10-80 Gy.
A method for using a miniaturized flash radiotherapy device mainly comprises the following steps:
an electron beam with high average flow strength of 1-100mA and long macropulse of 10 microseconds to 10 milliseconds is generated by a microwave grid-controlled hot cathode electron gun, and is injected into a normal-temperature radio-frequency linear accelerator 3 through a first transmission line 2 for energizing, so that the electron energy is increased to 9-15 MeV. Electrons emitted by the normal temperature radio frequency linear accelerator 3 bombard the X-ray target 6 through the second transmission line 5 to generate long-macropulse X-rays with high dosage rate which can reach 106Gy/s to 108Gy/s, the dose of a single macropulse can reach 10-80 Gy. The energy of the rays is adjusted by adjusting the energy of the electron beams, the time length of the rays is adjusted by adjusting the pulse length of the electron beams, the X-ray dose rate is adjusted by adjusting the current intensity of the electron beams, and the adjusted X-rays are incident to a target through a collimator to perform related treatment or experiment.
Example 2
As shown in fig. 1, a miniaturized flash radiotherapy apparatus using a high-voltage pulse grid-controlled hot cathode electron gun as a grid-controlled electron gun 1 includes a high-voltage pulse grid-controlled hot cathode electron gun for generating an electron beam, a normal temperature radio frequency linear accelerator 3 for energizing the electron beam, a time domain synthesis radio frequency power source 4 for providing a long pulse microwave power to the normal temperature radio frequency linear accelerator 3, an X-ray target 6 for generating an X-ray by electron beam bombardment, and a collimator 7 for adjusting an X-ray irradiation area. The high-voltage pulse grid-control hot cathode electron gun transmits electron beams to a normal-temperature radio-frequency linear accelerator 3 through a first transmission line 2, the normal-temperature radio-frequency linear accelerator 3 transmits the electron beams to an X-ray target 6 through a second transmission line 5, the electron beams bombard the X-ray target 6 to generate X-rays, and the X-rays irradiate a target 8 needing radiotherapy through a collimator 7. The high-voltage pulse grid-control hot cathode electron gun and the normal-temperature radio-frequency linear accelerator 3 operate together to generate relativistic electron beams, the high-voltage pulse grid-control hot cathode electron gun generates a low-energy electron beam EB1 with first energy E1, the low-energy electron beam EB1 is incident on the normal-temperature radio-frequency linear accelerator 3 through the first transmission line 2 and penetrates through the normal-temperature radio-frequency linear accelerator 3 to obtain energy increase delta E to become high-energy electron beam EB2, the high-energy electron beam EB2 bombards the X-ray target 6 along the second transmission line 5 to generate X-rays, the X-rays irradiate the target 8 through the collimator 7, and the collimator 7 filters the X-rays outside the target 8 area to adjust the distribution of the X-rays. In this embodiment, the target 8 is a human body requiring radiotherapy.
The high voltage pulse gated hot cathode electron gun comprises a cathode 10, a grid 11 and an anode 12 as shown in fig. 2. Electrons emitted by the cathode 10 are focused by the grid 11 with a certain shape and are extracted by the direct current high voltage between the anode 10 and the cathode 12. A pulse high-voltage field is formed between the cathode 10 and the grid 11, and electrons are pulled out from the cathode 10 by using the axial field of the electric field in the pulse high-voltage field and are energized to be led out of the grid.
The temporal structure of the low-energy electron beam EB1 is shown in fig. 3, where the electron beam macropulse is composed of a plurality of sequential micropulses a. The length of the dc high voltage pulse between the anode and cathode determines the macropulse length of the low energy electron beam EB 1. The high-voltage pulse grid-controlled hot cathode electron gun can generate long-pulse low-energy electron beams, the time length T of the macro pulse can be adjusted, and in the embodiment, the pulse length is 10 microseconds to 10 milliseconds. In some embodiments, the pulse length may be made longer.
The time domain synthesis rf power source 4 comprises a plurality of rf power sources 13 and a time domain synthesis circuit 14, as shown in fig. 4. A single radio frequency power source generates short-pulse microwave pulses, and a time domain synthesis circuit synthesizes a plurality of short microwave pulses into a long microwave pulse. The microwave pulse generated by the single radio frequency power source 13 can independently drive the normal temperature radio frequency linear accelerator, and the time length of the microwave pulse is short and is in the order of several microseconds to sub-milliseconds. The microwave pulses generated by the plurality of rf power sources 13 are combined in the time domain by the time domain combining circuit 14, and a long pulse microwave pulse can be generated, as shown in fig. 4. The long-pulse microwave pulse drives the normal-temperature radio-frequency linear accelerator 3, and can energize long-pulse electron beams.
The electron beams comprise a low-energy electron beam and a high-energy electron beam, the low-energy electron beam is an electron beam EB1 directly generated by the high-voltage pulse grid-control hot cathode electron gun, and the high-energy electron beam is an electron beam EB2 generated by the low-energy electron beam EB1 after being energized by the normal-temperature radio-frequency linear accelerator 3.
The normal temperature rf linear accelerator 3 is composed of a plurality of rf resonant cavities 9 for increasing the energy of the electron beam, as shown in fig. 5. The radio frequency resonant cavity 9 is distributed along the axis of the normal temperature radio frequency linear accelerator 3, the time domain synthesis radio frequency power source 4 is electrically connected with the radio frequency resonant cavity 13, the resonant frequency of the radio frequency resonant cavity 9 is the same as the frequency of the time domain synthesis radio frequency power source 4, the resonant frequency of the radio frequency resonant cavity 9 can be an L wave band, an S wave band, a C wave band or an X wave band, and the frequency of the corresponding time domain synthesis radio frequency power source 4 can be an L wave band, an S wave band, a C wave band or an X wave band. The low-energy electron beam EB1 penetrates through the normal-temperature radio-frequency linear accelerator 3 to obtain energy increase delta E, and the energy increase delta E becomes a high-energy electron beam EB2, the delta E is determined by the scale and the performance of the normal-temperature radio-frequency linear accelerator 3, the more the number of the radio-frequency resonant cavities 9 is, the larger the field gradient is, and the larger the energy increase delta E is. The energy gain Δ E in this example is 9-15 MeV. In this embodiment, the type and number of the rf resonant cavities 9 are determined, and the energy increase Δ E can be adjusted by adjusting the intensity of the rf field to meet the X-ray energy requirements for different radiotherapy treatments. The strength of the radio frequency field is positively correlated with the feed-in power of the power source, the power of the power source is adjusted, and the field gradient of the radio frequency field in the room-temperature radio frequency cavity is correspondingly changed.
The X-ray target 6 is a high atomic number material ray target, the electron beam and the high atomic number material interact to generate X-rays, and the X-ray target 6 is also provided with a heat dissipation assembly which is an automatic rotating disk or a cooling water pipe. The core component of the X-ray target 6 is a high atomic number material, i.e., a high Z material, which generates X-rays through the interaction of the electron beam and the high Z material. The X-ray target 6 generally uses tungsten or tantalum which is hard in texture, fast in heat transfer, and high in melting point. In this embodiment, the X-ray target 6 is a tungsten target. Part of the energy of the electron beam is converted into the energy of the X-ray, and the other part is deposited on the X-ray target 6 in the form of thermal energy, so that the X-ray target 6 has a cooling function to prevent the energy of the electron beam deposition from burning out the X-ray target 6. In this embodiment, a water pipe is laid in the target body, and cooling water is introduced to dissipate heat.
The long macro-pulse electron beam bombards the X-ray target to generate X-rays and generate long macro-pulse high-dosage-rate X-rays. The dosage rate can be up to 106Gy/s to 108Gy/s, the dose of a single macropulse can reach 10-80 Gy.
A method for using a miniaturized flash radiotherapy device mainly comprises the following steps:
the high-voltage pulse grid-controlled hot cathode electron gun 1 generates an electron beam with high average current intensity of 1-100mA and long macropulse of 10 microseconds to 10 milliseconds, and the electron beam is injected into a normal-temperature radio-frequency linear accelerator 3 through a first transmission line 2 for energizing to increase the electron energy to 9-15 MeV. Electrons emitted by the normal temperature radio frequency linear accelerator 3 bombard the X-ray target 6 through the second transmission line 5 to generate long-macropulse X-rays with high dosage rate which can reach 106Gy/s to 108Gy/s, the dose of a single macropulse can reach 10-80 Gy. The energy of the rays is adjusted by adjusting the energy of the electron beams, the time length of the rays is adjusted by adjusting the pulse length of the electron beams, the X-ray dose rate is adjusted by adjusting the current intensity of the electron beams, and the adjusted X-rays are incident to a target through a collimator to perform related treatment or experiment.
The "lifetime" of the electron beam can be summarized as follows: the energy is transmitted to a normal-temperature radio-frequency linear accelerator through a first transmission line to obtain energy increase, then the energy is transmitted to an X-ray target through a second transmission line, X-rays are generated through interaction of electron beams and the X-ray target, partial energy of the electron beams is converted into energy of the X-rays, and the other part of the energy is deposited on the X-ray target in the form of heat energy. The technical scheme adopts the combination of the grid-controlled electron gun and the normal-temperature radio-frequency linear accelerator, and compared with a photocathode electron gun, a laser system does not need to be driven; compared with a superconducting linear accelerator, the device does not need a low-temperature system, has simpler structure and smaller scale, is more suitable for the treatment method of 'flash radiotherapy' popularized in hospitals, reduces the purchase cost of equipment and further reduces the treatment cost of patients.
The above detailed description of the present invention is only a preferred embodiment of the present invention, and the scope of the present invention should not be limited thereto, i.e. all the equivalent changes and modifications made according to the scope of the present invention should be covered by the present invention.
Claims (10)
1. A miniaturized flash radiotherapy device, characterized by: the grid-controlled electron gun transmits the electron beam to the normal-temperature radio-frequency linear accelerator through a first transmission line, the normal-temperature radio-frequency linear accelerator transmits the electron beam to the X-ray target through a second transmission line, the electron beam bombards the X-ray target to generate X-rays, and the X-rays irradiate a target to be treated through the collimator.
2. The miniaturized flash radiotherapy device of claim 1, wherein: the pulse time length of the electron beam generated by the grid-control electron gun is adjustable, and the adjustment range is 10 microseconds to 10 milliseconds.
3. The miniaturized flash radiotherapy device of claim 1 or 2, wherein: the grid-controlled electron gun is a microwave grid-controlled hot cathode electron gun or a high-voltage pulse grid-controlled hot cathode electron gun.
4. The miniaturized flash radiotherapy device of claim 3, wherein: the microwave grid-controlled hot cathode electron gun generates an electron beam with high average flow strength of 1-100mA and long macropulse of 10 microseconds to 10 milliseconds, and the electron beam is injected into a normal-temperature radio-frequency linear accelerator through a first transmission line to be energized, so that the electron energy is increased to 9-15 MeV.
5. The miniaturized flash radiotherapy device of claim 3, wherein: the electron beams comprise low-energy electron beams and high-energy electron beams, the low-energy electron beams are electron beams directly generated by a grid-control electron gun, and the high-energy electron beams are electron beams generated after the low-energy electron beams are energized by a normal-temperature radio-frequency linear accelerator.
6. The miniaturized flash radiotherapy device of claim 5, wherein: the electron beam emitted by the normal temperature radio frequency linear accelerator bombards the X-ray target through a second transmission line to generate a long macropulse 106Gy/s-108Gy/s dose rate, the dose of a single macropulse is 10-80 Gy.
7. The miniaturized flash radiotherapy device of claim 1, wherein: the time domain synthesis radio frequency power source is composed of a plurality of radio frequency power sources according to a time domain synthesis circuit.
8. The miniaturized flash radiotherapy device of claim 1, wherein: the normal-temperature radio-frequency linear accelerator comprises a plurality of radio-frequency resonant cavities for increasing the energy of electron beams, the radio-frequency resonant cavities are distributed along the axis of the normal-temperature radio-frequency linear accelerator, and a time domain synthesis radio-frequency power source is electrically connected with the radio-frequency resonant cavities.
9. The miniaturized flash radiotherapy device of claim 1, wherein: the X-ray target is a high atomic number material ray target, and the electron beam and the high atomic number material interact to generate X-rays.
10. The miniaturized flash radiotherapy device of claim 1, wherein: and the X-ray target is also provided with a heat dissipation assembly, and the heat dissipation assembly is an automatic rotating disk or a cooling water pipe.
Priority Applications (1)
| Application Number | Priority Date | Filing Date | Title |
|---|---|---|---|
| CN202021062058.6U CN212214394U (en) | 2020-06-10 | 2020-06-10 | Miniaturized flash radiotherapy device |
Applications Claiming Priority (1)
| Application Number | Priority Date | Filing Date | Title |
|---|---|---|---|
| CN202021062058.6U CN212214394U (en) | 2020-06-10 | 2020-06-10 | Miniaturized flash radiotherapy device |
Publications (1)
| Publication Number | Publication Date |
|---|---|
| CN212214394U true CN212214394U (en) | 2020-12-25 |
Family
ID=73930287
Family Applications (1)
| Application Number | Title | Priority Date | Filing Date |
|---|---|---|---|
| CN202021062058.6U Active CN212214394U (en) | 2020-06-10 | 2020-06-10 | Miniaturized flash radiotherapy device |
Country Status (1)
| Country | Link |
|---|---|
| CN (1) | CN212214394U (en) |
Cited By (1)
| Publication number | Priority date | Publication date | Assignee | Title |
|---|---|---|---|---|
| CN111481840A (en) * | 2020-06-10 | 2020-08-04 | 中国工程物理研究院应用电子学研究所 | Miniaturized flash radiotherapy device |
-
2020
- 2020-06-10 CN CN202021062058.6U patent/CN212214394U/en active Active
Cited By (1)
| Publication number | Priority date | Publication date | Assignee | Title |
|---|---|---|---|---|
| CN111481840A (en) * | 2020-06-10 | 2020-08-04 | 中国工程物理研究院应用电子学研究所 | Miniaturized flash radiotherapy device |
Similar Documents
| Publication | Publication Date | Title |
|---|---|---|
| CN111481840A (en) | Miniaturized flash radiotherapy device | |
| CN111481841A (en) | Flash radiotherapy device | |
| US12027341B2 (en) | Three-dimensional beam forming X-ray source | |
| Linz et al. | What will it take for laser driven proton accelerators to be applied to tumor therapy? | |
| US8045679B2 (en) | Charged particle cancer therapy X-ray method and apparatus | |
| US7940894B2 (en) | Elongated lifetime X-ray method and apparatus used in conjunction with a charged particle cancer therapy system | |
| EP3517172B1 (en) | Neutron capture therapy system | |
| US8625739B2 (en) | Charged particle cancer therapy x-ray method and apparatus | |
| US8229072B2 (en) | Elongated lifetime X-ray method and apparatus used in conjunction with a charged particle cancer therapy system | |
| CN108811297A (en) | A kind of medical proton heavy ion avcceleration | |
| WO2015102681A2 (en) | Methods and systems for rf power generation and distribution to facilitate rapid radiation therapies | |
| US11154729B2 (en) | High brightness electron beam based precise radiation therapy method and system | |
| CN212214394U (en) | Miniaturized flash radiotherapy device | |
| CN107998517B (en) | Neutron capture therapy system | |
| Vretenar et al. | JACOW: Conceptual design of a compact synchrotron-based facility for cancer therapy and biomedical research with helium and proton beams | |
| CN212522747U (en) | Flash radiotherapy device | |
| JP7555399B2 (en) | Ultra-high dose rate radiation therapy equipment | |
| Bagne et al. | Radiation contamination and leakage assessment of intraoperative electron applicators | |
| Lennox | Hospital-based proton linear accelerator for particle therapy and radioisotope production | |
| RU2826821C1 (en) | Ultra-high dose rate radiation therapy device | |
| Jin et al. | Design of a magnetic-modulated target-selective X-ray tube to improve the energy distribution for intraoperative radiotherapy | |
| Kutsaev et al. | Feasibility study of high-power electron linac for clinical X-ray ROAD-FLASH therapy system | |
| WO2023281539A1 (en) | Multi purpose compact apparatus for the generation of high-flux of neutrons, particularly for intraoperative radiotherapy | |
| Kumada et al. | The CBS-the most cost effective and high performance carbon beam source dedicated for a new generation cancer therapy | |
| Minnaar | The first two electron linear accelerators in South Africa |
Legal Events
| Date | Code | Title | Description |
|---|---|---|---|
| GR01 | Patent grant | ||
| GR01 | Patent grant | ||
| TR01 | Transfer of patent right | ||
| TR01 | Transfer of patent right |
Effective date of registration: 20230717 Address after: 621900 No. 1, Kexue City Avenue, Youxian District, Mianyang, Sichuan Province Patentee after: Zhongjiu Flash Medical Technology Co.,Ltd. Address before: 621900 No.64, Mianshan Road, Youxian District, Mianyang City, Sichuan Province Patentee before: INSTITUTE OF APPLIED ELECTRONICS, CAEP |