The X-ray generator includes a booster for boosting a first DC voltage supplied from a voltage source to a second DC voltage higher than the first DC voltage, at least one capacitor for receiving the second DC voltage and generating a charging voltage on the basis of the second DC voltage, a converter for converting the charging voltage into a driving voltage, an X-ray source for receiving the driving voltage and emitting X-rays according to the driving voltage, and a controller for controlling the booster, the converter, and the X-ray source. The controller calculates a cooling time required for cooling the X-ray source to a predetermined temperature or lower, determines the magnitude of the second DC voltage according to the cooling time, and applies the second DC voltage to the capacitor for the cooling time.

CROSS REFERENCE TO RELATED APPLICATION

The present application claims priority to Korean Patent Application Nos. 10-2019-0176499 and 10-2020-0143124, respectively field on Dec. 27, 2019 and Oct. 30, 2020, the entire contents of which are incorporated herein for all purposes by this reference.

BACKGROUND OF THE INVENTION

Field of the Invention

The present invention relates to an X-ray generator.

Description of the Related Art

In recent years, with the development of semiconductors and information processing technology, X-ray radiography is rapidly being replaced with digital radiography (DR) that uses digital sensors. However, X-ray radiography technology is also being continuously researched and developed for specific applications.

An example is intra-oral X-ray radiography which is mainly used in dental clinics.

Intra-oral X-ray radiography is an X-ray imaging technology for obtaining an X-ray image of a limited area in a subject's oral cavity. An X-ray sensor is placed in the subject's oral cavity, X-rays are emitted from an X-ray generator disposed outside the oral cavity, and an X-ray image of a tooth and the surrounding tissue is acquired. These intra-oral X-ray images have the advantage of low distortion, high resolution and sharpness, and relatively low radiation exposure. Therefore, X-ray images are mainly used for implant procedures or root canal treatments that require a high resolution image.

X-ray radiographic devices for intra-oral X-ray radiography are referred to as portable X-ray generators, and in most cases, X-ray images are usually taken with an X-ray radiographic device held in an operator's hand. To increase the convenience and accuracy of intra-oral X-ray radiography and to improve the utilization of X-ray generators, it is required to reduce the weight and size of the X-ray generators for intra-oral X-ray radiography.

SUMMARY OF THE INVENTION

An objective of the present invention is to provide an X-ray generator with a light and compact body and an increased efficiency resulting from a reduced waiting time for X-ray generation.

In order to accomplish the objective of the present invention, there is provided an X-ray generator including: a booster configured to boost a first DC voltage supplied from a voltage source to a second DC voltage higher than the first DC voltage; at least one capacitor configured to receive the second DC voltage and generate a charging voltage from the second DC voltage; a converter configured to convert the charging voltage into a driving voltage; an X-ray source configured to receive the driving voltage and emit X-rays; and a controller configured to control the booster, the converter, and the X-ray source. The magnitude of the second DC voltage is variable. The controller calculates a cooling time required for cooling the X-ray source to a predetermined temperature or lower, determines the magnitude of the second DC voltage according to the cooling time, and applies the second DC voltage to the capacitor during a cooling operation performed for the cooling time.

The booster may include: an input terminal including an inductor connected to the voltage source; an output terminal including a diode connected to the capacitor; and a switching element that turns on and off such that the input terminal and the output terminal are connected to each other or disconnected from each other. The controller may control an ON and OFF cycle of the switching element to control the magnitude of the second DC voltage.

The booster may include an input capacitor connected in parallel between the voltage source and the inductor and an output capacitor connected in parallel between the diode and the capacitor.

The controller may stop the driving voltage from being applied to the X-ray source during the cooling operation performed for the cooling time.

The X-ray generator may further include a user input unit that receives X-ray imaging information including X-ray imaging mode and imaging conditions from a user, and the controller may calculate the cooling time on the basis of the X-ray imaging information.

The X-ray generator may further include a temperature sensor that measures a temperature of the X-ray source, and the controller may calculate the cooling time on the basis of the temperature of the X-ray source.

The present invention has the effect of providing a highly efficient X-ray generator having a light and compact body for convenience of use and being capable of increasing imaging efficiency by minimizing a waiting time for X-ray generation.

DETAILED DESCRIPTION OF THE INVENTION

The above-described objectives, features, and advantages will be more clearly understood with reference to the accompanying drawings and embodiments described below.

In the following description, the specific structural or functional descriptions for exemplary embodiments according to the concept of the present invention are merely for illustrative purposes. The embodiments according to the concept of the present invention may be implemented in various forms and should not be construed as being limited to the embodiments described in the specification of the present application.

Embodiments in accordance with the concept of the present invention can undergo various changes to have various forms, and only some specific embodiments are illustrated in the drawings and described in detail in the present disclosure. While specific embodiments of the present invention are described herein below, they are only for illustrative purposes and should not be construed as limiting to the present invention. Thus, the present invention should be construed to cover not only the specific embodiments but also cover all modifications, equivalents, and substitutions that fall within the concept and technical spirit of the present invention.

It is to be understood that when any element is referred to as being “connected to” or “coupled to” another element, it may be connected directly to or be coupled directly to another element or be connected to or coupled to another element indirectly, having a further element intervening therebetween. In contrast, it should be understood that when an element is referred to as being “directly coupled” or “directly connected” to another element, there are no intervening elements present between them. Other expressions for describing the relationship between the constituent elements, such as “between” and “directly between” or “adjacent to” and “directly adjacent to” should be interpreted in the same manner.

Hereinafter, modes of the present invention will be described in detail with reference to the accompanying drawings. Through the drawings, like reference symbols denote like elements.

FIG. 1is a perspective view illustrating the overall construction of an X-ray generator according to one embodiment of the present invention.

As illustrated, an X-ray generator according to one embodiment of the present invention includes: a main body10in which a power supply unit, a controller, a converter, an X-ray source, etc., which will be described later, are mounted; a handle12provided on one side of the main body10to allow a user to hold the X-ray generator for use; an emission switch14provided on one side of the main body10, preferably near the handle12, for user manipulation to emit X-rays; an emission port20provided on one side of the body10and configured to emit X-rays; a shielding plate22provided along the periphery of the emission port20to minimize exposure of a user of the X-ray generator to the X-rays emitted from the X-ray generator; an instrument panel32provided on one side of the main body10and configured to display operation-related messages and various types of X-ray imaging information including an imaging mode and imaging conditions; a user input unit32such as a button or a dial for user manipulation; and a power connector36provided on one side of the handle12, connected to an external power source to receive external power as a voltage source or as a battery charging power source.

Therefore, the user holds the handle12and lifts the main body10to set an appropriate imaging mode and imaging conditions using the instrument panel32and the user input unit34, aims the emission port20of the main body10at an X-ray application target position, and manipulates the emission switch14to emit X-rays to the X-ray application target position through the emission port20. The shielding plate22blocks X-rays scattering backward to minimize the exposure of the user to X-rays. The power connector36is connected to an external power source through an adapter, etc. to receive external power as a voltage source or a battery charging power source.

FIG. 1is only an example of an X-ray generator according to the present invention, and the shape or structure of the X-ray generator may vary.

FIG. 2is a diagram illustrating the internal construction of the X-ray generator according to the present invention and shows only parts necessary for description of the present invention.

As illustrated, the X-ray generator according to the present invention includes a power supply unit100, a controller200, a power converter300, and an X-ray source400.

The power supply unit100includes a voltage source102, a switch104, a booster106, and a capacitor108.

The voltage source102may be a battery. Specifically, it may be a commercial primary battery or a commercial secondary battery. The battery may be one battery or two or more batteries. When the battery is a primary battery and it is exhausted, the primary battery will be replaced with a new battery. When the battery is a secondary battery and it is exhausted, the battery will be charged to be used as a voltage source again. Under the premise that the battery is a commercial primary battery, the only configuration that needs to be replaced or changed in the X-ray generator according to the present invention is the battery. As will be described later, the requirement for the voltage source in the X-ray generator according to the present invention is a voltage source that can provide a low voltage (for example, 2.5V to 4.2V). The present invention uses a commercial primary battery or a commercial secondary battery as a voltage source, thereby minimizing the size of the X-ray generator and minimizing the charging time of the battery.

For reference, the voltage source102of the X-ray generator according to the present invention may be an external power source instead of a battery. When the voltage source102is a secondary battery, the secondary battery may be charged with an external power source.

The switch104controls electrical connection and disconnection between the voltage source102and the capacitor106, thereby blocking or allowing a first DC voltage supplied from the voltage source102to the booster106.

The switch104turns on and off according to an ON/OFF control signal received from the controller200.

The switch104receives an ON control signal from the controller200and electrically connects the voltage source102and the booster106to each other. When the voltage source102and the booster106are connected via the switch104, the first DC voltage output from the voltage source102is boosted to a second DC voltage by the booster106, and the capacitor108is charged with the secondary DC voltage. When the switch104receives an OFF control signal from the controller200, it electrically disconnects the voltage source102and the booster106from each other. When the switch41turns off so that the voltage source102and the booster103are disconnected from each other, the voltage supply from the voltage source102is cut off, and the charging of the capacitor104is stopped.

Although the description in the present disclosure and the illustration in the illustration show that the switch104is interposed between the voltage source102and the booster106, it is possible that the switch102is positioned between the booster106and the capacitor108.

The booster106is connected to the voltage source102via the switch104, and the driving of the booster106is controlled according to the ON/OFF state of the switch104. The booster106converts the first DC voltage output from the voltage source102into the second DC voltage used to charge the capacitor108. For example, the booster106receives a low voltage within a range of from 2.5 V to 4.2 V from the voltage source102as the first DC voltage, boosts the first DC voltage to a high voltage in a range of from 24 V to 30 V (0.15 A) as the second DC voltage suitable for charging the capacitor108. To this end, the booster106includes a DC/DC booster converter circuit.

FIG. 3is a circuitry diagram illustrating the power supply unit of the X-ray generator according to the present invention in a case where the power supply unit100includes a voltage source102and a capacitor108composed of a plurality of series-connected capacitor elements. The booster106receives the first DC voltage (in a range of 2.5 V to 4.2 V) which is an output voltage of the voltage source102and converts the first DC voltage into the second DC voltage (in a range of 24 V to 30 V) which is a voltage suitable for charging the capacitor108.

FIG. 4is a circuitry diagram illustrating the booster106of the X-ray generator according to the present invention. The booster106includes an input terminal106aincluding an inductor In connected in series with the voltage source102, an output terminal106bincluding a diode D connected in series with the capacitor108, and a switching element SW for controlling the electrical connection and disconnection between the input terminal106aand the output terminal106b. An input capacitor Cin is connected in parallel between the voltage source102and the inductor In, and an output capacitor Cout is connected in parallel between the diode D and the capacitor108.

When the switching element SW is turned on, the inductor In is supplied with current. Conversely, when the switching element SW is turned off, the current is discharged from the inductor In, and the second DC voltage higher than the first DC voltage is transmitted to the capacitor108. The magnitude of the second DC voltage varies depending on the ON and OFF cycle of the switching element SW, and the controller200to be described later controls the ON and OFF cycle of the switching element SW to adjust the magnitude of the second DC voltage.

Referring toFIG. 2, the capacitor108is charged with the secondary DC voltage supplied from the booster106. The capacitor108supplies a charging voltage to the converter300for X-ray emission. The capacitor108is composed of at least two capacitor elements connected in series. For example, assuming that electric power of 65 kV, 3 mA, and 1 second is required to drive the X-ray source400, the capacitor108is composed of 2 to 4 capacitor elements having a capacitance of 25F to 30F. Assuming that the capacitor element is an electrolytic capacitor having a size of about 12.5 ø×20 mm, the size of the capacitor108may be about 25×50×20 mm.

In the case of a comparative example in which only the first DC voltage output from the battery is used, since it is necessary to instantaneously apply a high voltage to the X-ray source400, the lifespan and durability of the battery are reduced. In the case of another comparative example in which only the DC voltage of a specially designed high-power battery is used, there is a problem in that the cost of the X-ray generator increases. In the case of a further comparative example in which only the DC voltage of a high-capacity capacitor is used, since the energy density of the capacitor is low, the size of the power supply unit is increased. Therefore, it is inconvenient to use the X-ray generator. In the case of a further comparative example in which only the DC voltage of a low-capacitance capacitor is used, there is a problem in that it is inconvenient to use the X-ray generator due to the hassle of charging the battery of the X-ray generator each time a radiographic operation is performed.

However, in the present invention, the power supply unit100is composed of the voltage source102implemented with a high energy-density battery and the capacitor104implemented with multiple capacitor elements connected in series to output a high voltage. Since the size of the power supply unit is reduced, the X-ray generator is accordingly light and compact, and thus the imaging efficiency and convenience of the X-ray generator are increased.

The controller200controls the overall operation of the X-ray generator apparatus according to the present invention.

Referring toFIG. 1, the controller200controls the converter300to adjust the magnitude of the voltage supplied to the X-ray source400when the user inputs X-ray imaging information such as an imaging mode and imaging conditions through the instrument panel32and the user input unit34. When the user inputs an X-ray emission command through the emission switch14, the controller200controls the converter300to supply the driving voltage to the X-ray source400so that X-rays are emitted. When X-ray imaging is completed, the controller200controls the converter300not to supply the driving voltage to the X-ray source400, thereby stopping the X-ray emission. However, during a cooling operation performed for a cooling time which is a predetermined period of time after the X-ray emission is performed, even though the user inputs an X-ray irradiation command through the emission switch14, the controller200controls the converter300not to supply the driving voltage to the X-ray source400so that the X-ray source400can be cooled to a preset temperature or below.

In particular, the controller200of the X-ray generator according to the present invention calculates the cooling time on the basis of the temperature of the X-ray source400or on the X-ray imaging information after the X-ray imaging information is input through the user input unit34or after the X-ray emission is performed by the X-ray source400. In addition, the controller controls the driving time of the booster106and the magnitude of the second DC voltage such that the charging voltage of the capacitor108becomes equal to or higher than a predetermined voltage within the calculated cooling time. In addition, the controller200displays an X-ray emission ready message notifying that it is ready for X-ray imaging, on the instrument panel32when the temperature of the X-ray source400is lowered to the predetermined temperature or below after the charging of the capacitor108is completed.

In addition, the controller200of the X-ray generator according to the present invention monitors the status of the power supply unit102and the X-ray source400and provides an appropriate notification message to the user through the instrument panel32on the basis of the monitoring result. For example, the controller200measures the residual voltage of the battery and displays a battery replacement request message on the instrument panel32to request the replacement of the battery when the residual voltage of the battery is equal to or less than a preset level. In addition, the controller200checks whether the capacitor108is overcharged or short-circuited and displays a capacitor abnormality notification message on the instrument panel32when overcharging or short-circuit is detected. In addition, the controller measures the real-time temperature of the X-ray source400and displays the real-time temperature on the instrument panel32.

FIG. 5is a diagram illustrating the construction of the controller of the X-ray generator according to one embodiment of the present invention.

As illustrated inFIG. 5, the controller200includes a logic operation circuit210equipped with a series of control algorithms for X-ray radiography, a memory240for storing X-ray imaging information used to calculate a cooling time or for storing cooling time information for each temperature of the X-ray source400, a power supply unit100, a power supply management module230for monitoring the voltage source102and the capacitor108, a temperature sensor250for measuring the temperature of the X-ray source400, and a timer260for counting time for the cooling time.

The operations of the controller200will be described in detail below.

Referring toFIG. 2, the converter300boosts the charging voltage supplied from the capacitor108to a driving voltage (for example, 65 kV, 3 mA) for driving the X-ray source400and supplies the driving voltage to the X-ray source400.

The converter300includes an inverter302, a primary booster304and a secondary booster306. For example, the primary booster304includes a transformer, and the secondary booster304includes a Cockcroft-Walton generator. The Cockcroft-Walton generator is composed of an n-fold voltage rectifier circuit or a Cockcroft multiplier and rectifier circuit.

The converter300is controlled by the ON/OFF control signal transmitted to the inverter302from the controller200. When the ON control signal is transmitted to the inverter302from the controller200, the converter300converts the charging voltage supplied from the capacitor108of the power supply unit100into a driving voltage and supplies the driving voltage to the X-ray source400so that the X-ray source400emits X-rays according to the driving voltage.

After X-rays are emitted in accordance with the X-ray imaging information, the controller200cuts off the driving voltage supplied to the X-ray source by transmitting the OFF control signal to the inverter302so that the X-ray emission of the X-ray source is stopped.

The X-ray source400receives the driving voltage from the converter300, generates X-rays, and emits the X-rays to a subject. The X-ray source400is a field emission X-ray source including a cathode electrode having an emitter, an anode electrode having an X-ray target surface, and a gate electrode controlling the field emission of the emitter.

FIGS. 6 and 7are flowcharts showing a method of driving an X-ray generator, according to the present invention. The driving method of the X-ray generator and specific operations of the controller will be described with reference toFIGS. 1 to 5andFIGS. 6 and 7. The present disclosure provides two embodiments for the driving method of the X-ray generator according to the present invention. The operations that are common between a first embodiment and a second embodiment will be described in a section labeled “First Embodiment”, and only a difference between the first embodiment and the second embodiment will be described in a section labeled “Second Embodiment” below.

First Embodiment

In order to drive the X-ray generator according to the present invention, the controller200controls the power management module230to check whether the charging voltage of capacitor108is equal to or greater than a predetermined voltage value (ST10).

When the charging voltage of the capacitor108is higher than or equal to the predetermined voltage value, the controller200turns off the switch104of the power supply unit100to block the second DC voltage so that the second DC voltage cannot be transmitted to the capacitor108, thereby preventing overcharging of the capacitor108(ST15). Conversely, when the charging voltage of the capacitor108is lower than the predetermined voltage value, the controller200turns on the switch104of the power supply unit100so that the second DC voltage can be transmitted to the capacitor108, thereby charging the capacitor108with the second DC voltage (ST22). The magnitude of the second DC voltage is preset.

The controller200checks the temperature of the X-ray source400with the use of the temperature sensor250(ST20). When the temperature of the X-ray source400is higher than a predetermined temperature, the controller200maintains the cooling operation by which the transmission of the driving voltage from the converter300to the X-ray source400is blocked even though the user inputs an X-ray emission command through the emission switch14(ST25). On the other hand, when the temperature of the X-ray source400is lower than the predetermined temperature, the controller200displays the X-ray emission ready message on the instrument panel32and waits for a user's X-ray emission command (ST30).

Subsequently, when the user inputs X-ray imaging information such as imaging mode and imaging conditions through the user input unit34and inputs an X-ray emission command by pushing the X-ray emission switch14, the controller200adjusts the driving voltage of the converter300and transmits the adjusted driving voltage to the X-ray source400so that an X-ray imaging operation is performed according to the X-ray imaging information, and the X-ray source400emits X-rays (ST35, ST40). When an X-ray emission command is not input, the controller200waits for an X-ray emission command input through the X-ray emission switch14(ST30).

In addition, when the X-ray imaging information is input through the user input unit34, the controller200calculates an estimated cooling time on the basis of the cooling time information stored in the memory240. The cooling time information is configured in the form of a table in which each piece of the X-ray imaging information is associated with a specific cooling time (ST50).

For reference, the first embodiment differs from the second embodiment described below in that the charging time for the capacitor108is maximized and, to this end, the controller200calculates the estimated cooling time on the basis of the X-ray imaging information input through the user input unit34. Therefore, the controller200calculates the estimated cooling time after the X-ray imaging information is input through the user input unit34. That is, the calculation of the estimated cooling time is performed before the X-ray emission is performed by the X-ray source400or at the time when the X-ray emission of the X-ray source400is performed.

Next, the controller200determines the magnitude of the second DC voltage for completing the charging of the capacitor108within the calculated cooling time (ST55). The cooling time and the magnitude of the second DC voltage have an inverse linear correlation. Therefore, a predetermined function for calculating the magnitude of the second DC voltage according to the cooling time is stored in the memory240, or second DC voltage information that is present in the form of a table in which second DC voltages and cooling times are respectively matched is stored in the memory240.

Next, the controller200controls the booster106to boost the first DC voltage to the second DC voltage having a magnitude that is determined in step ST55. To this end, the controller200adjusts the ON and OFF cycle of the switching element SW of the booster106. The ON time and the magnitude of the second DC voltage are in a proportional correlation.

On the other hand, when the X-ray emission in accordance with the desired X-ray imaging information is completed, the controller200blocks the driving voltage transmitted from the converter300to the X-ray source400, thereby stopping the X-ray emission (ST60).

Next, the controller200maintains the cooling operation for the cooling time. During the cooling operation, even though the user inputs an X-ray emission command through the emission switch14, the driving voltage output from the converter300is blocked not to be input to the X-ray source400. In this case, the controller200displays a message notifying that the cooling operation is performed on the instrument panel32. Preferably, the remaining cooling time may be displayed together with the message on the instrument panel32(ST25).

Next, when the cooling time is over, the controller200performs step ST10and the subsequent steps again.

In the present embodiment, the controller200determines a cooling time for each imaging condition included in the X-ray imaging information and the magnitude of the second DC voltage at the same time as the X-ray emission and controls the second booster106to output the second DC voltage having the determined magnitude. With this operation, the maximum charging time for the capacitor108can be secured, which minimizes the waiting time for X-ray generation. The waiting time occurs due to an insufficient charging voltage for the capacitor108after the cooling time is over.

Second Embodiment

A method of driving the X-ray generator, according to the second embodiment, is substantially the same as the method of the first embodiment in terms of operations of step ST10to step ST40. However, the driving method of the second embodiment differs from the driving method of the first embodiment in that the controller200measures the real-time temperature of the X-ray source400with the use of the temperature sensor250after the completion of the X-ray emission (ST60) instead of calculating the cooling time for each imaging condition included in the X-ray imaging information when starting the X-ray emission of step ST40.

Next, the cooling time for each temperature of the X-ray source400is calculated on the basis of the real-time temperature of the X-ray source400. To this end, per-temperature cooling time information that is a table in which temperatures of the X-ray source400and cooling times are respectively matched is stored in the memory240.

Next, the controller200determines the magnitude of the second DC voltage for charging the capacitor108to a predetermined voltage value or more during the calculated cooling time (ST55) and adjusts the ON and OFF cycle of the switching element SW of the booster106such that the booster106outputs the second DC voltage having the magnitude determined in step ST55(ST60).

The second embodiment is advantageous over the first embodiment in that the cooling time can be determined according to the real-time temperature of the X-ray source400but is disadvantageous in that the charging time for the charging the capacitor108is slightly reduced because the calculation of the cooling time is performed after the completion of the X-ray emission.

Therefore, preferably, the controller200controls the booster106to output a second DC voltage having a predetermined magnitude at the same time as the time at which the X-ray emission is performed, and, after step ST55, the controller200adjusts the second DC voltage to have a magnitude determined in step ST55.

While the present invention has been particularly illustrated and described with reference to exemplary modes, it is to be understood that the present invention is not limited to the disclosed exemplary modes and that modifications thereto are possible.

Therefore the scope of the present invention should not be limited to the preferred embodiments but should be defined by the claims and their equivalents.

EXPLANATION OF REFERENCE NUMERALS