DESCRIPTION

Field of the invention

In its most general aspect, the present invention refers to the technical field of radiology and, specifically, to a radiological device.

More specifically, the invention refers to a so-called monoblock or a power monoblock radiological device for digital applications, comprising an X-ray insert and a high-voltage power supply which are combined with each other inside the same case or casing.

State of the art

In the technical field of radiology, radiological devices are known to comprise an X-ray insert and a high-voltage power supply electrically connected to the X-ray insert for applying a difference of potential to the X-ray insert.

The X-ray insert, which is also known as the X-ray tube, and the high-voltage power supply may be arranged in two separate cases or may be mounted in a single case.

In the first case, since the respective cases are separated from each other, the X- ray insert and the high-voltage power supply are electrically connected to each other by suitable cables which ensure the insulation of the high voltage to earth, and which precisely allow them to be used even when these cases are physically distant from each other.

However, in the case of two separate cases, we are referring of tubehead for the case comprising the X-ray insert, and of generator for the one comprising the high- voltage power supply, whereas for the second case, in which only one case is provided, we are referring to monoblock.

It should be noted that we are referring more correctly of tubehead, of high- voltage generator or generator, and monoblock when the respective case is filled with dielectric oil in order to ensure electric insulation between the inner parts of the case and outwards. In particular, the X-ray insert and high-voltage power supply are immersed in dielectric oil which ensures electric insulation between high voltage parts and zero potential parts such as the aforesaid case, which for this purpose is provided with an earth connection.

Briefly, a high-voltage power supply essentially comprises a transformer and a rectifying circuit or a multiplier consisting of diodes and capacitors, which develops a direct voltage, i.e. a constant difference of potential (d.o.p.) between its poles, whereas an X-ray insert essentially comprises a filament and a small disc inserted in a glass “ampoule”, in which, in use, the former is connected to a lower potential and the latter to a higher potential. That is, filament and small disc constitute two poles connected to the aforesaid high-voltage power supply, cathode and anode respectively.

In detail, the filament consists of a small metal spiral which is passed through by electric current and is heated by joule effect, which is mounted inside a metal cup known as the filament optics, whereas the small disc is a metal disc with a high melting temperature.

By thermoelectric effect, the filament emits electrons which are accelerated by the electric field produced by the applied difference of potential and which, through the optics of the filament, are conveyed to a single point on the small disc (or, in any case, a single area), called focus. Once the small disc is reached, the electrons interact with the material of which the same small disc is made and, in practice, are “braked” by it.

The deceleration of the electrons causes the X-radiation, called braking radiation, which corresponds to the radiation mainly used in radiology.

The two-case solution allows to use X-ray inserts larger than those used in a monoblock and, consequently, is more suitable for applications in which there are particularly high powers or workloads.

It should be noted, in fact, that only about 1% of all the energy used in this process is converted into X-rays, the remaining part is transformed into heat which is transferred to the small disc which, by cooling down, transfers the accumulated heat to the tubehead or monoblock, on a case by case basis.

In this regard, it should be noted that, generally speaking, the larger an X-ray insert, the greater its heat capacity and the greater its continuous heat loss which characterises it.

In detail, the heat capacity of the X-ray insert, or more properly the maximum anode heat load, defines the maximum amount of energy that can be stored by the anode or, more precisely, defines the energy that the small disc can store until it reaches the maximum allowable temperature for the anode mass and, therefore, this parameter must be well evaluated depending on the type of application.

The continuous heat loss of the X-ray insert, or more specifically of the anode, rather expresses the maximum amount of heat that the anode is able to continuously dissipate in the unit of time. In other words, the continuous heat loss of the anode is the average power that is transferred through radiation to the external environment by the anode when its temperature is maintained at an admissible value for continuous operation.

It should be noted that, according to the known art, continuous heat loss values around 500/600 W are on average to be considered high.

As regards the difference of potential applicable to an X-ray insert between anode and cathode, it should be mentioned that, in principle, tubeheads with X-ray inserts having large dimensions can operate with voltages up to 150 kV whereas, in monoblocks, the working voltage does not exceed 120 kV-125 kV.

Another significant magnitude when forming images in radiology, in addition to the difference of potential, is the current flowing through the X-ray insert, which is generated by electrons moving from cathode to anode and impacting the small disc.

In detail, for a given X-ray insert, the applied difference of potential and the current flowing through it determine the power supplied to the X-ray insert. At the same power supplied, the difference of potential and current may be different: e.g., making a continuous exposure at 150 kV voltage and 3.33 mA current or at 50 kV voltage and 10 mA current, the power supplied to the X-ray insert is in both cases 500 W, but the working voltages and currents are very different. If the continuous heat loss is assumed to be 500 W, this means that when the anode reaches the maximum admissible operating temperature, no more than 500 W can be supplied continuously so that the temperature does not increase (regardless of whether one is working at 50 kV/10 mA or 150 kV/3.33 mA). Since the power applied to the X-ray insert is almost entirely supplied to the anode in the form of heat, the continuous heat loss of the anode results in a limit to the continuously applicable power; if, on the other hand, the exposure to X-rays does not take place continuously but for a short interval of time, the maximum applicable power will be higher and, by definition, corresponds to the rated anode power defined for a single load and an exposure with a duration of 100 ms. Of course, the maximum power that can be executed at different intervals and for multiple repetitions should be evaluated depending on the repetitions and duration of the exposures.

In accordance with the above, if the currents are low, longer times are needed to achieve a good radiographic image but, if the times are short, at the same voltage, the current and, therefore, the power (Vxl) of the exposure should be increased.

A monoblock, therefore, offers more limited performance, since the power output (voltage and current) and the maximum anode heat load are generally lower than those achieved by solutions comprising a tubehead and, therefore, the scope of use of a monoblock is currently confined to applications with low to medium powers although, as a whole, a monoblock allows to simplify the structural and, therefore, mechanical implementation of radiological devices, especially due to the use of a single case instead of two separate cases.

It should be added that the development of systems for digital acquisition and processing of images now allows a three-dimensional reconstruction of the parts involved in the X-ray investigation and also allows to highlight certain areas or organs that can be difficult to distinguish in classic or standard radiographs.

In this regard, from the point of view of the intensity and duration of emission, the modes for executing a classic or standard radiological exposure may be divided into Radiography and Fluoroscopy.

Radiography is characterised by providing short duration exposures (in the order of hundreds of ms) with high currents (tens or hundreds of mA), taking advantage of the maximum power available for a given X-ray insert.

Fluoroscopy is characterised by providing continuous and long exposures, with low currents (in the order of magnitude of mA) in order to remain below the continuous heat loss value of the X-ray insert or, in general, of the X-ray device (monoblock or tubehead) in which it is inserted.

The increasing application of digital acquisition and computer processing technology has led to a widespread use of X-ray sources, both monoblocks and tubeheads, in pulsed mode in so-called Pulsed Fluoroscopy, which is characterised by providing short exposures (in the order of tens of ms) repeated in rapid succession with powers higher than classic Fluoroscopy, i.e. non-pulsed, and generally lower than those used in Radiography. In practice, in Pulsed Fluoroscopy, pulses of intensity higher than classic Fluoroscopy are generated, which are synchronised with the acquisition times of designated detectors or sensors. In addition, as with Radiography, also in Pulsed Fluoroscopy, the power of the single pulse may be greater than the continuous heat loss threshold, but because the pauses between one pulse and the next are very short and do not allow the total heat loss, there will still be heat storage on the anode.

In particular, in Pulsed Fluoroscopy, the acquisition of a sequence of images of the same patient from multiple angles in short times is often required in order to achieve the desired results, as it takes place in Tomography, in which an X-ray source and a detector move around a patient while staying opposite each other.

In this regard, in order to generate an X-ray pulse, i.e. an emission lasting a short and defined time interval, the power supply of the high-voltage power supply is usually removed at a given instant, i.e. the latter is operated in ON/OFF mode by removing and supplying the power supply necessary to generate the high voltage at the ends of the X- ray insert.

This way, the difference of potential at the ends of the X-ray insert is lost and thus the phenomenon originating the X-rays is stopped; more precisely, signals with rather short rise times but more or less long fall times are obtained, because they are governed by the capacitor discharge phenomenon. In this second step, there is the emission of X-rays that are not useful for examinations’ purposes and unnecessarily increase the amount of radiation to which the user, i.e. in this case a patient, is subjected. In practice, the voltage at the ends of the X-ray insert at the beginning of the rays’ pulse reaches the set value in about 1 ms, while at the end of the rays’ pulse, the time required for it to drop to zero is governed by the discharge of the electric circuitry in the high-voltage power supply and may have durations of several milliseconds, depending on the parameters set.

The emission of X-rays occurs until the difference of potential at the ends of the X-ray insert is almost zero so, inconveniently, there is some emission of X-rays even during the so-called power-off transient of the circuit or high-voltage power supply. Disadvantageously, this last emission of X-rays is an unwanted emission because it unnecessarily increases the exposure of patients to X-rays with no advantage in creation of the image; in this regard, see figure 1, which depicts what is described above, and figure 2, which shows a corresponding ideal, hence theoretical, state.

In addition, the pulsed mode generally results in increased electrical stress in high-voltage circuits, due to the repetition of charging and discharging transients with each pulse.

In summary, in order to have a good image in the case of pulsed mode, due to the shortness of the pulses, the power of the emission must be higher than classic Fluoroscopy, and so the amount of heat transferred to the X-ray insert in the unit of time is also greater than in the case of classic Fluoroscopy.

Therefore, since the amount of heat that an X-ray insert can store and the maximum working voltage also depend on its dimensions, in applications where pulsed mode is required, there is a need to use X-ray inserts that have dimensions greater than those used in classic Fluoroscopy, which results in a general orientation toward tubehead and generator solutions that, however, disadvantageously involve a more complex structure of the radiological device as a whole, and a more difficult use by the user, compared to monoblocks.

Disadvantageously, in accordance with the known art, the pulsed mode can be executed with monoblocks in a limited number of applications, such as a head CT scan, whereas for some categories of examinations, the monoblocks’ characteristics are not adequate, especially when considering examinations on obese patients where it is necessary to take advantage of particularly penetrating X-rays.

In this regard, it should be recalled that X-rays are all the more penetrating the higher their energy and this is related to the working voltage of the X-ray tube which, according to the known art, does not exceed 120 kV-125 kV in the monoblocks, as mentioned above.

Summary of the invention

The technical problem underlying the present invention has been that of providing a radiological device adapted to digital applications, which has technical characteristics such as to overcome one or more of the drawbacks mentioned above with reference to the known art.

In accordance with the invention, the aforesaid problem is solved by a so-called monoblock or power monoblock radiological device, comprising: a casing, an X-ray insert for the emission of X-rays, a high-voltage power supply electrically connected to the aforesaid X-ray insert for applying a difference of potential to the X-ray insert, and a dielectric medium, wherein the aforesaid X-ray insert, and the aforesaid high-voltage power supply, are immersed in said dielectric medium within the aforesaid casing, wherein the aforesaid X-ray insert comprises a filament and a small disc, wherein during the emission of X-rays, the aforesaid filament constitutes a cathode as it is connected to a lower potential of the aforesaid high-voltage power supply (negative pole) and the aforesaid small disc constitutes an anode as it is connected to a higher potential of the aforesaid high-voltage power supply (positive pole), which is characterised in that the aforesaid X-ray insert has a maximum power of about 50 kW and a maximum working voltage (maximum d.o.p.), between cathode and anode, of about 150 kV, in that it comprises stopping means for stopping the emission of X-rays, which are adapted to interpose a potential lower than the aforesaid lowest potential between the aforesaid filament and the aforesaid small disc, without varying the aforesaid difference of potential applied to the aforesaid X-ray insert, and in that it comprises an optoelectronic system adapted to make a nonelectrical optical connection for transmitting, internally to the aforesaid casing, an activation or deactivation signal to the aforesaid stopping means, which is imparted by a user.

In practice, in order to generate the aforesaid lower potential, the present monoblock comprises a dedicated circuit that produces a difference of potential independent of that applied to the aforesaid X-ray insert by means of the aforesaid high- voltage power supply for the emission of X-rays.

Preferably, the aforesaid high-voltage power supply comprises a transformer or first transformer and a rectifying circuit, or multiplier, comprising diodes and capacitors, which develops a direct voltage, i.e. a constant difference of potential between its poles, between 40 kV and 150 kV.

Preferably, the aforesaid high-voltage power supply comprises insulating means arranged between the so-called primary and the so-called secondary of the aforesaid first transformer, wherein preferably the aforesaid insulating means comprise polypropylene spools.

Preferably, the aforesaid stopping means comprise: a stopping element adapted to be biased to the aforesaid potential lower than the potential of the aforesaid filament, a transformer or second transformer, a voltage rectifier stage connected to the aforesaid transformer, and a switch or first switch, wherein a higher-potential output of the aforesaid rectifier stage is connected to the aforesaid filament and a lower-potential output of the aforesaid rectifier stage is connected, through the aforesaid switch, to the aforesaid stopping element, wherein when the aforesaid switch is open, the aforesaid device emits X-rays, and when the aforesaid switch is closed, the aforesaid stopping element is biased to the aforesaid potential lower than the potential of the aforesaid filament and the aforesaid device does not emit X-rays.

Preferably, the aforesaid stopping means and, in particular, the aforesaid voltage rectifier stage, when the aforesaid switch is closed, generate a constant difference of potential between about -2 kV and about -3 kV.

Preferably, the aforesaid stopping means comprise insulating means arranged between the so-called primary and the so-called secondary of the aforesaid second transformer, wherein the aforesaid insulating means preferably comprise polypropylene spools arranged between the primary and secondary of the aforesaid second transformer.

Preferably, as a whole, between the primary and secondary of the aforesaid second transformer, insulation equal to at least the maximum expected difference of potential between the aforesaid stopping element and the earth is provided.

Preferably, the aforesaid device comprises a second switch, wherein when the aforesaid first switch is open, the aforesaid second switch is closed, and when the aforesaid first switch is closed, the aforesaid second switch is open, and wherein the aforesaid second switch connects the aforesaid stopping element to the aforesaid filament (cathode) thus zeroing, when closed, the difference of potential between the aforesaid stopping element and the aforesaid filament.

In accordance with the invention, the aforesaid second switch makes it particularly quick to switch the potential of the aforesaid stopping element from the aforesaid potential lower than zero with respect to the cathode, thus zeroing the difference of potential between the aforesaid stopping element and the aforesaid filament when it is desired to emit X-rays.

In practice, in accordance with the invention, the aforesaid first switch and the aforesaid second switch are actuated by the same control.

Preferably, the aforesaid stopping element is a so-called grid interposed between the aforesaid filament and the aforesaid small disc, or it is a metal cup within which the aforesaid filament is arranged.

In accordance with the invention, the aforesaid stopping element is in practice a third electrode positioned in the X-ray insert, which when is biased to a voltage lower than the aforesaid cathode, repels the electrons moving toward the aforesaid anode, thus preventing the emission of X-rays and, therefore, in the scope the present invention, the expressions “grid electrode”, “grid” and “third electrode” are equivalent to each other and can be used to identify the aforesaid stopping element in an equivalent way.

In other words, in accordance with the invention, in order to eliminate unwanted radiation emission, e.g. in the case of digital applications operating in pulsed mode, the high-voltage power supply is not operated in ON/OFF mode but the present device allows the creation of an electric field in the proximity of the filament, which is opposite the electric field produced between the anode and the cathode, i.e. between the same filament and the small disc of the X-ray insert. This way, when desired, the electrons are prevented from accelerating and ending up impinging the small disc, thus generating X-radiation by braking.

In practice, in accordance with the invention, the aforesaid user-controlled optoelectronic system is active on the aforesaid switches and, in particular, in the case that a single switch (first switch) is provided, the optoelectronic system, as needed, controls the closure or opening of the aforesaid first switch, whereas in the case the second switch is also provided, the optoelectronic system, as needed, controls the closure of the first switch and the opening of the second switch, or controls the opening of the first switch and the closure of the second switch, wherein in both cases, therefore in the presence of one or two switches, when the aforesaid first switch is open, the aforesaid X-ray insert emits X-rays and when the aforesaid first switch is closed, the aforesaid X-ray insert does not emit X-rays even if actively connected to the aforesaid high-voltage power supply.

Preferably, the aforesaid optoelectronic system comprises a transmitter, a receiver and an optical fibre extended between the aforesaid transmitter and the aforesaid receiver.

Preferably, the aforesaid transmitter is an emitting diode.

Preferably, the aforesaid receiver is a photo-transistor.

Advantageously, the aforesaid optoelectronic system ensures sufficient insulation with respect to earth. Preferably, the aforesaid cathode is at a potential lower than the earth potential and the aforesaid anode is at a potential higher than the earth potential.

Preferably, the aforesaid difference of potential is equally divided, with respect to earth, between the aforesaid anode and the aforesaid cathode and, therefore, preferably the aforesaid anode operates at a positive potential between +20 kV and +75 kV and the aforesaid cathode operates at a negative potential between -20 kV and -75 kV.

Preferably, the aforesaid device comprises means intended for cooling the aforesaid dielectric medium, wherein, more preferably, the aforesaid dielectric medium comprises or consists of a mineral oil, a silicone oil or a synthetic oil.

Preferably, the aforesaid means intended for cooling the aforesaid dielectric medium, which are also identified herein as cooling means, comprise a heat exchanger, more preferably a heat exchanger selected from the group comprising dual-circuit heat exchangers, such as plate heat exchangers and pipe heat exchangers, radiators, mechanically ventilated radiators.

Preferably, the aforesaid heat exchanger comprises at least one first circuit for circulating the aforesaid dielectric medium and, possibly, a second circuit adapted to the circulation of a cooling fluid coming from a cooling system external to the present radiological device, wherein the aforesaid first circuit and the aforesaid second circuit are in thermal contact with each other.

Preferably, the aforesaid thermal contact between the aforesaid first circuit and the aforesaid second circuit takes place within the aforesaid casing, however, the possibility of providing the aforesaid thermal contact between the aforesaid first circuit and the aforesaid second circuit at least partially external to the aforesaid casing is not excluded, just as the cooling of the aforesaid dielectric medium takes place outside the aforesaid casing in the case of radiators and mechanically ventilated radiators. In the latter case, the aforesaid hot dielectric medium is pumped into the radiator and the heat exchange takes place outside the aforesaid casing with surrounding air possibly blown by appropriate fans. Therefore, in accordance with the above, the aforesaid first circuit may possibly be at least partially extended outside the aforesaid casing. Preferably, therefore, the aforesaid device comprises a pump for circulating the aforesaid dielectric medium in the aforesaid heat exchanger, wherein more preferably the aforesaid pump is arranged within the aforesaid casing. In practice, the aforesaid pump is active on the aforesaid first circuit in which the aforesaid dielectric medium circulates.

Preferably, the aforesaid second circuit comprises an inlet opening and an outlet opening for the aforesaid cooling fluid, which are open to the outside of the aforesaid casing.

Preferably, the aforesaid casing is made of aluminium, aluminium alloy or steel.

Preferably, the aforesaid casing is an extruded profile or a casing obtained by chilled casting or die casting, however the possibility of making the aforesaid casing by deep-drawing or by bending and welding aluminium or steel sheets, is not excluded.

Preferably, the aforesaid radiological device is capable of dissipating up to 700 W continuously.

Brief description of the figures

Further characteristics and advantages of the invention will become clearer from the following detailed description of some preferred, but not exclusive, embodiments depicted only by way of non-limiting example with the aid of the accompanying drawings, wherein:

- Figure 1 shows schematically the X-ray emission of an X-ray insert versus time, depending on the voltage applied thereto, according to the known art;

- Figure 2 schematically shows the theoretical emission of X-rays of an X-ray insert with respect to time, depending on the voltage applied thereto, according to the known art;

- Figure 3 shows schematically a partial sectional view of a radiological device comprising an X-ray insert equipped with a grid electrode and a high-voltage power supply in accordance with the present invention;

- Figure 4 shows, versus time, the emission of X-rays driven by the aforesaid grid electrode of the radiological device according to the present invention, depending on the voltage applied to the aforesaid X-ray insert; - Figures 5a and 5b schematically show the trend, versus time, of the emission of X-rays driven by the aforesaid grid electrode of the radiological device according to the present invention and the trend versus time, respectively, of the difference of potential between the cathode of the aforesaid X-ray insert and the aforesaid grid electrode, for a given high voltage value applied between cathode and anode of the aforesaid X-ray insert;

- Figure 6 shows a conceptual diagram of a driving circuit of the aforesaid grid electrode according to a first configuration, in accordance with the present invention;

- Figure 7 shows a conceptual diagram of a driving circuit of the aforesaid grid electrode in accordance with an implementation variant of the present invention.

Detailed description of the invention

With reference to figures 3-6, a radiological device for digital applications according to the present invention is generally denoted by 1.

The radiological device 1, hereinafter also identified simply as device, is optimised for operating in pulsed fluoroscopy, i.e. for an operation that substantially provides short emissions of X-rays in rapid succession.

In detail, the radiological device l is a so-called monoblock or power monoblock and essentially comprises a casing 2, an X-ray insert 3 for the emission of X-rays, a high-voltage power supply 4 electrically connected to the X-ray insert 3 for the application of a difference of potential to the X-ray insert 3, a dielectric medium 5 which substantially fills the casing 2 and in which the X-ray insert 3 and the high- voltage power supply 4 are immersed, stopping means for stopping the emission of X- rays generally denoted by 6 and an optoelectronic system 7 for transmitting an activation or deactivation signal to the stopping means 6, input by a user.

In particular, the X-ray insert 3 comprises a filament 8 and a small disc 9 which, during the emission of X-rays, constitute a cathode and an anode, respectively. The filament 8 is in fact connected to the lowest potential of the high-voltage power supply 4, whereas the small disc 9 is connected to the highest potential of the high-voltage power supply 4.

In this regard, the high-voltage power supply 4 comprises a transformer also identified as first transformer, a rectifying circuit or multiplier comprising diodes and capacitors that generates a direct voltage, i.e. a constant difference of potential between its poles between 40 kV and 150 kV, and insulating means arranged between the so- called primary and the so-called secondary of the aforesaid first transformer, in which according to the present invention the aforesaid insulating means, in the examples of the figures not depicted, comprise polypropylene spools.

In accordance with the invention, therefore, the X-ray insert 3 has a maximum working voltage, between cathode and anode, of 150 kV and a maximum power of 50 kW.

As far as the stopping means 6 are concerned, it should be mentioned that they are adapted to interpose, between the filament 8 and the small disc 9, a potential lower than the aforesaid lowest potential corresponding to the potential of the cathode, without changing the aforesaid difference of potential applied to the X-ray insert 3.

In practice, in order to generate the aforesaid lower potential, the present radiological device comprises a dedicated circuit that produces a difference of potential independent of that applied to the X-ray insert 3 by means of the high-voltage power supply 4 in order to obtain the emission of X-rays.

In this regard, the stopping means 6 comprise a stopping element 10 which is adapted to be biased to the aforesaid potential lower than the potential of the filament 8, a transformer or second transformer, a voltage rectifier stage connected to the second transformer and a switch herein also identified as first switch and generally denoted by 11. On the other hand, the second transformer and the voltage rectifier stage connected thereto are generally depicted by 12, while the reference 13 identifies a potential output higher than the rectifier stage and connected to the filament 8, and the reference 14 identifies a potential output lower than the rectifier stage which is connected, through the first switch 11, to the stopping element 10.

When the first switch 11 is open, the device 1 emits X-rays, whereas when the first switch 11 is closed, the stopping element 10 is biased to the aforesaid lower potential, i.e. at a potential lower than the potential of the filament 8 and the device 1 does not emit X-rays. In practice, the stopping element 10 is a so-called grid which, in accordance with the example of the aforesaid figures, is interposed between the filament 8 and the small disc 9 and which essentially comprises a set of parallel metal wires placed on a support.

The aforesaid grid, thus the stopping element 10, substantially represents a third electrode positioned in the X-ray insert 3 which, when is biased to a voltage lower than the cathode voltage, repels the electrons moving toward the anode preventing the emission of X-rays and, therefore, in the scope of present invention, the expressions “grid electrode”, “grid” and “third electrode” are equivalent to each other and each identifies the stopping element 10 in an equivalent way.

Summarising the above, in order to eliminate unwanted radiation emission, e.g. in the case of digital applications operating in pulsed mode (pulsed fluoroscopy), the high-voltage power supply 4 is not operated in on/off mode but, in the present device, an electric field is created in the proximity of the filament 8, which is opposite the electric field produced between the aforesaid anode and cathode, i.e. between the same filament 8 and the small disc 9 of the X-ray insert 3. This way, when desired, the electrons are prevented from accelerating and ending up impinging the small disc 9, thus generating X-radiation by braking.

From an electrical point of view the aforesaid grid and, in any case, the stopping element 10 in any form it may be, represents a small capacitive load and, therefore, its potential can vary rapidly. Therefore, as soon as the grid electrode reaches the value of potential necessary to prevent the electrons from moving, the emission of X-rays ceases and this takes place even if the difference of potential set for the radiological examination remains at the ends of the X-ray insert 3. The emission of X-rays then resumes when the difference of potential between the stopping element 10 and the filament 8 is zeroed.

In order to achieve the same result as above, as an alternative to the aforesaid grid in the form of a set of parallel metal wires placed on a support, the aforesaid third electrode may consist of a metal cup inside which the aforesaid filament is arranged.

The filament of the X-ray insert in fact consists of a small metal spiral through which electric current is passed, which is heated by the Joule effect and which is mounted inside a metal cup defined “filament optics” that conveys the beam of electrons to a single point (or single zone) called focus, on the small disc which is nothing more than a metal disc with a high melting temperature. In this alternative embodiment of the aforesaid third electrode, in the examples of the figures not shown, it is therefore the aforesaid metal cup that is biased to a potential lower than the potential of the cathode, substantially in the same way as the aforesaid stopping element in the form of grid is biased. Of course, in the case where the filament optics are not used as stopping element for stopping X-rays, the aforesaid metal cup is electrically connected directly to the filament so as to be at the same potential as the filament.

In any case, regardless of how the aforesaid stopping element is made, the stopping means 6 and, particularly, the aforesaid voltage rectifier stage preferably generate a constant difference of potential between about -2 kV and about -3 kV, when the first switch 11 is closed.

In other words, the aforesaid third electrode varies its potential between 0 and about -2 kV or about -3 kV with respect to the cathode.

In this regard it should be added that, preferably, the aforesaid cathode, thus the filament 8, is at a potential lower than the earth potential and the aforesaid anode, thus the small disc 9, is at a potential higher than the earth potential.

In addition, preferably, the aforesaid difference of potential with respect to earth is equally divided between the aforesaid anode and the aforesaid cathode, and, therefore, preferably the anode operates at a positive potential between +20 kV and +75 kV and the cathode operates at a negative potential between -20 kV and -75 kV. Consequently, preferably, the aforesaid grid electrode will be brought, with respect to earth, to a value of about -75 kV/-78 kV, although the possibility of bringing the cathode to earth or distributing the difference of potential between cathode and anode differently, for example, is not excluded.

It should be added that the stopping means 6 also comprise insulating means arranged between the so-called primary and the so-called secondary of the aforesaid second transformer, which preferably comprise polypropylene spools arranged between the primary and the secondary of the second transformer. It should be mentioned that, as a whole, between the primary and secondary of the aforesaid second transformer, insulation equal to at least the maximum expected difference of potential between the aforesaid stopping element and the earth is provided.

As regards in detail the opening and closing of the first switch 11, it should be mentioned that, basically, it is controlled by the user through the aforesaid optoelectronic system 7 which makes a non-electrical optical connection for transmitting to the stopping means 6, within the casing 2, an activation or deactivation signal.

The optoelectronic system 7 preferably comprises a transmitter 7a, a receiver 7b and a fibre optic extended between the transmitter 7a and the receiver 7b wherein, preferably, the transmitter 7a is an emitting diode and the receiver 7b is a phototransistor.

Advantageously, the optoelectronic system 7 ensures a sufficient electric insulation with respect to earth, which, as a whole, must be at least equal to the maximum voltage between the stopping element and earth.

In practice, in accordance with the invention, the user intervenes with a control on the optoelectronic system 7 of the desired duration and frequency, which drives the interruption or resumption of the X-rays emission. For example, in accordance with the examples in the aforesaid figures, when the control signal is high, the first switch 11 closes thus connecting the stopping element 10 to the aforesaid lower-potential output of the aforesaid rectifier stage, resulting in the stopping of the emission of X-rays. On the other hand, when the control signal is low, the bias potential is removed from the stopping element 10 (third electrode) and the emission of X-rays takes place.

In accordance with the invention, in order to make the switching of the grid potential particularly quick, e.g. and in accordance with the above from -Vg to 0 V with respect to the cathode, the aforesaid stopping means may also comprise a second switch which connects the stopping means to the cathode and which closes at the moment when the first switch opens and vice versa, thus zeroing the difference of potential between the stopping means and the filament of the X-ray insert when the emission of X-rays is desired. A driving circuit of the grid electrode comprising a second switch 15, as mentioned above, is shown in the example of figure 7, in which parts corresponding to those described with reference to device 1 described above retain the same reference numbers used in the examples of figures 3-6, to the description of which reference is made.

Therefore, in the example of figure 7, are also denoted a filament 8, a small disc

9, an optoelectronic system 7, and stopping means 6 which comprise a stopping element

10, a second transformer connected to a voltage rectifier stage altogether denoted by 12, a first switch 11 and the aforesaid second switch 15, in which the reference 13 identifies a potential output higher of the rectifier stage which is connected to the filament 8, and the reference 16 identifies the connection of the stopping element 10 to the filament 8 through the second switch 15.

In practice, in accordance with the above, the grid driving circuit substantially consists of the stopping means 6 and the optoelectronic system 7.

As mentioned, when the first switch 11 is open, the second switch 15 is closed and in this case the difference of potential between the stopping element 10 and the filament 8 is zeroed, resulting in the emission of X-rays, whereas when the first switch 11 is closed, the second switch 15 is open and, in this case, the stopping element 10 is biased to a potential lower than that of the filament 8, resulting in the stopping of the emission of X-rays.

In accordance with the invention, the first switch 11 and the second switch 15 are actuated by the same control by the user, which control is given through the optoelectronic system 7. In practice, the aforesaid control that manages the two switchings of the aforesaid switches is still the same but the status of the two switches is different because when one is open the other is closed and vice versa. At the circuit level, the function of the aforesaid first switch and, if present, also the function of the aforesaid second switch, is carried out by the high-voltage MOSFET transistors.

In accordance with the invention, it should be added that the present radiological device preferably also comprises means intended for cooling the aforesaid dielectric medium which preferably comprises or consists of a mineral oil, a silicone oil or a synthetic oil.

The aforesaid means intended for cooling the dielectric medium, also identified herein as cooling means, comprise a heat exchanger and, preferably, a heat exchanger selected from the group comprising dual-circuit heat exchangers of the plate or tube type, radiators, mechanically ventilated radiators.

With reference to the example of figure 3, means intended for cooling the dielectric medium in the form of a plate heat exchanger are generally denoted by 17.

In detail, the means intended for cooling the dielectric medium 17 comprise a first circuit 18 for the circulation of the dielectric medium 5 and a second circuit 19 adapted to the circulation of a cooling fluid coming from a cooling system external to the present radiological device, wherein the first circuit 18 and the second circuit 19 are in thermal contact with each other.

As shown in the example of figure 3, the thermal contact between the first circuit 18 and the second circuit 19 takes place within the casing 2 of the present radiological device, however the possibility of providing the aforesaid thermal contact at least partially outside the casing 2 is not excluded. In the latter case, at least the aforesaid first circuit is at least partially extended outside the casing of the radiological device, as may take place, e.g., in the case of radiators and mechanically ventilated radiators.

In any case, the present radiological device, in order to facilitate the circulation of the dielectric medium 5 and, therefore, its cooling by the aforesaid cooling means 17, preferably comprises a pump 20 which is preferably arranged within the casing 2.

In practice, the pump 20 is active on the first circuit 18 in which the dielectric medium 5 circulates, whereas in accordance with the example of the figures, it should be added that the second circuit 19 comprises an inlet opening 21 and an outlet opening 22 for the aforesaid cooling fluid, which are open toward the outside of the casing 2.

This way, the present radiological device is capable of dissipating up to 700 W continuously.

In this regard, it should be added that the heat developed in the X-ray insert during the emission of rays is transferred to the dielectric medium, e.g. to the mineral oil, which is heated to a maximum temperature of about 70 °C, beyond which safety systems intervene that stop the operation of the present device in order to prevent failures.

Therefore, the more stored heat can be dissipated, the slower the monoblock temperature will rise and the greater the number of consecutive exposures or their duration will be.

In this regard it should be mentioned that, advantageously, the casing 2 is preferably an extruded profile or a casing obtained by chilled casting or by die casting, that is made of aluminium or an aluminium alloy which ensures its lightness and good thermal conductivity, however, the possibility of making the aforesaid casing by deep- drawing or by bending and welding aluminium or steel sheets, is not excluded.

The advantages of the present invention, which appeared evident in the description above, can be summarised by noting that a radiological device of the type known as a monoblock has been provided, which is also suitable for heavy-duty applications, in particular power applications with high thermal load, in pulsed fluoroscopy, all using a solution that is compact, with small encumbrance and simplified structure.

In order to meet contingent and specific needs, a person skilled in the art can make numerous changes and modifications to the present invention in the embodiments shown and described, all thereby comprised in the protection scope of the invention as defined by the following claims.