FIELD OF THE INVENTION

The present invention relates to an X-ray tube and a medical device comprising such X-ray tube. Particularly, the present invention is directed to an X-ray tube with a repelling top electrode in an emitter unit for accelerated switching-off of a primary electron beam.

BACKGROUND OF THE INVENTION

In conventional X-ray tubes, a beam of high-energy electrons is directed onto a target in order to generate X-rays. For this purpose, electrons are emitted at an emitter, for example by thermo -ionic emission due to heating an emitter's surface to elevated temperatures. The emitter usually comprises an electron emitting assembly and other electrodes for focusing. Therein, the emitter may form a part of the cathode of the X-ray tube. The electrons are then accelerated in a strong electrical field between the emitter and an anode. Finally the accelerated electrons impact onto a target thereby generating X-rays as "Bremsstrahlung".

When the X-ray tube is switched off, for example after an examination of a patient has been finished or, in case that an X-ray examination is performed using X- ray pulses, after each X-ray pulse, the electron beam should be cut-off as quickly as possible in order to avoid excessive radiation. Conventionally, upon switching-off the X- ray tube, i.e. when an external voltage supplied for generating the voltage between the emitter and the anode is switched off, the electron beam onto the target does not immediately stop but continues for a certain period of for example a few milliseconds, as the capacity of the high voltage circuit which supplies the tube is discharging. During this period, the energy of the electrons impacting onto the target decreases, as the voltage between the emitter and the anode decreases over time.

Accordingly, the X-rays generated have a lower energy as well and are referred to as soft X-rays. On the one hand, these soft X-rays are usually not penetrating a patient to be examined and do therefore not contribute to imaging the patient in a proper way. On the other hand, soft X-rays contribute to the overall dose to the patient as they are mainly absorbed in the patient's skin.

In order to cut-off the electron beam in the X-ray tube as quickly as possible after a desired X-ray pulse to avoid excessive radiation, an additional electrode, also referred to as top electrode, may be placed in front of the electron emitting surface of the emitter. Such top electrode may actively generate a repelling electrical field for the electron beam upon switching-off the X-ray tube by switching off the high voltage supply.

However, it may be difficult in X-ray tubes with high electron emitting capability to actively handle the large cut-off voltages needed.

SUMMARY OF THE INVENTION

Accordingly, there may be a need for an X-ray tube and a medical device comprising such X-ray tube wherein the X-ray tube is adapted for quick cut-off of an electron beam after switching-off the X-ray tube. Particularly, there may be a need for an X-ray tube being capable of generating an electron repelling field upon switching-off using a simple, reliable and cheap electrode arrangement.

These needs may be met by the subject-matter according to the independent claims. Advantageous embodiments of the present invention are described in the dependent claims.

According to a first aspect of the present invention, an X-ray tube comprising a emitter for emitting an electron beam along a beam path in a direction towards a target, a target electrode arranged adjacent to the X-ray beam and a housing enclosing the emitter and the top electrode is proposed. Therein, the top electrode is arranged and adapted such as to be electrically charged by electrons emitted from the emitter via thermo-ionic emission during operation of the emitter. Furthermore, the top electrode is essentially electrically isolated against the emitter.

A gist of the present invention may be seen as being based on the following ideas: The inventor of the present invention has realized that in conventional X- ray tube systems the actively controlled top electrode used for generating a repelling electrical field upon switching-off of the X-ray tube may be difficult to control. One of the reasons is that high voltages in the range of many kilo volts have to be quickly applied to the top electrode after switching-off the voltage supply between the emitter and the anode. Actively controlling such high voltages within a time frame of milliseconds using an external voltage control may be complicated and cost-extensive. Furthermore, the reliability of such actively controlled systems may suffer from voltage surges, generated by uncontrolled discharges between electrodes of the X-ray tube.

It has therefore been an idea of the inventor to create a passive top electrode for generating the repelling electrical field. The passive top electrode should be essentially electrically isolated against the emitter such that the top electrode may be on a different electrical potential than the emitter and/or the housing. On the other hand, the passive top electrode should be arranged and adapted such that it can be electrically charged by electrons emitted from the emitter, to give it e.g. a controlled electrical potential.

Accordingly, the top electrode may be charged to a negative electrical potential being only slightly less negative than the negative potential of the emitter. Therefore, the charged top electrode will not substantially influence the electron beam coming from the emitter during an "on"-state of the X-ray tube. However, upon switching-off the X-ray tube, the electrical potential of the top electrode which has previously been achieved by the charging of the top electrode during the "on"-state will remain at a relatively high level for a certain time period as the top electrode is essentially isolated against the emitter and may build up a capacitance to other electrodes of the tube, which may be on fixed potential, e.g. the tube frame. On the other hand, upon switching-off the X-ray tube, the electrical potential of the emitter in the cathode will decrease faster. For a certain transition period after switching-off the X-ray tube, the negative potential of the top electrode may be higher than the negative potential of the emitter and thereby generating a repelling field which limits or even cuts-off the electron emission at the emitter thereby limiting the generation of the X-rays.

As the proposed X-ray tube neither needs an external voltage control for actively controlling the electrical potential of the top electrode nor needs any electrical connection for externally applying a voltage to the top electrode, it may be at the same time easy to fabricate, inexpensive and reliable. In the following, possible features and advantages of embodiments of the proposed X-ray tube will be described.

The emitter may comprise heatable electron emitting assembly. It may comprise two electrical connections which can be used, on the one hand, for supplying an electrical current for producing heat in an electrical resistance of the emitter within the emitter and which, on the other hand, can be used for applying an appropriate negative potential to the emitter with respect to the anode of the X-ray tube. The emitter has an electron emitting surface directed towards a target within the X-ray tube. Electrons may be emitted from the heated emitting surface via thermo-ionic electron emission and may then be accelerated towards the target due to the potential difference between the emitter and the anode.

The top electrode shall be arranged adjacent to a path of the electrons accelerated towards the target. For example, the top electrode may be a ring shaped electrode surrounding the electron beam path. The top electrode may be arranged at a short distance of for example a few millimeters or less adjacent to the electron emitting surface of the emitter.

The housing which may also be referred to as frame envelope encloses the emitter and the top electrode and may also enclose the anode and the target. The housing may be on a same electrical potential as the anode. Furthermore, the target may also be on the same electrical potential as the anode. For example, the anode may be grounded such that the potential of the anode as well as optionally the potential of the housing and the target may be set at 0 V.

The top electrode may be arranged at a position within the X-ray tube such that electrons emitted from the emitter via thermo-ionic emission during operation of the emitter, i.e. as long as the emitter is heated, may reach a surface of the top electrode or a surface of an electron capturing part electrically connected to the top electrode. Thereby, the electrons arriving at the surface of the top electrode or the electron capturing part may charge the top electrode to a negative potential which is only slightly more positive, than the negative electrical potential of the emitter during operation of the X-ray tube. As will be described in further detail below with respect to a specific embodiment, the top electrode or the electron capturing part connected therewith may be arranged at a very short distance of a few millimetres or less to the electron emitting surface of the emitter.

While the top electrode may be charged via free electrons coming from the heated emitter, it shall be electrically isolated against the emitter. This may mean that there is no direct electrical connection for example in the form of an electrical conductor between the top electrode and the emitter. Instead, there may be a gap between the top electrode and the emitter. While this gap may be overcome by free electrons emitted at the emitter' surface, the gap may act as an electrical isolator for bound electrons such that essentially no electrons may flow between the emitter and the top electrode as long as the emitter is not heated or the potential at the top electrode is more negative than that of the emitter. The combination of top electrode and emitter act as a diode. In case there is a mechanical connection between the top electrode and the emitter, this mechanical connection should be realized using an electrically isolating material, like alumina ceramics. Furthermore, the top electrode may be essentially electrically isolated against the housing of the X-ray tube. While the top electrode may be mechanically supported by the housing, the mechanical connection between the housing and the top electrode should be performed by an electrically isolating material interposed between the housing and the top electrode. To maintain the negative voltage level of the top electrode despite of leakage currents through the supporting insulator and stray capacitance, the capacity between top electrode and tube frame should be maximized. This can be done by a proper geometry, e.g. by enlarging the surface area and / or by minimizing the distance or using insulating material with a high dielectric constant. According to an advantageous embodiment of the present invention, the top electrode is adapted to generate a repelling electrical field for repelling electrons emitted from the emitter after actively switching-off the high voltage supply to the emitter.

This capability of generating a repelling electrical field may arise from the specific arrangement of the top electrode and its electrical isolation against further components of the X-ray tube. While the top electrode may be electrically charged during an "on" state of the emitter, i.e. while the emitter is heated and supplied with high voltage, the built-up charge may remain on the top electrode also for a specific period after switching-off the voltage supply to the emitter, due to the essential electrical isolation of the top electrode against the emitter. Accordingly, the electrical potential of the top electrode may remain high within this time period while at the same time, the electrical potential of the emitter will decrease. Thereby, the repelling electrical field may be generated. This repelling field may hinder or prevent electrons which are still emitted from the emitting surface of the emitter from reaching the target. Thereby, the switching- off of the X-ray emission from the target upon switching-off of the emitter may be substantially accelerated, particularly during the initial phase after switch-off.

According to a further advantageous embodiment of the present invention, the top electrode is not wired for applying an external voltage to the top electrode. In other words, there are no electrical connections for example in the form of cables for applying an external voltage to the top electrode. While in prior art approaches, top electrodes are controlled by actively applying a voltage from an external control device to electrical connections of a top electrode, no such electrical connections are necessary in the X-ray tube according to this embodiment of the present invention. Instead, the electrical potential of the top electrode is not controlled actively by applying an external voltage but is set passively due to the fact that the top electrode is electrically charged by electrons emitted from the emitter during operation of the emitter. Such passive potential control may render the X-ray tube and its switching-off operation more reliable than in previous X-ray tube systems. According to a further advantageous embodiment of the present invention, an isolator having a resistance R _tc is arranged between the top electrode and the emitter and the emitter and the top electrode are capacitively coupled with a capacitance C _to , and the housing and the top electrode are capacitively coupled with a capacitance Qh ; wherein the top electrode and the isolator are arranged and adapted such that C _tc is smaller than Qh.

In other words, in the X-ray tube according to this embodiment, the top electrode is electrically isolated against both, the emitter and the housing using one or more isolators. Such arrangement with an isolator interposed between two conducting components, namely a top electrode on the one side and an emitter or a housing on the other side, may effect that the top electrode and the emitter are capacitively coupled to each other with a first specific capacitance Q _c and that the top electrode is coupled to the housing with a second specific capacitance Qh . When the top electrode and the isolator are arranged and adapted such that Q _c is smaller than Qh , the top electrode is more capacitively coupled to the housing than to the emitter. Accordingly, a change in the electrical potential of the emitter, e.g. due to switching of the voltage supply to the emitter, does hardly effect the electrical potential of the top electrode as the top electrode is mainly capacitively coupled to the housing and not to the emitter.

Furthermore, the isolator between the top electrode and the emitter prevents ohmic coupling, i.e. hinders electrical charges from flowing between the top electrode and the emitter.

According to a further advantageous embodiment of the present invention, a time constant R _tc • Q _c .is larger than a time constant with which a potential of the emitter decreases after switching-off a voltage supply connected to the emitter.

In other words, on the one hand, the top electrode may be charged to an electrical potential with respect to the emitter, and, on the other hand, electrical charges may flow from the top electrode to the emitter as the isolator there between is not ideal but has a specific resistance. Accordingly, the top-electrode-isolator-emitter arrangement may have its own time constant R _tc • C _tc ..

The idea may be now that this time constant R _tc • C _tc . is larger than a time constant with which a potential of the emitter decreases after switching-off a voltage supply connected to the emitter. In other words, the emitter may be connected to an internal or external voltage supply which, during operation of the X-ray tube, may set the emitter on a specific negative potential. Upon switching-off the voltage supply, the negative potential of the emitter will drop. Therein, the time constant, corresponding to the velocity with which the potential drops, may depend on several factors such as the capacitive coupling and the electrical isolation between the emitter and its environment, the wiring to the voltage supply, etc. If the time constant R _tc • C _tc of the top-electrode- isolator-emitter arrangement is large compared the time constant of the voltage drop at the emitter upon switching-off, the top electrode may stay on a more negative potential than the emitter during switching-off thereby enabling a temporary electron repelling electrical field.

In other words, such arrangement may result in the effect that, upon switching-off the emitter, the electrical potential of the emitter decays faster than the electrical potential of the top electrode. Accordingly, there may be a time period immediately after switching-off the emitter in which time period the electrical potential of the emitter is less negative than the electrical potential of the top electrode whereby an electron repelling electrical field between the emitter and the top electrode may be established.

According to a further advantageous embodiment of the present invention, the resistance R _tc of the isolator is larger than 10 Megaohm, preferably larger than 50 Megaohm, more preferably in an order of magnitude of 100 Megaohm. Furthermore, the capacitances C _tc , Qh are in a range between 1 pF and 100 pF, preferably in a range between 3 pF and 20 pF. Such values for the resistance and the capacitances have been proven in respect of the inventive X-ray tubes to enable an effective repelling electrical field for accelerating the switching-off process of the X-ray tube.

According to a further advantageous embodiment of the present invention, the top electrode is electrically connected to an electron capturing part for capturing electrons emitted from the emitter via thermo -ionic emission during an operation of the emitter. The electron capturing part is arranged at a distance to the emitter, i.e. with a gap, of less than 5 mm, preferably less than 3 mm and more preferably a distance of between 0.2 and 2 mm.

The electron capturing part may be made with an electrically conductive material. It should be arranged in close neighbourhood to the emitter. For example, it may be arranged parallel to a heatable surface of the emitter from which, during operation, electrons are thermally emitted. However, the electron capturing part should be arranged such that it does not negatively influence the emission and acceleration of the electrons of the primary beam of the X-ray tube towards the target, i.e. it should not "stand in the way" of the electrons of the primary beam. For example, the electron capturing part may be arranged at a side of the emitter opposite to the side on which the top electrode is arranged. In other words, while the top electrode is arranged adjacent to the beam path of the primary electron beam, the electron capturing part can be arranged close to a surface of the emitter opposite to the side of the target. Accordingly, the electron capturing part can be arranged close to the surface of the emitter

As the electron capturing part is electrically connected to the top electrode, charges collected by the electron capturing part are at least partially transferred to the top electrode effecting a built-up of an electrical charge within the top electrode. The top electrode may act as an anode of a diode, the emitter as a cathode of it. According to a further aspect of the present invention, a medical device comprising an X-ray tube according to embodiments of the present invention as described above is proposed. The medical device can be any X-ray system such as for example a computer tomography (CT) system or a C-arm system and can be used for imaging and examining a patient by transmitting and detecting X-rays therethrough.

Expressed in other words, features of the invention and its embodiments may be summarized as follows: An electron beam from an emitter within an X-ray tube may need to be cut-off as quickly as possible after a desired X-ray pulse to avoid excessive radiation. An electrode may be placed in front of the emitting surface of the cathode which electrode may generate a repelling electrical field for electron beam cutoff. The repelling field is only generated after switching-off ("off ') of the negative high voltage which is otherwise actively supplied to the emitter. To achieve this, the top electrode of the cathode (e.g. "cathode head") may have to be isolated from the high voltage (h/v) connection to the h/v generator (centre wire of the h/v cable) and the connected emitting surface. The top electrode may be substantially electrically floating and coupled to ground only through its geometric capacity C (for example approximately 10 pF). C should be larger than the stray capacity between the arrangement of emitter and centre wire of the high voltage cable on one end and the top electrode on the other end. The insulation may have a very low conductivity (resistivity R of for example some hundred Megaohms) such that the time constant R • C is at least in the range of the length of the high voltage tail after switch-off (e.g. some milliseconds). During beam generation ("on"), part of the emission current from the emitter is directed towards the top electrode by placing a connection wire closely (for example approximately 0.3 - 2 mm) above to the surface of the emitter such that it is acting as the anode of a diode during "on" and the voltage of the top electrode is only a little more positive than that of the emitter. In addition, this small gap may act as a voltage limiter and protect the insulation in case of h/v discharge of the top electrode to the ground (tube arching). Immediately after switch off, the top electrode first stays negative (due to its capacity to ground) while the emitter and the cable are already discharging. For a certain transition period, a repelling field is generated which limits the electron emission or even cuts it off and so limits the production of X-rays. The proposed system may be beneficial as it is inexpensive, that no external connection is needed and that no separate grid supply from the high voltage supply unit is needed.

It has to be noted that features and advantages of the present invention have been described with reference to different embodiments of the invention and, partly, also with respect to manufacturing properties of the inventive device. However, a person skilled in the art will gather from the above and the following description that, unless other notified, in addition to any combination of features belonging to one embodiment also any combinations between features relating to different embodiments or to a manufacturing method is considered to be disclosed with this application.

BRIEF DESCRIPTION OF THE DRAWINGS

Features and advantages of the present invention will be further described with respect to specific embodiments as shown in the accompanying figures but to which the invention shall not be limited.

Fig. 1 shows a cross section through an X-ray tube according to an embodiment of the present invention. Fig. 2 shows a specific cathode arrangement for an X-ray tube according to an embodiment of the present invention. Fig. 3 shows a graph indicating the time-depending negative potential/emission current for a conventional X-ray tube and for an X-ray tube according to an embodiment of the present invention.

Fig. 4 shows a medical device with an X-ray tube according to an embodiment of the present invention.

The drawings in the figures are only schematically and not to scale. Similar elements in the figures are referred to with similar reference signs. DETAILED DESCRIPTION OF PREFERRED EMBODIMENTS

Fig. 1 shows an embodiment of an X-ray tube 1 according to the present invention. The X-ray tube 1 comprises an emitter 3 for emitting an electron beam 5 along a beam path in a direction towards a target 7. The emitter 3 is connected via cables 9 to two external voltage connections 11, 13. To these connections 11, 13, a negative potential of e.g. -125 kV can be applied. Furthermore, these connections 11, 13 can be used to induce a heating current into the emitter 3 thereby heating an electron emitting surface of the emitter 3 to desired elevated temperatures of more than 2000 ⁰ C. Adjacent to the beam path of the electron beam 5, a top electrode 15 is arranged. The top electrode comprises a cylindrical electrode which is arranged such that its longitudinal axis coincides with the electron beam path.

The emitter 3 together with the top electrode 15 and the target 7 are enclosed by a housing 17. The housing 17 and the target 7 are both grounded thereby setting its electrical potential to 0 V.

Upon impact of the electron beam 5 onto a focal spot on the target 7, an X-ray beam 19 is generated which exits the housing 17 through a specific window 21.

As only schematically shown in Fig. 1 and as also indicated by the circuit diagram included on the upper left side in Fig. 1, the emitter 3 and the top electrode 15 are geometrically arranged with respect to each other and with respect to the housing 17 and isolators 23, 25 are provided in such a way that specific capacitive coupling and resistive coupling conditions are met. The emitter 3 is mechanically attached to the housing 7 via a large high voltage insulator 25. Furthermore, the top electrode 15 is isolated against the emitter 3 using an additional insulator 23. The additional insulator 23 may be mechanically attached to a supporting structure 29. Furthermore, the top electrode 15 may be insulated against the housing 17 with a vacuum gap 27 and a combination of the insulators 23 and 25 which may be established as solid-state insulators.

Accordingly, as indicated in the circuit diagram of Fig. 1, the additional isolator 23 having a resistivity R _tc and being interposed between the top electrode 15 and the emitter 3, taking into account its geometric relation to the top electrode, creates a capacity C _tc which may also be referred to as the stray-capacity of the emitter. This capacity C _ta may be for example 5 pF.

Furthermore, the geometrically optimized arrangement of the top electrode 15 with respect to the housing 17 creates a capacity C _t h being for example in the range of 15 pF.

While the top electrode 15 is mechanically attached to the support structure 29 enclosing the emitter 3, the top electrode 15 is electrically essentially isolated against the emitter 3 via the isolator 23. Accordingly, in a non-heated state of the emitter 3, no charges will flow from the emitter 3 to the top electrode 15 or vice versa.

However, an electron capturing part 31 is connected to the top electrode

15. This electron capturing part 31 extends into a region on the rear side of the emitter

3. A small gap of 0.3 - 2 mm separates the electron capturing part 31 from the emitter 3. Furthermore, the electron capturing part 31 is electrically isolated against the emitter 3 such that, in a cold state of the emitter 3, no charges flow from the emitter 3 to the electron capturing part 31 or vice versa.

However, in a hot state of the emitter 3, i.e. while operating the X-ray tube 1 , electrons are also emitted from the electron emitting surface in a direction towards the electron capturing part 31. These electrons can be captured by the electron capturing part 31 thereby electrically charging both the electron capturing part 31 and the top electrode 15 connected therewith.

Accordingly, the close arrangement of the emitter 3 and the electron capturing part 31 may be seen as creating a thermo-ionic diode, i.e. an electron receiving "anode" e.g. on the backside of the emitter acting as a potential control of the top electrode during an "on" state of the X-ray tube 1. The distance of the emitter to the

"diode electrode" may depend on a desired repelling potential. For example, for a potential difference of -10 kV, a distance of approximately 1 mm is optimal. Such gap would have to hold 10 kV for some milliseconds. This means, upon switching-off the emitter 3, the top electrode 15 would remain charged at a relatively high negative potential for some milliseconds while the negative potential of the emitter 3 already steeply decreases. Thereby, a sufficient repelling electrical field 33 might be created which might reduce the emission of the X- ray tube for example by more than 80 % within a certain time period. Thereby, the overall radiation dose to a patient can be significantly reduced.

Fig. 2 schematically shows a emitter head for an X-ray tube according to an embodiment of the present invention. An emitter 3 and a top electrode 15 are enclosed in a housing 17. The top electrode 15 is provided as a ring electrode surrounding the path of the electron beam 5 in the neighbourhood of the emitter 3. A large high voltage isolator 25 is provided for isolation of the emitter 3 against the housing 17. At the same time, insulation is provided between the top electrode 15 and the emitter 3. A distance d between parts of the top electrode 15 and the housing 17 is minimized in order to maximize a capacity to the grounded housing while at the same time providing sufficient insulation against both, the housing and the emitter 3. An electron capturing part (not shown in Fig. 2) is placed at a distance of about 1 mm behind the emitter 3 and is connected to the top electrode 15.

In Fig. 3, exemplary conditions with respect to a negative potential of the emitter and the top electrode and with respect to an emission current in dependence of a time during "on"-state and after switching-off are schematically shown. After switching "on", both, the negative potential 41 of the emitter and the negative potential 43 of the top electrode increase up to a specific maximum. Therein, during the "on"-state, the negative potential 43 of the top electrode is only slightly more positive than the negative potential 41 of the emitter as, during the "on"- state, the top electrode 15 is increasingly charged by electrons coming from the emitter 3 and being captured by the electron capturing part 31 connected to the top electrode 15. At the same time, the emission current 47 within the X-ray tube increases up to a maximum value during the "on"-state.

After switching-off the high voltage supply to the emitter, both, the negative potential 41 of the emitter as well as the negative potential 43 of the top electrode will start to decrease. However, the negative potential 43 of the top electrode will decrease more slowly than the negative potential 41 of the emitter due to the fact that the top electrode arrangement has a larger time constant R _t h • Qh than the time constant of the voltage drop at the emitter arrangement, wherein the emitter and the high voltage cable are discharged by the primary electron beam

Accordingly, after switching-off the emitter, a repelling electrical field 45 will be established between the top electrode and the emitter. This repelling electrical field 45 will help to quickly cut off the emission current 47 of the X-ray tube.

For comparing purposes, the emission current 49 for the case of a conventional X-ray tube in which no repelling electrical field is established after switching-off the emitter is also indicated in Fig. 3. It can be seen that in such conventional X-ray tube, the emission current will decrease much more slowly than in the X-ray tube in accordance with the present invention and forms a tail which corresponds to the production of undesired excessive emission of soft X-rays. In Fig. 4, an example of a medical device 200 with an X-ray tube according to an embodiment of the present invention is shown. Fig. 4 shows the main features of a CT scanner, namely an X-ray source 220 comprising an X-ray tube 1 , a radiation detector 210 and a patient couch 230. The CT scanner may rotate around the object to be observed and may acquire projection images by means of radiation detection using the detector 210.

Finally, it should be noted that the terms "comprising", "including", etc. do not exclude other elements or steps and the terms "a" or "an" do not exclude a plurality of elements. Also, elements described in association with different embodiments may be combined. It should also be noted that reference signs in the claims should not be construed as limiting the scope of the claims.