WO2013131628A1

WO2013131628A1 - Compact x-ray sources for moderate loading with x-ray tube with carbon nanotube cathode

Info

Publication number: WO2013131628A1
Application number: PCT/EP2013/000603
Authority: WO
Inventors: Roberto Molteni
Original assignee: Roberto Molteni
Priority date: 2012-03-05
Filing date: 2013-03-01
Publication date: 2013-09-12
Also published as: ITVR20120035A1; EP2823502A1

Abstract

New kinds of moderate-load X-ray source devices are described, with reduced weight and size, thus being especially suitable for compact portable X-ray sources, such as those required for hand-held dental and veterinarian applications. The devices comprises an X-ray tube (1) with carbon nanotube cathode (3) and electrically grounded anode (5), whereas the overall design of the device and of the X- ray tube therein are adapted so to take advantage of the special characteristics of the carbon nanotube cathode, instead of replicating the design concepts that have been used for long time with filament- based X-ray tubes and with high-load X-ray sources.

Description

COMPACT X-RAY SOURCES FOR MODERATE LOADING WITH X-RAY

TUBE WITH CARBON NANOTUBE CATHODE.

*****

FIELD OF THE INVENTION

The invention concerns X-ray devices and methods of using thereof. More particularly, the invention concerns X-ray tubes used in X-ray sources especially suitable for portable X-ray devices, containing special X-ray tubes with carbon nanotube cathode, and to design methods apt to integrate such special X-ray tubes and take optimal advantage of their characteristics.

BACKGROUND

X-ray sources for high-load radiography are customarily composed of the following: an X-ray tube, often of rotating anode type; a housing or sealed enclosure to enclose the above-mentioned X-ray tube immersed in an insulating and heat-conductive media (usually mineral oil), such housing accommodating high-voltage connector sockets; shielding to block the X-ray except at an exit window; a X-ray collimator; additional filter(s) to determine the total X-ray filtration; high-voltage cables; and a high-voltage generator controlled by a control system that, among else, regulates the radiological technique factors (aJ .a. loading factors) to the X-ray tube (kV, mA, and s).

The majority of heavy-load medical radiography use this kind of design, which benefit from easily replacement of X-ray tube when exhausted that may happen several times over the apparatus lifetime due to the intense workload in such applications.

However, many radiographic applications make use of moderate electrical and radiological loads, such that the X-ray tube is subject to moderate stress and may well last over the whole life of the radiographic system without need of replacement. In these cases the X-ray tube is generally of fixed anode type (i.e. not rotating), and the X-ray source is based upon a so-called tube-head (a.k.a. mono-block) design. With such design, a sealed housing contains in its interior the fixed-anode X-ray tube and a step-up high-voltage transformer and other electrical circuitry needed to generate the high voltage required for X-ray tube to operate. As in the case previously described, the parts inside the tube-head are immersed in an insulating and heat-conductive media. The advantages of the tube-head design respect to that with external high-voltage power source and re-loadable X-ray tube include, among else, avoidance of bulky high- voltage cables and connections. These parts must insulate at their interior the extremely high voltages required for medical radiology, which may reach or even exceed 150 kV, whereas their external surface shall be safe to touch.

A classic X-ray tube requires a supply of high voltage (at relatively moderate current) across the anode and cathode - this is called the anodic circuit - and a circuit to supply the cathode, which consist of a filament (typically made out of tungsten) with a low voltage but relatively high current sufficient to heat-up the filament and cause it to emit electrons by thermionic effect, i.e. the filament circuit or cathodic circuit. X-ray tubes may also include a third electrode, called wehnelt, kept at a negative electrical potential respect to the cathode in order to focus and stabilize the electron flow.

X-ray sources, either of re-loadable X-ray tube or tube-head types, are further subdivided into two categories: AC (Alternate Current) and DC (Direct Current).

An AC tube-head X-ray source is supplied directly from the electrical mains power, i.e. from an alternate-current source at low frequency. Its simplest embodiment consists of the X-ray tube, a step-up transformer with a very high transformer ratio, the necessary electrical connections, and a sealed housing that contains the insulating media and provides the X-ray shielding in all direction but at an exit window. The X-ray tube self-rectifies the applied AC high-voltage, hence the electron flow and X-ray emission occurs in peaks only during the positive half-wave. The high-voltage to the anodic circuit is often supplied in bipolar mode, i.e. the difference of potential between anode and cathode is obtained by supplying the anode with a voltage respect to ground corresponding to approximately half of the total difference of potential, and the cathode with a voltage respect to ground corresponding to the remaining approximate half of the total difference of potential, thus limiting the maximum electrical potential respect to ground inside the tube-head. Supply of the cathodic filament circuit is often obtained simply by adding further windings to the secondary coil beyond the tap for the cathode; the end of such additional winding connects to the opposite side of the cathode filament. In this way a single transformer with one primary coil sustains both the main secondary winding for the anodic circuit and a smaller secondary winding for the cathodic circuit. AC type tube-heads are traditional, old-style design in use since many decades. In recent years this design has gradually been supplemented, or supplanted, by the more modern DC type design.

A DC tube-head X-ray source is usually supplied from a power electronic converter (or an inverter if the input electrical power is a direct current). The converter is typically powered from the electrical mains and produces a stabilized square-wave output at medium frequency, generally in the range between a few thousand hertz to a few hundred thousand hertz. Inside the tube-head there is a stepping-up transformer with a ferrite magnetic circuit. Such ferrite magnetic circuit can be utilized due to the relatively high frequency used, and makes possible a compact and lightweight construction of the transformer. It is not necessary that the transformer elevates the voltage all the way up to the value required to operate the X-ray tube because at such relatively-high frequencies it is practically possible to interpose a voltage multiplier or booster i.e. an array of high- voltage diodes and capacitors (e.g. of Cockroft- Walton type), which elevates many-folds the voltage output from the secondary of the transformer up to the value required across the X-ray tube. Furthermore the voltage at the output of voltage multiplier is fully rectified, i.e. is a true direct voltage (DC), which results into harder X-radiation of more consistent quality.

In DC-type X-ray sources, usually the power to filament cathode is supplied and controlled independently from the anodic voltage, in pursue of an accurate regulation of the anodic current, also to preheat the filament cathode so that the onset of anodic current occurs immediately at the application of anodic voltage and very short irradiation times can be accurately controlled. Usually one side of the filament cathode is directly connected to the electrical ground, and the potential difference anode-to-cathode is obtained with a positive voltage respect to ground at the anode. This configuration requires isolation at the anode side twice higher than in case of the bipolar configuration previously described, but offers the advantage that the cathode filament can be driven directly from low-voltage circuits with no need of special insulation, since one side of it is directly connected to ground.

Compared to older AC design, the DC design features better control and stabilization of radiographic technique factors (kV, mA, and s) independently from tolerances and fluctuations of the mains power supply and of X-ray tube aging, accurate setting of very short irradiation times, and lighter construction.

These advantages are exploited in portable X-ray sources where compactness and light weight are paramount. For instance, such design is used in hand-held X-ray sources, especially suited for dental radiography, including in also rechargeable batteries, inverter circuitry, and electronic control circuits. In this set, high-voltage electrical insulation can be advantageously achieved by means of insulating gel or putty incorporating compounds of high atomic weight (such as e.g. barium oxide) to further provide X-ray shielding with no need of separate specific radiation shield. Such kind of design is described for instance in US patents 2000 6038287 (Miles) and US patent 2009 7496178 (Turner).

Starting from the 90's, new types of X-ray tubes has been described, prototyped, and industrially produced, that is X-ray tubes where cathodic electron emission is not obtained by thermionic effect with a heated filament, but by room-temperature electric field emission from materials having suitably-low work function, typically carbon nanotubes. The flow of electrons is usually enhanced and controlled via a takeoff electrode (or "grid") interposed between anode and cathode, and very close to the cathode, whose electrical potential determines the electric field at the cathode surface that causes the extraction of electrons.

This technology offers advantages and disadvantages. The advantages include:

· No circuit is necessary to glow the filament. Therefore, also no electrical power is required for this purpose.

• Almost-instantaneous emission of electrons, and therefore production of X-ray, at the moment of application of the high voltage to the anodic circuit, which makes possible the emission of X-ray pulses at rather high frequency.

Disadvantages include:

• Stability in long-terms performances, and repeatability among different exemplars of supposedly identical X-ray tubes, may be questionable, being determined by nano- scale microscopic features of carbon nanotube cathode, which are deemed difficult to control in a repeatable manner with the current process technology know-how.

· Control of the X-ray tube anodic current may be critical. • A significant portion of the electrons extracted from the cathode is captured by the grid and lost for the X-rays production.

X-ray tubes with carbon nanotube cathode are commercially available. However their design - and the overall design of the X-ray system incorporating them - essentially replicates that of conventional filament-based X-ray tubes and systems except, of course, for the cathode being made of a carbon nanotube strip instead of a heated filament. Instead, the overall system design should be reshaped in view of the distinctive advantages underlying the properties of carbon nanotube cathodes.

For example, the anode of the x-ray tube can be connected to the electrical ground and all the electrical potential applied to the cathode, as a high voltage negative potential, without need of the cathode-side circuitry for driving the filament, which would requiring special provision for high-voltage insulation and isolation. Connecting the anode to ground is disclosed for instance in US Patent 2002 6456691 (Takahashi et al.) and in US Patent 2003 0002627 Al (Espinosa et al.), however these documents do not address the ways by which such feature can be advantageously exploited.

As previously stated, by the present technology the control of current on anode-cathode circuit, i.e. the electrons flow in the x-ray tube and therefore the rate of x-ray generated at target, is difficultly afforded in accurate and reliable way, being influenced by the sub- microscopic-scale properties of carbon nanotubes.

For instance, US Patent 2002 6456691 (Takahashi et al.) discloses a design with an x-ray source from carbon nanotube cathode, whose main scope is the regulation of the electrical current anode-to-cathode by variously and independently controlling the voltage potential respect to cathode of a so-called a takeoff electrode (or"grid") and of the wehnelt, and/or by controlling the temperature of the carbon nanotube cathode.

Japanese Patent 4566501 B (2010; Matsumoto et al.) discloses a method to "control the operation of x-ray tube" having a carbon nanotube cathode by means of an electronic switch in series with the cathode, between the cathode itself and ground, based upon a feedback signal from a radiation detector. However, that method has several limitations:

• The type of x-ray tube described, with a conical anode made out of copper having the tip as the electrons target and X-ray focal source, is not suitable for use in normal diagnostic medical radiology, where a stripe focal source on a tungsten target disposed at an angle with the x-ray beam axis is needed to produce enough x-ray flux. This type of design is suitable for applications other-than medical radiology, e.g. for industrial testing or material analysis only.

• Controlling the x-ray production by a switching component along the anode-cathode circuit requires a switch capable to withstand very high commuting voltages, and can only be conceived for relatively low high-voltage ranges. In facts, the inventive scope has been limited to high voltages up to 20 kV, which is largely insufficient for any medial radiological application (except mammography).

So the teachings of Japanese patent 4566501 can only be applied to x-ray devices for special applications, not generally including medical or dental radiology, and require anyway the use of a critical and expensive component (the high-voltage electronic switch).

In conclusion, the use of carbon nanotube cathodes in X-ray tubes is already well known in the art, but the teaching from prior literature and patent documents does not include methods to reduce the overall size and weight of the X-ray source by taking full advantage of the specific features and properties of the carbon nanotube cathode with respect to the conventional thermionic filament cathode. This document addresses such shortcoming, with special attention to the requirements for X-ray sources for medical and dental diagnostic imaging.

BRIEF DESCRIPTION OF THE DRAWINGS

Figure 1 is a schematic view of the X-ray device wherein hooded-type anode of the tube is mechanically, electrically, and thermally intimately connected to a part of the housing having large thermal capacity, and a takeoff electrode (or "grid") is present between the carbon nanotube cathode and the anode, in close proximity with the former.

Figure 2 is a schematic view of the X-ray device wherein no takeoff electrode is present in proximity of the carbon nanotube cathode and the extraction of electrodes from the cathode is obtained directly by the anode-to-cathode electrical field.

Figure 3 is a schematic view of the X-ray device wherein a radiation detector, situated at a position subject to irradiation from the X-radiation, provides a feedback signal to a control circuit which interrupts the energizing of the X-ray tube, and consequently the emission of X-radiation, when a preset amount of radiation dose is reached. Figure 4 is a schematic view of the X-ray device wherein the anode of the x-ray tube is shaped in the customary way for fixed anodes, i.e. is not "hooded", and is not bolted to the inside wall of the housing.

SUMMARY OF THE INVENTION

The present invention features compact and lightweight X-ray sources for moderate radiological load, especially suitable for portable X-ray systems, utilizing a specially- designed X-ray tube with carbon nanotube cathode, whereas the overall design of the X- ray source is adapted to take advantage of the fact that no electrical power and no active circuitry is required to cause the cathode to emit electrons.

One aspect of the invention includes having the anode connected to ground and preferably shaped so to conform to a suitable metallic part of the housing (typically one of its internal surfaces) and fasten to it, thus achieving an intimate physical contact that maximizes the heat flow from the focal spot in the target and its dissipation. This is especially advantageous when the electric insulation inside the tube-head is augmented with a medium other than mineral oil, thus missing the additional heat dissipation that oil convection provides. The anode is preferably of hooded type, that is the target which is hit by the accelerated electrons is recessed inside a deep hollowing of the anode body; in this way the majority of the stray X-ray - that is, the portion of X-radiation that do not belong to the primary beam exiting from the exit window and useful for radiographic imaging - is absorbed by the anode itself, thus reducing the needs for specific radiation shielding, and consequently the size and weight of the tube-head.

In one embodiment, the present invention features a radiation detector that detects the air kerma (the radiation "dose"), or more commonly its rate, and feeds a signal to a control circuit which interrupts the supply of power to the x-ray tube - therefore X-ray emission - when a certain threshold is reached as set by an operator, thus relieving the needs to accurately regulate anodic current which may be particularly difficult with X-ray tubes with carbon nanotube cathode, especially when a takeoff electrode (or "grid") is not used. DETAILED DESCRIPTION

The terms and expressions used herein match the terminology officially adopted by the International Electrotechnical Commission and defined in IEC publication TR 60788 - 2nd Ed. 2004-02 Medical Electrical Equipment - Glossary of Defined Terms - and in other IEC standards, and is preferably used when it exists and unless otherwise specified.

In X-ray tube technology the term "cathode" generally means simultaneously an electrode with more negative voltage potential during the x-ray generation phase and the source of electrons accelerated by electrical field and causing emission of x-radiation by hitting a target. The term "anode" generally encompasses the electrode with more positive voltage potential during x-ray generation, which incorporates a target that becomes the source of X-rays when hit by electrons accelerated by the electrical field.

The present invention consists of a x-ray source of the tube-head type (alias mono- block), comprising, among else, (i) an enclosure, (ii) an x-ray tube, (iii) a converter or inverter circuit to convert input electrical power into a pulsating current at relatively high frequency and voltage, (iv) a "voltage multiplier" electrical circuit to boost the voltage from the converter or inverter up to the high voltage required for the operation of the X-ray tube, (v) means to insulate the high voltage parts, (vi) means to shield all unwanted X- radiation except at the output window for the primary-ray beam, (vii) optionally radiographic filters for the primary X-ray beam, (viii) an electronic control circuit (that may or may not be inside the above-mentioned enclosure) to control the irradiation events under the supervision of the operator, and (ix) other components needed and customarily included in tube-head-type x-ray sources.

The anode of the X-ray tube, incorporating a target that emits X-rays when hit by the electrons of the cathode-to-anode current, is of fixed type, i.e. not rotating, and is connected to ground and preferably shaped so to conform to a suitable relatively-large metallic part of the housing (typically one of its internal surfaces) and fasten to it, thus achieving an intimate physical and thermal contact that maximizes the heat flow from the focal spot in the target and its dissipation. The anode is preferably of hooded type, i.e. the target which is hit by the accelerated electrons is recessed inside a deep hollowing of the anode body; in this way the majority of the stray X-ray - that is, the portion of X-radiation that does not belong to the primary beam exiting from the exit window and is not useful for radiographic imaging - is absorbed by the anode itself, thus reducing the needs for specific radiation shielding, and consequently the size and weight of the tube-head. The electrical insulation inside the tube-head is augmented by pouring the volume surrounding the high voltage parts (body of the x-ray tube, cathode and associated electrical circuitry) with a medium other than mineral oil, for instance gel or putty or elastomeric rubber, since the additional heat dissipation that oil convection provides is here unnecessary due to the intimated thermal contact of the anode with a mass having large thermal capacity. Such insulating media is charged with compounds of high atomic number elements, such as for instance barium sulphite or bismuth sulphite, which shield the unwanted secondary X- radiation not self-absorbed by the anode.

The X-ray tube may be with or (preferably) without takeoff electrode (or "grid"), in the latter case the electrical field required to extract electrons from the carbon nanotube cathode (typically a few thousands V/mm) is advantageously provided directly by the electrical potential between cathode and anode, which is preferably in the range to 20kV to 100 kV). According to recent literature, direct electron extraction from the tip of carbon nanotubes of suitable characteristics is possible with electric field as low as 2000 V/mm, or less, which is achievable with the anode-to-cathode electron accelerating potential itself, given the proper distance anode-to-cathode.

In a further aspect of the invention, a detector of X-radiation is included with the tube- head, preferably (but not necessarily) mounted in proximity to the X-ray exit window, so to be exposed to the X-radiation, and provides a feedback to the control circuitry to stop the application of electrical power to the X-ray tube, and the production of X-ray, when a given amount, or threshold, of kerma is reached, thus overcoming the need to control the anodic current (mA) and the irradiation time (s) independently and accurately. In general, the radiation detector provides a feedback signal proportional to the kerma rate, and this signal is integrated by the control circuit to obtain the cumulative kerma for that irradiation event, as preset by the operator. The total air kerma is directly proportional to the current- time-product (mAs) and to approximately the square of the anode-to-cathode electrical potential.

Traditionally, great emphasis is placed on the accurate control of all the variable technique factors - i.e. the high-voltage potential (kV), the anodic current (mA), the irradiation time (s) - independently from each other, as an indirect way to accurately control the amount of air kerma and the patient radiation absorbed dose in order to obtain proper radiographic imaging without unnecessarily over-exposing the patient. However, those concepts are the legacy of an obsolete technology, where the real-time measurement of kerma and kerma rate with small, fast, and reliable solid state-detectors was impossible or impractical, and imaging was based upon radiographic films as a detectors instead of the digital electronic detectors thoroughly adopted nowadays. What is actually and ultimately necessary to control for good radiographic practice is the radiation absorbed dose (which is directly proportional to the air kerma), and to a lesser degree the "quality" or "hardness" of radiation i.e. the spectral distribution of the X-ray photons that determines the image contrast; however the contrast can be corrected by means of digital post-processing in systems with electronic image detector, therefore it is not a factor of paramount importance when digital image detectors are used.

The old technology had influenced the formulation of the prevailing technical performance standards, which in turn have reinforced the emphasis onto accurate individual control of each separate technique factor. However the technical standard are continuously evolving and have begun accepting the concept that the imparted dose, and consequently the (air) kerma, is ultimately the one important factor to control and regulate. An important advantage of the last aspect of the invention is that the X-ray tube current (or "anodic current") - and, to a lesser extent, the X-ray tube voltage - do not need to be regulated to a high degree of accuracy, which is of especial relevance with x-ray sources based upon carbon nanotube cathodes that do not have specific provisions for accurately controlling the cathode-to-anode current, and more specifically with those that do not have and use a takeoff electrode (or "grid'), since their current can be subject to significant variations among x-ray tubes and cathodes that supposedly have the same design and fabrication but is influenced by the properties of the carbon nanotubes at nanoscopic scale which may be very difficult to accurately control and replicate.

Another advantage of the last aspect of the invention is that the of amount radiation dose, or kerma, produced at an irradiation event is independent from effects eventually caused by aging of the X-ray tube, such as for instance the coating of the exit window by a layer of high-atomic-number atoms due to sputtering from the target, because they are counteracted by the feedback operation.

Note that the advantages offered by the latter aspect of the invention apply whether or not the anode of the x-ray tube is of the hooded type, and whether or not it is in intimate physical and thermal contact with other grounded metallic parts having a larger thermal capacity

This invention is especially advantageous for small radiographic sources, such as the portable hand-held ones used for dental radiology, veterinarian radiology, emergency in- field operation, non-destructive testing, etc., where the independent control of each technique factor is unnecessary, whereas lightness, compactness, and simplicity of use are of paramount importance.

Those skilled in the art would clearly understand that the various aspects of the invention here described can be practiced either in their totality or in any partial combination, without the efficacy and benefits of each individual aspect being impaired.

A preferred embodiment of the invention is depicted in fig. 1

An X-ray tube 1 with carbon nanotube cathode is mounted inside the tube-head housing 2. The carbon nanotubes in cathode 3 are mechanically supported by a suitable backing, such as e.g. a small platelet (not shown) of nickel, silicon, or metal-coated ceramics, which is attached to the glass envelope of the x-ray tube and which also mechanically supports the wehnelt 4 that is required for the focalization of electrons onto the target 5, typically made out of tungsten, and a takeoff electrode 7 (which may be e.g. a wide-meshed metallic grid).

The anode 5 is directly secured (typically bolted) to an internal surface of the housing 2, to a part of the housing which is made out of metal and is connected to the electrical ground of the system. The anode is of the hooded type, i.e. it is shaped in such a way that the target 5 hit by the flow of electrons of the anodic current - and consequently emitting X-rays - is recessed inside a deep hollowing or cavity, whereas an opening 8 on the side of the cavity provides an passage for the wanted X-rays. The flow of electrons is electromagnetically shaped by the wehnelt 4 and by the general design of the x-rray tube in such a way that the area of the target 6 that they hit, and which emits x-radiation, has the shape of a narrow strip (see from a side in fig. 1). As in conventional X-ray tubes, the target 5 on the inner face of the anode is at an angle a respect to the X-ray beam nominal axis; consequently when observed from the direction of the X-ray beam nominal axis (i.e. from the exit window) the focal strip on the surface of the target appears shortened in length (by a factor sin a) into a quasi-dot-like focal spot. The X-ray beam is filtered by a radiation filter 10 at the exit window 9. With such construction, a large part of the X-ray photons produced, that do not belong to the primary X-ray beam - the primary X-ray beam being constituted by those X-ray photons that travel through the exit window - are shielded and absorbed directly by the body of the anode, thus reducing the need for further radiation shielding.

In this embodiment, the enhancement of electrical insulation inside the tube-head is preferably achieved by non-liquid insulating media, such as for example by filling the sealed tube-head with an insulating gel, or by coating the high voltage parts with a thick layer of rubber or elastomer (not shown), since the low thermal resistance between the anode and the housing, resulting from the intimate physical contact between them, improves the dissipation of the heat produced at the focal spot, hence making unnecessary the contribution to thermal dissipation of thermal convection in a liquid insulating medium. The insulating media is charged with a compound including at least one element of high atomic number, such as barium sulphite, bismuth sulphite, or lead oxide, which completely shields the portion of stray X-radiation that is not self-absorbed in the hooded anode.

In this particular embodiment, the X-ray tube features a control grid 7, (also known as a takeoff electrode) whose electrical potential is set by a passive network such as a voltage divider (exemplified here with a network of four resistors), or by a combination of voltage divider and Zener diodes (not shown), and/or other passive electrical components (i.e. components that do not require to be powered with a separate power supply in order to perform their function); the same passive electric network provides the required difference of potential between the cathode and the wehnelt (which must be more negative at the wehnelt).

The difference of potential between anode and cathode, which consists only of a negative voltage to the cathode since the anode is at ground potential, is obtained from the input power source via an converter 13 (or an inverter if the input power source is direct current), a high- voltage step-up transformer 12, and a voltage booster 11 (also known in the art as a "diodes pump" or "voltage multiplier", in this case of Cockroft- Walton type). In this example the number of stages of the voltage booster is six, but any convenient number of stages can be used depending upon the requirements of the detailed design. The inverter is preferably of the resonant type. The operation of the inverter (for instance, the start and duration of irradiation events) is regulated by a control circuit 14, which is under the control of the operator (for instance, in setting the technique factors and initiating and irradiation).

Another preferred embodiment is further depicted in fig. 2

In this embodiment there is no takeoff electrode or grid in the X-ray tube, and the electrical field for extracting electrons from the cathode is provided directly by the anode- to-cathode potential, thus limiting the total circuitry on the cathode side to two wire, one bearing the high voltage for the cathode-anode circuit, and the other a negative potential respect to the cathode fort the wehnelt. Such difference of potential between cathode and wehnelt can be obtained with one or more zener diode(s), as generically shown in fig. 2, or a combination of zener diode(s), resistors, and/or other passive electrical components.

Furthermore, a small plate or sheet 15 made of low atomic number material, for instance aluminium or magnesium or beryllium or an alloy thereof, in inserted into the opening 8 so to protects the glass wall of the X-ray tube from long-term damage potentially caused by ions sputtered from the anode, at the same time providing at least part of the required radiation filtering to the X-ray beam.

Another preferred embodiment is further depicted in fig. 3

In this embodiment, an X-ray radiation detector 16 is mounted at a convenient location in the tube-head. Such detector provides a signal whose amplitude is directly proportional to the rate of X-radiation produced, and consequently to the air kerma rate of the primary X- ray beam. It is to be noted that the detector does not need to be positioned directly in the primary X-ray beam, or in its close proximity as shown in fig. 3 (for instance, at the exit window); in fact it can be positioned anywhere it is hit by a sufficient amount of X- radiation, since, for any given geometry, there is strict proportionality in the value of kerma rate (and kerma) at all places.

The signal from the detector is fed back to the control circuit of the high-voltage generator. When a desired total (integral) air kerma is reached, the high-voltage is shut off and X-ray production interrupted. The readout signal from the detector can be calibrated against the actual air kerma rate of the primary X-ray beam, which can easily be done e.g. in factory during the final test and calibration process. Fig. 4 depicts another preferred embodiment of the invention, analogous in all aspects to the embodiment described in fig.3 except that the anode is not of the hooded type and is not in intimate physical and thermal contact with a metallic surface of the housing (although electrically connected to the ground). This construction retains the benefits of controlling the irradiation basing upon the radiation kerma without requiring accurate control of anodic current, and may be advantageous with respect to size in those applications where the radiological load is sufficiently small that the heat dissipation from the anode is not a significant concern, and also the radiation dose load is small enough so that the additional radiation self-shielding of the hooded anode is unnecessary.

The above-described preferred embodiments and examples of the present invention are not intended to limit the scope of the invention. Modifications or alterations may be made therein without departing from the spirit and the scope of the invention as set forth in the appended claims.

Claims

An X-ray device of tube-head type, comprising:

a. a housing, including an x-radiation window to permit the x-ray beam to exit the tube-head;

b. electrical circuitry to elevate the voltage from a source of electrical power up to the high voltage required for the x-ray tube to operate;

c. means to shield a substantial portion of the unwanted x-radiation anywhere but at the x-radiation window.

d. an X-ray tube;

wherein the X-ray tube is characterized in that:

i. the electrode acting as cathode includes a multitude of carbon

nanotubes which emit electrons under the action of an electrical field; ii. the electrode acting as anode is electrically connected to the electric ground;

iii. the electrode acting as anode (ii) is in intimate physical contact with at least one metallic part of the housing having thermal capacity larger than that of the anode;

iv. the electrode acting as anode is of the hooded type, i.e. the included target for the emission of x-radiation is recessed in a deep hollowing of the anode body;

v. optionally, a takeoff electrode (or "grid"), operating at positive

electrical potential with respect to the cathode thus determining the strength of the electrical field at the surface of the cathode which causes the extraction of electrons from the cathode, is situated in close proximity of the cathode;

An X-ray device of tube-head type, comprising:

b. electrical circuitry to elevate the voltage from a source of electrical power up to the high voltage required for the x-ray tube to operate; c. means to shield a substantial portion of the unwanted x-radiation anywhere but at the x-radiation window.

d. a radiation detector that provides a feedback signal to a control circuit so that the electrical power to the anodic circuit and the X-ray irradiation is interrupted when the integrated signal from the radiation detector reaches a preset threshold which is proportional to the kerma produced by the irradiation event at a given location.

e. An electronic control circuit, that may be located either inside or outside the said housing, which controls and causes the interruption of the x-ray irradiation event basing upon the parameters set by an operator and the feedback from the said radiation detector.

f. an X-ray tube;

wherein the X-ray tube is characterized in that:

i. the electrode acting as cathode includes a multitude of carbon

iii. optionally, the electrode acting as anode (ii) is in intimate physical contact with at least one metallic part of the housing; iv. optionally, a takeoff electrode (or "grid"), operating at positive

electrical potential with respect to the cathode thus determining the strength of the electrical field at the surface of the cathode which causes the extraction of electrons from the cathode, is situated in close proximity of the cathode.

The device of claim 4 wherein the control of the amount of radiation imparted with an irradiation event is based upon the operator selection of the desired current-time- product (mAs).

The device of claim 4 wherein the control of the amount of radiation imparted with an irradiation event is based upon operator selection of the desired air kerma.