EP0529505A1 - Method for producing high contrast diagnostic x-ray exposures and circuit arrangement therefor - Google Patents

Abstract

In order to obtain high-contrast, diagnostic X-ray photographs using an X-ray generator which is connected to an AC voltage network and has a converter, high-voltage rectifier and X-ray tube connected to the output of the latter, and having a control unit, the converter increasing the frequency of the AC current to the converter frequency, as a frequency converter, and increasing its voltage to the desired high voltage, as a voltage converter, and the switching unit controlling the tube current, tube voltage and exposure duration, the voltage of the X-ray tube is reduced during the exposure time at least once from an upper limit value which can be predetermined to a lower limit value which can likewise be determined; a circuit arrangement for the purpose is furthermore proposed, in which at least one measurement input (E1; E2; E3) for the voltage to the X-ray tube (3), as well as a modulation stage (5), are allocated to the converter (2), which modulation stage (5) passes on the measurement of the tube voltage as a control signal to a modulator (4) which produces the control signal for reducing the tube voltage and increasing it again and is connected to the control section (2.2) and to the grid of the electron tube (7) which is connected upstream of the X-ray tube (3), and to the grid of the X-ray tube (3'). <IMAGE>

Description

The invention relates to a method for increasing the contrast of diagnostic X-ray recordings with an X-ray generator connected to an electrical AC voltage network with a high-voltage rectifier and an X-ray tube connected to it and with a switching unit that controls the tube current and tube voltage as well as exposure time, in particular those with a converter, on the one of which is provided with a charging capacitor High-voltage output, the X-ray tube is connected, the converter as frequency converter to the frequency of the alternating current supplied to the rectifier Converter frequency and as a voltage converter whose voltage is increased to a value corresponding to the desired high voltage; it also relates to a circuit arrangement therefor.

X-ray generators for diagnostic purposes are operated in different voltage ranges and with different current intensities and with different switch-on times, the level of the voltage applied to the X-ray tube determining the spectral composition of the X-ray radiation and radiation yield, the current strength determining the radiation intensity and both together determining the dose rate, while the switch-on duration determines the Exposure time and thus the dose is given. All three factors determine the exposure of the X-ray film and thus influence the quality of the X-ray images. This influence is essentially due to the change in the spectral composition of the radiation and thus in connection with the different absorption capacities of the types of tissue lying in the beam path with their different absorption behavior characterized by the half-value layer thickness. These different absorptions ultimately provide the different blackening of the x-ray image and are therefore a prerequisite for the contrasts occurring in the x-ray image, which - because of the spectral dependence of the tissue absorption - but also from the spectral distribution of the x-ray light and - because of the gradation of the radiation dose X-ray film (if necessary with intensifying film) - depend on the dose. The variations in the spectral distribution of the X-ray light and the variations in the dose therefore lead to different image qualities that can be used in a meaningful manner for the different recording conditions. For example, the rib structure with a lower tube voltage and the soft tissue structure with a higher tube voltage will appear more clearly in a chest scan, the latter because the radiation of the strongly absorbing ribs with the harder radiation is otherwise present due to the absorption Compensates for loss of information.

These factors are not independent of each other, since a certain density of blackening of the X-ray film can be assigned to a certain dose of radiation, the tube voltage, tube current and duty cycle are necessarily related the 5th power of the tube voltage can be assumed. It goes without saying that the properties of the X-ray film must also be taken into account. With these considerations to ensure the necessary contrast, however, the radiation exposure of the body, in particular its skin surface, must not be overlooked with an X-ray dose.

Generally, an increase in the tube voltage leads to a reduction in the duty cycle and thus avoids blurring due to movement; it allows the focal spot on the tube anode to be reduced, with the success of a greater sharpness of the image due to improved radiation geometry; it enables a greater distance between the X-ray tube and the patient, which is associated with less distortion, less geometric blur and less secondary radiation. Lowering the tube voltage means an increase in the duty cycle. In addition, a reduction in the tube voltage also means a shift in the spectral energy distribution towards a softer radiation, which increases the stress on the skin. Optimal image quality will therefore have to start from a tube voltage that corresponds to the organ to be displayed and the thickness of the patient to be irradiated. This optimization is limited on the one hand by the structure to be represented and on the other hand by the resilience of the X-ray tube with tube limit voltage and tube limit current. Modern converter technology is often used to generate the tube voltage, with one initially corresponding to the mains voltage DC voltage is generated, which has a ripple corresponding to the type of rectification, the converter frequency lying in the range of approximately 25 kHz already enabling smoothing with a (relatively) small charging capacitor. The voltage present at the charging capacitor is now conducted via the X-ray tube, which is discharged from the tube current, and the recharging is carried out by feeding in rectified converter current, the supply resistance determining the upward and internal resistance of the X-ray tube to determine the discharge. In equilibrium there is a constant tube voltage, which only has a slight ripple during the charging phase. This shows, however, that a converter high voltage - particularly at higher powers - comes very close to a smoothed DC voltage in terms of the type of high voltage. This means that the radiation behavior of the X-ray tube can largely be regarded as brake radiation which can be assigned to this tube voltage and which contains a certain proportion of characteristic radiation from the anode metal, with the result that the contrast behavior drops significantly. Such a reduced contrast behavior is, however, only desired and desired in the case of hard radiation images, for example the thorax.

This leads to the technical problem on which the invention is based, according to which a method is to be specified with which high-contrast x-ray images are obtained using converter technology; in addition, a circuit arrangement is to be specified for this.

This object is achieved according to the invention in that the voltage applied to the X-ray tube is reduced at least once from a predeterminable upper limit value to a likewise predeterminable lower limit value during the exposure time. The tube voltage thus runs through all voltage values between an upper limit value and a lower limit value in an exposure interval, the spectral composition of the emitted X-rays follows this voltage change, so that the spectral composition required for the exposure situation and necessary for maximum contrast is thereby achieved.

The lowering is advantageously carried out several times, preferably quasi-periodically, with the tube voltage being raised to the upper limit value after the lowering. It is advantageous if the interruptions of the higher-frequency alternating current supplied to the rectifier during the exposure time take place quasi-periodically, the frequency to be assigned to this quasi-periodic interruption being in the range from 1/100 to 1/5 of the converter frequency. At normal converter frequencies of 20-40 kHz, this means an interruption to which a frequency of 0.2-8 kHz can be assigned. With recording times in the millisecond range, this ensures that the tube voltage runs through the entire voltage interval at least once during the exposure, that is, it oscillates between the specified maximum and the likewise specified minimum value. It is advantageous if the time constants for the drop in the tube voltage and its increase are chosen to be the same. The inductance which interacts with the charging capacitor, in general the load-dependent leakage inductance of the high-voltage transformer, forms an oscillating circuit, the switching elements in the oscillating circuit also being controllable.

To lower and raise the tube voltage again, the higher-frequency alternating current which causes the charging of the charging capacitor and is fed to it at the converter frequency is interrupted quasi-periodically and this interruption is canceled when the predetermined upper limit value of the tube voltage has been reached, or when the predetermined one is also reached lower limit of the tube voltage has been reached. Alternatively, the negative grid bias of one of the X-ray tubes is used to lower and re-raise the tube voltage upstream electron tube quasi-periodically increased and decreased when the predetermined upper or the likewise predetermined lower limit value of the tube voltage is reached. Again, as an alternative to this, the negative grid bias of a grid control of the x-ray tube is increased or decreased quasi-periodically to lower and raise the tube voltage when the predetermined upper or the likewise predetermined lower limit value of the tube voltage has been reached. If the charging voltage of the charging capacitor or the voltage on the X-ray tube is measured, when the maximum value of the anode voltage of the X-ray tube is reached, either the oscillation can be suppressed so that the charging capacitor discharges to the lower limit value, or the measured value is converted into a grid bias, which either controls the electron tube upstream of the x-ray tube or controls the x-ray tube itself.

In the first alternative, the charging capacitor is charged so that the upper limit is reached even under load; the variation of the voltage is brought about by the discharge of the charge through the X-ray tube and the inflow of charge from the high-voltage rectifier, the duration of a reduction interval, together with that determined by the temperature of the cathode of the X-ray tube, determining the amount of the reduction, given the capacitance of the charging capacitor. When the lower limit is reached, the interruption is canceled and the charging capacitor starts charging again. In the other two alternatives, the upstream grid-controlled tube acts as a series resistor, the current flowing in the high-voltage circuit is a saturation current determined by the temperature of the cathode of the X-ray tube, the grid bias changes the internal resistance of the upstream electron tube, so that its voltage drop causes the voltage across the X-ray tube changed, the charge of the charging capacitor is kept constant and adjusted so that in the time interval the lowering to the desired lower Limit of the tube voltage is reached. If the X-ray tube is not operated in the area of saturation, but the tube current is determined by the bias of the grid of the X-ray tube, the third alternative can be used: Here, the variation of the grid bias causes an increase or decrease in the tube current, so that the charging capacitor more or unloaded less quickly. Since the voltage at the charging capacitor is equal to the tube voltage and this is due to the respective balance between charging (or recharging) and removal, this must fluctuate between a predefinable upper and a likewise predeterminable lower limit value.

The desired degree of ripple in the anode voltage of the X-ray tube is achieved with these variants: The tube voltage passes through a voltage interval which specifies the spectral compositions of the X-ray radiation. As the tube voltage is lowered, the short-wave cutoff frequency of the radiation spectrum is shifted towards longer wavelengths, and the radiation becomes softer. In this case, even if the voltage required for the excitation of a characteristic radiation is undershot, this can be omitted, but this is of subordinate importance because of the prevailing intensity of the braking radiation. For the sake of simplicity, this interplay of lowering and raising is referred to here as ripple or modulation, the degree of ripple (or modulation) being assumed as the ratio of maximum voltage to minimum voltage. It is also advantageous if means are provided with which the curve shape of lowering and increasing the voltage can be influenced.

It is preferably provided that the values for the lowering and re-raising of the tube voltage, for its quasi-periodic frequency and / or for the "modulation" degree of the X-ray tube voltage and thus for the ripples that can be achieved in accordance with the type of Desired x-ray image are stored in an application memory and retrieved from the memory and are preferably fed to the processor for controlling the interruption in preparation for the image. Automation is thus achieved, which enables exposures or fluoroscopes to be carried out in a simple manner, even under special circumstances, with only the exposure situation having to be specified based on the stored values. It is advantageously provided that the application memory of the processor additionally contains correction values which can be called up in connection with a desired x-ray image in addition to the application data for superimposing them. On the one hand, this prevents the X-ray tube from being overloaded, on the other hand, the radiation exposure of the patient can be kept clearly within limits, especially since repeated procedures can generally be dispensed with in the procedure. In addition, monitoring of the x-ray system itself is also possible if the measured values for voltage variation and degree of ripple are fed back to the control unit, for comparison with the preset values for the upper and lower limit values, taking into account the maximum values of the x-ray tube voltage and current as well as the permissible current flow time and to correct them.

A circuit arrangement which is advantageous for carrying out the method is provided in that the converter has at least one measurement input for the tube voltage, which forwards the measured tube voltage as a control signal to a modulator generating the control signal for lowering and raising the tube voltage, which is connected to the control part or to the Grid of the electron tube upstream of the X-ray tube or connected to the grid of the X-ray tube. With this arrangement, the signal necessary for the control of the tube voltage can be obtained and either to interrupt the converter frequency, to generate the grid bias of the grid-controlled electron tube upstream of the x-ray tube or finally to generate the grid bias of the x-ray tube self.

Advantageously, the control connections of the thyristors of the converter are connected to the outputs of a phase control, the current pulses which ignite the thyristors are phase-shiftable with respect to the converter oscillation in accordance with the desired upper limit value of the voltage for setting their upper limit value, and that the phase control has means which each after reaching this upper limit value blocks the delivery of these ignition pulses until the lower limit value of the voltage is reached, for which purpose the phase gating control on the one hand provides a connection for taking over the in-phase converter frequency and on the other hand makes a connection for taking over the anode of the x-ray tube, preferably the one connected to the high-voltage rectifier Charging capacitor applied voltage is supplied directly or via a voltage converter. Such a circuit arrangement operates under the specified conditions as a frequency generator and is thus able to control the thyristors of the converter accordingly, its frequency being specified by a series resonant circuit contained in the converter. The quasi-periodic charging and discharging of the charging capacitor results in voltage fluctuations which extend from the predetermined upper limit value to the likewise predetermined lower limit value. These voltage fluctuations run with their own time constants, which depend on the external and internal resistances. This type of circuit leads to a type of "modulation" of the converter frequency, since the converter alternating current of higher frequency, which charges the charging capacitor via the high-voltage rectifier, is only present quasi-periodically. It is therefore only possible to generate a charging current for the charging capacitor in the phase position following the lower limit value of the tube voltage (the modulation frequency of the tube voltage), as a result of which its charging is limited in time and thus the upper limit value is predetermined, although the output voltage of the charge affects the upper limit. The range for the "open" phase angle (ie for the phase angle range in which a charging current is generated) must therefore be determined taking into account the lower limit value. Therefore, an advantageous further development is given in that the converter has a timer for setting the opening time and thus the charging time of the smoothing capacitor. With this timer, the time constant acting for charging the charging capacitor can be changed. In a further embodiment, the x-ray tube circuit has a controllable resistor, preferably a triode, connected upstream of the anode of the x-ray tube, the internal resistance of which and thus the time constant of the discharge of the charging capacitor can be controlled. Alternatively, it is proposed that the x-ray tube have a control grid for controlling its internal resistance and thus the time constant of the discharge of the charging capacitor via the grid bias. With this embodiment, the time constant determining the discharge of the charging capacitor can be changed. If both changes are combined, the two time constants can be adjusted to each other. This creates an at least sine-like voltage curve on the charging capacitor. In order to further approximate the waveform of a sine wave, it is proposed that an inductor be provided in the line carrying the charging current of the charging capacitor, for deforming the waveform of the current emitted by the converter, the inductor and the capacitor forming a resonant circuit with a resonance frequency close to that Break frequency. The resonance frequency of this resonant circuit determines the frequency with which the charging and discharging of the charging capacitor take place, and thus the repetition frequency for the "ripple". With the help of the procedure according to the invention, all desired curve shapes can be simulated.

The converter or the modulation stage is preferably one Assigned to a microprocessor to control the interruption frequency and thus the "modulation" level of the X-ray tube voltage. It is also advantageously provided that the microprocessor is assigned a working memory and a further mass memory, the mass memory containing files in which recording or fluoroscopy parameters are stored, which are required for a desired recording in the working memory for comparison with the set or measured values Values are transferable. The device is brought so far that it can interact with a computer or even contains an integrated computer. The connections are made via parallel bus structures, whereby serial connections are not excluded. With the help of the stored values, the usual recording situations can be specified so far that they can be called up and are thus available. This is advantageously carried out in connection with a computer, for example a personal computer, the keyboard of which enables the necessary commands to be entered and the monitor of which allows the entire recording or screening process to be monitored. It goes without saying that the computer can also be integrated, with a keyboard and monitor.

In an advantageous development, the processor is connected via at least one internal connection bus to an internal mass memory designed as a permanent memory, the output signals of which are carried out together with those of the working memory via an internal data bus, these being connected to the inputs and the output the intended digital / analog or digital / analog converter. In this design, the direct interaction of the control device with the X-ray system is clear: the input values for tube voltage, tube current and tube heating are implemented in such a way that the variations necessary for the exposure situation are generated. The memories can record and save new recording situations, they are learnable.

Fig. 1:

Kurvenform des Verlaufs der Röntgenröhrenspannung bei einer 6-Puls-Gleichrichtung an einem Drehstromnetz;

Fig. 2:

Kurvenform des Verlaufs der Röntgenröhrenspannung bei Konverterspeisung;

Fig. 3:

Kurvenform des Verlaufs der Röntgenröhrenspannung bei Konverterspeisung mit quasi-periodischer Auf-und Entlung des Glättungskondensators (Konverterfrequenz relativ niedrig gegenüber Unterbrechungsfrequenz);

Fig. 4:

Kurvenform entsprechend Fig. 3, jedoch höhere Unterbrechungsfrequenz;

Fig. 5:

Prinzip-Schaltbild einer mit einem Konverter gesteuerten Röntgenröhre, Konverter gesteuert;

Fig. 6:

Prinzip-Schaltbild wie Fig. 5, jedoch mit Gittersteuerung anstelle der Steuerung des Konverters,
Fig. 6a: Mit der Röntgenröhre vorgeschalteter Triode,
Fig. 6b: Mit gittergesteuerter Röntgenröhe;

Fig. 7:

Schaltbild eines gesteuerten Konverters mit Ansteuerung durch einen PC;

Fig. 8:

Prinzip-Schaltbild eines Röntgengenerators mit Microprozessor zur Steuerung der Aufladung des Glättungskondensators,
Fig. 8a: Steuerteil, Fig. 8b: Ein- und Ausgänge.

The essence of the invention is explained in more detail with reference to the circuit diagrams shown in Figures 1 to 8 and the waveforms of the voltage profiles; show:

Fig. 1:

Curve shape of the course of the X-ray tube voltage in a 6-pulse rectification on a three-phase network;

Fig. 2:

Curve shape of the course of the X-ray tube voltage with converter supply;

Fig. 3:

Curve shape of the course of the X-ray tube voltage with converter supply with quasi-periodic opening and unloading of the smoothing capacitor (converter frequency relatively low compared to the interruption frequency);

Fig. 4:

3, but higher interruption frequency;

Fig. 5:

Principle circuit diagram of an X-ray tube controlled by a converter, converter controlled;

Fig. 6:

Schematic diagram as in FIG. 5, but with grid control instead of control of the converter,
6a: Triode connected upstream with the X-ray tube,
Fig. 6b: With grid-controlled X-ray height;

Fig. 7:

Circuit diagram of a controlled converter controlled by a PC;

Fig. 8:

Principle circuit diagram of an X-ray generator with a microprocessor for controlling the charging of the smoothing capacitor,
Fig. 8a: control part, Fig. 8b: inputs and outputs.

FIGS. 1 and 2 show the typical course of the high voltage on the X-ray tube when using a conventional rectification with a 3-phase full-wave rectifier to generate the DC voltage applied to the anode (FIG. 1) and when using a converter known per se ( Fig. 2). The main difference lies in the ripple of the DC voltage (also "humming"), which the 6-pulse voltage from the 50 Hz three-phase network is around 30%, and the converter drops to below 10%. This is due to the use of a (relatively) high frequency of around 20-30 kHz, with which the high voltage is generated here. The advantages achieved in this way are the savings in iron and copper in the area of the transformer, which, at higher frequencies, manages to transmit the same power with the same transmission ratio on the one hand with a smaller core and on the other hand with fewer turns. In addition, a small charging capacitor is sufficient at the higher frequencies to reduce the ripple.

FIGS. 3 and 4 show the curve shapes of voltages applied to X-ray tubes and generated by means of a converter, which are provided with a ripple according to the invention. These curves consist of the rising charging branches and the falling discharging branches of the charging capacitor (each related to the voltage at the X-ray tube anode). For the sake of clarity, the ripple resulting from the converter frequency, which overlaps the charging branches in accordance with the illustration in FIG. 2, has not been drawn in, although it goes without saying that the falling branches do not show any such ripple. While FIG. 3 shows a curve shape with a degree of ripple (or modulation) of approximately 50% (FIG. 3a) or approximately 75% (FIG. 3b) with an interruption frequency of approximately 1 kHz, the interruption frequency lies in the voltage curve according to FIG 4 significantly higher at about 3 kHz, the degree of ripple (or modulation) corresponding approximately to that of FIG. 3a. The choice of the interruption frequency and degree of ripple determine external parameters, which depend on the type of recording situation and the X-ray generator, the maximum voltage at the X-ray tube also being able to be determined via the charging pulses resulting from the higher-frequency AC voltage, which for a few pulses when the anode voltage reaches the maximum value suspend until the anode voltage has dropped slightly below the value of this maximum value.

Figures 5 and 6 show basic circuits for the connection of X-ray tubes, in which a converter takes over the generation of the "modulated" anode voltage (Fig. 5), in which the discharge and thus the anode voltage of the X-ray tube is achieved by control with a lattice-controlled high-vacuum tube ( 6a, 6b). In all cases, voltage is supplied from a three-phase network (not designated in more detail) via the circuit breaker 1, which can also be designed with controlled semiconductor switching elements. The voltage is fed to the converter 2, in which the voltage converted in the vibration generator 2.1 to the converter frequency in the transformer and rectifier part 2.3 is converted into the anode voltage of the x-ray tube 3. In the embodiment with a controlled converter (FIG. 5), the voltage applied to the charging capacitor connected downstream of the rectifier is detected by the voltage measuring set 4 and reported to the modulator 5, which in turn releases the modulation control 2.2 until the maximum voltage value at the charging capacitor as specified is reached, then the converter oscillation stops and only releases it again when the voltage has reached a (likewise predetermined) lower voltage value at the charging capacitor.

In the embodiments with control via a grid-controlled high vacuum tube instead of controlling the converter (or the converter vibrations), as shown in FIGS. 6, a triode 7 is connected upstream of the x-ray tube as a grid-controlled high vacuum tube, the voltage at the output corresponding to the anode voltage of the x-ray tube 3 7.1 is removed from the triode and fed to the voltage measuring set 4, which evaluates the value of the applied voltage as described above and reports to the modulator 5, which then in turn adjusts the grid bias of the triode 7 so that the internal resistance of the triode 7 and thus the voltage drop across it the anode voltage of the X-ray tube 3 assumes the desired value. If an X-ray tube 3 'with a control grid is used, their anode current can be controlled with the grid voltage, the anode current representing a load for the charging capacitor and, depending on the excess charge or deficit, increasing or decreasing the anode voltage. In both cases, the grid control intervenes in the discharge of the charge capacitor, the discharge of the charge capacitor predominating when the grid is "open" and its charge predominating when the grid is sufficiently biased negatively, the grid control itself acting as a "switch" that interrupts the Converter vibration and thus charging voltage replaced. In each of the cases, the anode voltage of the x-ray tube 3 is set by means of the control described so that the x-ray radiation lies in the desired spectral range.

FIG. 7 shows a (schematic) circuit of a converter which can be controlled by a PC and is supplied from the three-phase network via the connections U, V, W. The center conductor M _P ensures a connection to the neutral point of the three-phase network and also allows a corresponding AC voltage to decrease, for controlling the main switching relay HR or the heating circuit HK for heating the cathode of the X-ray tube RR. The entire control of the X-ray generator is effected with the aid of a computer, for example a PC with an input terminal and monitor as an output unit with a comprehensive control panel STP. After the main switch (not shown) is switched on, a voltage is provided for the heating circuit HK, and further the voltage for switching through the main switching relay HR for activating the X-ray tube RR. In addition, the tube voltage required for the present case is preselected on the control panel STP and the heating current required for the required tube current is set, which - since the cathode must be preheated due to its inertia - flows before the start of the recording, since the cathode temperature for the anode current of the X-ray tube is decisive. These values as well as the preselection values corresponding to the shooting situation can be displayed on the monitor.

When the recording is triggered, the main switching relay HR then connects the mains voltage to the high-voltage generator, which is then converted directly into the mains rectifier NGL, which is designed as a 3-phase double-path rectifier, and fed to the current transformer part STW. Here, the voltage symmetrized by the same capacitors C1, C2 is applied to the thyristors Th1 and Th2, to which the diodes D1 and D2 are connected antiparallel, the circuit via one of the capacitor C3 and the (leakage) inductance of the power transformer TR1 formed resonant circuit is closed. Due to the series resonant circuit, the direct voltage is converted into a higher-frequency alternating voltage, the frequency of which is determined by the capacitor capacitance C3 and the leakage inductance of the transformer TR, which is large when idling and decreases with increasing load, which causes an increase in frequency with increasing Load and thus a reduction in ripple means. The AC voltage is converted in the power transformer TR to the voltage required for the desired operation of the X-ray tube RR, which is then rectified in the high-voltage rectifier HGL (indicated only as a diode) and smoothed with the capacitor C4. The increasing frequency with increasing load together with this smoothing causes the (relatively) low ripple of the curve shape of the direct voltage generated by a conventional converter circuit (see FIG. 2). It goes without saying that the converter circuit is not limited to the use of a thyristor pair, but in the same way e.g. can also be implemented with a bridge circuit.

Since the power transformer has a fixed transformation ratio, in order to adapt the required tube voltage and thus the spectrum of the X-ray radiation as required, the phase of the control current pulses which control the thyristors Th1 and Th2 is shifted with the aid of a phase shifter PHS corresponding to a phase control, so that in the end result an output voltage corresponding to the desired voltage at the charging capacitor C4 results. With this firm and (almost) smooth tube voltage, there is now a spectrum of X-ray radiation which is not suitable for all exposures and fluoroscopy, but is the optimum for the selected exposure situation, whereby in the event of patient-related deviations from the "ideal" exposure situation necessary or desired corrections can be entered via the input terminal, if necessary using selection tables or selection menus stored in the computer.

If the control of the thyristors Th1 and Th2 is interrupted quasi-periodically with the aid of an interval switch IVS, the high voltage available for the X-ray tube RR to the smoothing capacitor HGL drops after each interruption until it is recharged by opening the thyristors Th1 and Th2. In this way, an alternating voltage with a preselectable ripple that is adapted to the recording or the fluoroscopy and thus a shift of the center of gravity of the radiation spectrum lying in the desired sense is achieved. The switch-off interval can take place both via the fall time of the voltage at the charging capacitor HGL (time constant of the charging capacitor and internal resistance of the X-ray tube) and via the direct determination of the value of the falling voltage at the charging capacitor. When determining the time interval, the preselected heating system must be taken into account, which determines the saturation current of the X-ray tube via the cathode temperature, which thus receives an internal resistance that is dependent on the heating current. For safety reasons, in order to avoid exceeding the maximum voltage permitted for the X-ray tube, the upper voltage value is measured and monitored; In order to achieve the ripple that is favorable for high-contrast recording, this monitoring is extended to the lower voltage value. Therefore, the opening of the thyristors Th1 and Th2 is advantageously coupled to this voltage measurement and thus the way is opened via the control panel STP the upper and lower limits to specify for the voltage so that during the recording or fluoroscopic interval the tube voltage drops at least once from the high voltage specified on the control panel STP to the likewise specified low voltage. These specifications also allow the circuit to be adapted to the circuits of X-ray generators which work with a grid-controlled vacuum tube (see FIG. 6a) connected upstream of the anode of the X-ray tube, or in which the X-ray tube RR 'itself has a grid control (see FIG. 6b ). In these cases there is no need to interrupt the converter oscillation, the charging voltage is constant here. The anode voltage is changed here by reducing or increasing the discharge of the charging capacitor, since here the current flowing in the anode circuit is not the saturation current corresponding to the temperature of the cathode of the x-ray tube, but is determined by the grid bias.

FIGS. 8 (where FIG. 8b continues FIG. 8a in accordance with the entered interfaces) show the computer-aided control of the converter in a schematic basic circuit. The microcontroller U1 as CPU with external power supply (here indicated by a mains rectifier NGL with outputs for the supply voltage VCC and for a reference voltage VRZ) and external clock generator (with quartz TG) is the heart of the control; it works according to a program stored in a ROM memory U4, which is read in via an external I / O connection, this I / O connection being connected to a computer, for example a PC, but at least to a mass storage device (with the latter, the CPU also controls this mass storage as a microprocessor). When connected to a PC, it can be entered and read directly via its input unit (keyboard or mouse) and its output unit (monitor); It goes without saying that when using the controller without its own PC, such an input and output option should be provided in order to operate the controller effectively to enable. For this purpose, additional external connections can be provided via the I / O bus I / O or the external control bus STB (ex). Incidentally, here the electronic components are connected to one another in the manner of switching technology customary in computers, the connections for the supply voltage VCC or the reference voltage VRZ being merely indicated by full circles.

The externally incoming commands lead to the activation of the CPU U1, which issues the corresponding commands via the control bus output "S" and the bus outputs "C" and "B". The outgoing signals via the address bus ADB are fed to the latch U2 and the ROM U4. With the help of the versions for exposure processes and limit loading of the X-ray tube stored in the ROM U4, the corresponding control commands are output via the data bus DDB, which on the one hand go back to the CPU U1 and on the other hand are fed to the inputs and outputs. The signals arriving at a first, programmable logic gate U3 are converted into control signals for the ROM U4, the RAM of the RAM U5, the I / O unit U6 and the general read / write signal R / W and via appropriate bus or line connections output: The ROM U4 receives the control signal via the control line ROSL, the RAM U5 via RASL and the I / O unit U6 via IOSL. The read / write signals for all are output via the control line RWSL. A second programmable logic gate U6 receives the corresponding signals for measurement and setpoints via the data bus DDB and via the signal which switches the converter on and off. From this, the second programmable logic gate U6 derives both the START signal, which is output via a separate START line, and the input / output signal, which runs over the input / output bus and the input and output units U7, U9, U11 and U14 activated.

The input connections E1, E2 and E3 for the actual values of tube voltage U _tube , tube current I _tube and heating current I _heating (or in general _Heating power N _heating ) are essentially of the same design: each has an operational amplifier connected to the supply voltage and to the reference voltage, the two inputs of which are provided with a resistance network, for adaptation to the connected circuit, the operational amplifier adjusting the signal level at the desired output impedance manufactures that can be processed by the downstream analog / digital converters U7, U9, U11. The input signals from E1, E2 and E3 digitized in this way run via the data bus DDB to the CPU U1 and to the RAM U4, in order to be compared there as actual values with the desired values. The CPU U1, with the RAM U5 taking on the role of the main memory required for the calculation, calculates the necessary control commands that are sent to the digital / analog converter U14 and, via this, a process amplifier U13 as an adaptation element and level converter, from the deviation detected in this comparison is connected to the converter via output A1. It goes without saying that the number of inputs and outputs need not be limited to 3 or 1 in accordance with the selected exemplary embodiment. For this purpose, the corresponding continuations of the control lines and the data bus are indicated by dashed lines in FIG. 8b. It goes without saying that the address bus can also continue to be used if further RAMs are used (indicated by dashed lines).

The general rule here is that the operation of the microcontroller U1 takes place according to the operating system stored in the ROM U4 in connection with the program stored in it and called up or a program read into the RAM U5 from an external source. Particularly for the version of the lattice-controlled X-ray tube or the one with an upstream control triode, the ROM U4 also has the current program as well as the limit values for tube voltage, tube current and anode power loss that apply to the connected X-ray tube or the upstream control triode (or tetrode) are stored so that they can be called up. It goes without saying even that operating values for constantly recurring exposure situations are stored as such, advantageously with a short command. The microcontroller U1 is controlled via a (possibly externally connected) control bus (with the individual control lines for RESET, IQR, R / W, E; START, WR, RD, LOAD, MSB, LSB), which controls the microprocessor U1 with the connected electronic components, the (edge-triggered) latch U2, the first programmable logic gate PAL U3, the ROM U4 for the operating system of the microprocessor U1, the RAM U5 with those read in by the PC via the I / O connection Program parts, the second programmable logic gate PAL U6, and the analog / digital converters U7, U9, and U11 as well as the digital / analog converter U15 and their OP amplifiers U8, U10 provided in the input and output for adaptation, U12, U16 connects with them and with each other. An input bus (A-Bus of the MC U1), which is connected to an input terminal, allows the selection of the operating state of the X-ray tube desired for the next exposure and the entry of specifications that deviate from the standard values. The data exchange between the CPU U1, the ROM U5 and the main memory as well as the input A / D converters U7, U9, U11, U13 and the output D / A converter U15 takes place via an internal data bus connecting these modules DDB (B-Bus of the CPU U1), the associated memory addresses via an internal data bus ADB (C-Bus of the CPU U1) at least to the latch U2, which generates address information from the data information, which allows the read-only memory of the Addressing the ROM U4 and the RAM of the RAM U5 address-wise. The values for the tube voltage kV-IN (E1), the tube current mA-IN (E2), the heating current A _H -IN (E3) and for in are also fed via the internal input bus DDB via the analog inputs of the control RAM U5 for comparison with the specifications obtained from the ROM U4 or entered via the terminal, these representing the TARGET values and that of the ACTUAL values. It goes without saying that further inputs E4 to EN can be provided, wherein the number of which could only be given by an address limit. In the case of grid-controlled X-ray tubes or if a grid-controlled vacuum tube, such as a triode or a tetrode, is connected upstream, this control is sufficient, since here a high voltage coming from any high-voltage source is present at the anode of the X-ray tube.

For X-ray generators that are equipped with a converter generator and that already work in X-ray systems and that are to be converted by installing a modulation stage, the control system must be varied: The maximum high voltage that builds up with the highest charging frequency (approx. 25 kHz) is achieved via a certain time measured by the control unit. The steepness of the voltage rise in the voltage on the charging capacitor gives the possibility of calculating the interruption frequency with which the high voltage can be modulated (and thus the charging of the charging capacitor can be interrupted). The average voltage at the charging capacitor corresponds to the equilibrium between charging and discharging, the ripple or modulation being determined by the individual time constants.

The modulation takes place in such a way that the high-voltage ACTUAL value is measured up to its maximum (via kV-IN E1). Up to this point, the output value of the modulation level is equal to the input value, so it is kV-OUT = kV-IN. When the setpoint of the high voltage at the charging capacitor is reached, the freely oscillating converter is blocked for a number of recharging pulses. The drop in the voltage at the charging capacitor is monitored, so that this means for the converter that the charging of the charging capacitor stops up to a lower discharge point corresponding to the minimum value of the anode voltage (FIGS. 3a, 3b). If the modulation unit now holds the high-voltage value kV-OUT at the maximum value for a calculated time, it pretends to the converter that the voltage is above the value at which reloading must begin; therefore there is no reloading as long. The actual high voltage thus drops more or less quickly to a point specified by the modulation unit in accordance with the specification, for example to 30% of the high voltage value kV-OUT, in relation to the flowing tube power which is taken from the charging capacitor. However, since the modulation unit measures the course of the high voltage, it can recognize when this point has been reached. At this point in time, the "real" ACTUAL value of the high voltage kV-ACTUAL is output as the kV-OUT value, which influences the recharge control of the converter in such a way that the high voltage is brought back to its maximum value as quickly as possible or depending on the preselected ripple, the waveform and / or steepness of the voltage rise can also be selected.

In the case of a converter or DC voltage generator to be newly constructed, the modulation unit can of course be integrated into the basic concept. The duration of exposure is obtained via a timing element (not shown in more detail), which acts on a switching element which generates or deactivates the medium-frequency low voltage. It seems obvious that both further input A / D converters can be provided for further input parameters and further output D / A converters if there is a need for this. Such a need can be given, for example, by the fact that the setting of the exposure duration is also taken over by the converter controller, with the dose rate sensed by a radiation indicator being additionally used as an input parameter for limitation. In this case, for example, the second output signal would act on the converter in such a way that the higher-frequency low voltage is only generated during the "open" time specified by this output signal, or that the pulses necessary for conversion to the high voltage are generated outside of this "open" Time be suppressed. It goes without saying that this can also be done with a Grid control of the X-ray tube can be achieved, in which the current flow is suppressed when a permissible dose rate is exceeded via a far negative bias of the grating, this far negative grid bias being reduced during the "open" time, so that the X-ray tube has tube current during this time leads, which can itself be regulated via the amount of the grid bias. Finally, it goes without saying that the same can also be done with a triode (or tetrode) connected in series with the x-ray tube. The ripple of the edges of the "modulations" is determined by the converter frequency: When the charging capacitor is charged, the converter voltage is applied to it via a rectifier with a ripple of a few percent (hum) corresponding to the converter frequency, whereas the converter frequency is when the charging capacitor is discharged generally suppressed so that the flanks corresponding to the discharge have no ripple in this case.

Claims (19)

Method for increasing the contrast of diagnostic X-ray recordings with an X-ray generator connected to an electrical AC voltage network with converter, high-voltage rectifier and X-ray tube, which is connected to its high-voltage output provided with a smoothing capacitor, and with a control unit, the converter acting as a frequency converter, the frequency of the alternating current supplied to the rectifier to converter frequency and as a voltage converter, the voltage of which increases to a value corresponding to the desired high voltage, and wherein the switching unit controls the tube current and tube voltage as well as the exposure time, characterized in that during the exposure time the voltage applied to the x-ray tube at least once from a predeterminable upper limit value to one predeterminable lower limit is also lowered. A method according to claim 1, characterized in that the lowering takes place several times, preferably quasi-periodically, with the tube voltage being raised to the upper limit value after the lowering. Method according to claim 1 or 2, characterized in that the quasi-periodic lowering and re-raising of the tube voltage takes place at a frequency which is in the range from 1/100 to 1/5 of the converter frequency. Method according to one of Claims 1 to 3, characterized in that, for lowering and raising the tube voltage again, the higher-frequency alternating current, which causes the charging capacitor to be charged via the high-voltage rectifier and is supplied with the converter frequency, is quasi-periodically interrupted when the predetermined upper limit value of the tube voltage is reached and this interruption is canceled when the likewise specified lower limit value of the tube voltage is reached. Method according to one of Claims 1 to 3, characterized in that, in order to lower and raise the tube voltage again, the negative grid bias of an electron tube upstream of the X-ray tube is increased and decreased quasi-periodically when the predetermined upper or the likewise predetermined lower limit value of the tube voltage is reached. Method according to one of claims 1 to 3, characterized in that to lower and raise the tube voltage again, the negative grid bias of a grid control of the X-ray tube is increased or decreased quasi-periodically when the predetermined upper or the same predetermined lower limit of the tube voltage is reached. Method according to one of claims 1 to 6, characterized in that the values for the reduction and re-increase of the tube voltage, for its quasi-periodic frequency and / or for the "modulation" degree of the x-ray tube voltage and thus for the ripples that can be achieved accordingly the type of the desired x-ray image are stored in an application memory and retrieved from the memory and are preferably fed to the processor for controlling the interruption in preparation for the image. Method according to Claim 7, characterized in that the application memory of the processor additionally contains correction values which can be called up in connection with a desired x-ray image in addition to the application data for superimposing them. Method according to one of Claims 1 to 8, characterized in that the measured values for the interruption frequency and degree of ripple are fed back to the control unit, for comparison at least with the preset values for the upper and the lower limit value, taking into account the maximum values of the X-ray tube voltage and current as well as the permissible values Current flow duration and its correction. Circuit arrangement for carrying out the method according to one of Claims 1 to 9, characterized in that the converter (2; KOV) has at least one measuring input (E1; E2; E3) with voltage measuring set (4) or voltage value converter (U7; U9; U11) for the voltage on the x-ray tube (3) as well a modulation stage (5) is assigned, which forwards the measured value of the tube voltage as a control signal to a modulator (4) which generates the control signal for lowering and re-raising the tube voltage and which uses the control part (2.2) or the grid of the X-ray tube ( 3) upstream electron tube (7) or connected to the grid of the X-ray tube (3 '). Circuit arrangement according to Claim 10, characterized in that the control connections of the thyristors (Th1, Th2) of the control part of the converter (KOV) are connected to the outputs of a phase control (PHS), the current pulses of which ignite the thyristors (Th1, Th2) in relation to the converter oscillation of the desired upper limit value of the voltage for setting its upper limit value are phase-shifted, and that the leading edge control (PHS) has means which, after reaching this upper limit value, blocks the delivery of these ignition pulses until the lower limit value of the voltage is reached, for which purpose the leading edge control (PHS) on the one hand a connection for taking over the phase-correct converter frequency and on the other hand a connection for taking over the voltage present at the anode of the X-ray tube (3), preferably at the charging capacitor (C4) connected downstream of the high-voltage rectifier (HSG), directly or via the voltage he is fed a voltage converter. Circuit arrangement according to Claim 10 or 11, characterized in that the converter (KOV) has a timing element for setting the time interval for lowering and raising the tube voltage again. Circuit arrangement according to claim 12, characterized in that the x-ray tube circuit has a controllable resistor, preferably a triode (R1), connected upstream of the anode of the x-ray tube (3), the grid of which controls the internal resistance and thus controls the time constant of the discharge of the charging capacitor (C4) is connected to the modulator stage (5) , which reduces the negative grid bias when the predeterminable upper limit value of the tube voltage is reached and increases it again when the predeterminable lower limit value is reached. Circuit arrangement according to claim 12, characterized in that the x-ray tube (3 ') has a control grid for controlling its internal resistance and thus the time constant of the discharge of the charging capacitor (C4) via the grid bias, the grid controlling the internal resistance and thus the time constant of the discharge of the charging capacitor is connected to the modulation stage (5), which reduces the negative grid bias when a predetermined upper limit value of the tube voltage is reached and increases it again when a predetermined lower limit value of the tube voltage is reached. Circuit arrangement according to one of Claims 10 to 14 , characterized in that an inductor is provided in the line carrying the charging current of the charging capacitor (C4), for deforming the curve shape of the current emitted by the converter, the inductor and the charging capacitor forming an oscillating circuit with Resonance frequency close to the break frequency. Circuit arrangement according to one of claims 10 to 15, characterized in that the converter or the modulation stage is assigned a processor (U1) for modulation of the interruption frequency and thus the "modulation" level of the X-ray tube voltage. Circuit arrangement according to Claim 16, characterized in that the processor (U1) is assigned a working memory (U5) and a further mass memory (U4), the mass memory containing files in which recording or fluoroscopy parameters are stored which are necessary for a desired recording can be transferred to the working memory for comparison with the set or measured values. Circuit arrangement according to Claim 17, characterized in that the processor can be connected via an external connection bus to a computer PC, on the keyboard of which the inputs are made, and on the monitor, the output of both the specified values and parameters and the acknowledgment for the specifications can be output, the computer / PC preferably having a wet memory designed as a hard disk or as a floppy disk drive, via which the parameters relevant for the recording or the fluoroscopy can be input. Circuit arrangement according to Claim 17 or 18, characterized in that the processor is connected via at least one internal connection bus (ADB) to an internal mass memory (U4) in the form of a permanent memory, the output signals of which, together with those of the main memory (U5), are connected via an internal one Data bus (DDB) are guided, which interact with the digital / analog or digital / analog converter provided in the inputs (E1, E2, E3) and the output (A1).

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