RU2575050C2 - State of space controller with feedback in area of digital data for source of power supply of gradient coil for magnet-resonance imaging - Google Patents

Abstract

FIELD: measurement equipment.

SUBSTANCE: invention relates to a state of space controller with a feedback. The method of operation of a system of magnet-resonance imaging (MRI) with a gradient magnet for the generation of gradient magnetic fields by means of excitation of currents of a gradient coil in gradient coils of the system with the gradient magnet, besides, currents of the gradient coil are controlled with the help of a controller, which is implemented in the form of a state of space controller with a feedback in the field of digital data and which generates a digital output value, with the help of which they generate width-pulse-modulated voltage to excite currents of the gradient coil via the gradient coils with a preset wave form, besides, the control deviation is integrated, which is generated in the form of a difference between the current of the gradient coil and the reference current, the digital outlet value (u) is generated by the integrated control deviation, and the digital outlet value (u) is sent back and delayed at least by one, two and three cycles of synchronisation, and then distract from the digital outlet value (u).

EFFECT: increased quality of imaging in the MRI system.

4 cl, 9 dwg

Description

FIELD OF TECHNOLOGY

The invention relates to a method and a digital amplifier system for operating a gradient magnet system for MRI (magnetic resonance imaging) for generating gradient magnetic fields, in particular for spatial coding of MR relaxation signals (magnetic resonance).

BACKGROUND OF THE INVENTION

In an MRI or MR scanner, the object under investigation, usually the patient, is exposed to a uniform main magnetic field (field B ₀ ) so that the magnetic moments of the nuclei inside the object under investigation tend to rotate around the axis of the applied field B ₀ (Larmor precession) with a certain resulting magnetization of all nuclei parallel to the field B ₀ . The precession rate is called the Larmor frequency, which depends on the specific physical characteristics of the involved nuclei, namely their gyromagnetic ratio and the applied field strength B ₀ . The gyromagnetic ratio is the ratio between the magnetic moment and the spin of the nucleus.

By transmitting a radio frequency (RF) exciting pulse (field B ₁ ), which is orthogonal to the field B ₀ , which the RF transmitting antenna generates, and when the Larmor frequency of the nuclei of interest coincides, the nuclear spins are excited and phase matched, and their resultant deviation magnetization from the direction of the field B _{0 is} obtained in order to generate a transverse component depending on the longitudinal component of the resulting magnetization.

After the completion of the RF excitation pulse, relaxation processes of the longitudinal and transverse components of the resulting magnetization begin until the resulting magnetization returns to its equilibrium state. MR relaxation signals that are emitted by the transverse relaxation process are detected by the MR / RF receiving antenna.

The Fourier transform of the received MR relaxation signals, which are amplitude signals in time, into frequency-based MR spectrum signals and processing to generate MR images of the nuclei of interest inside the object under study are carried out. In order to achieve spatial selection of the slice or volume inside the object under study and spatial coding of the received MR relaxation signals originating from the slice or volume of interest, gradient magnetic fields are applied to the field B ₀ , which have the same direction as the field B ₀ but have gradients in the orthogonal directions x, y and z. Due to the fact that the Larmor frequency depends on the magnetic field strength that is superimposed on the nuclei, the Larmor frequency of the nuclei correspondingly decreases along with the decreasing gradient (and vice versa) of the general field B ₀ obtained as a result of superposition, so that by appropriate adjustment of the transmitted frequency RF of the exciting pulse (and by appropriate adjustment of the resonant frequency MR / RF of the receiving antenna) and by appropriate control of the gradient magnetic fields can be achieved selection of nuclei within a slice at a specific location along each gradient in the x, y, and z direction, and thereby, generally, within a specific voxel of an object.

To generate gradient magnetic fields, a gradient magnet system is provided that contains a plurality of gradient magnets in the form of coils (gradient coils), which typically operates using a gradient amplifier system to generate an electric current to power the gradient coils. Typically, such gradient coil currents have a specific waveform that the gradient amplifier system must produce very accurately. The waveform is, for example, a trapezoidal pulse with a pulse duration of, for example, approximately 40 ms, a rise and fall time of pulses, for example, approximately 0.2 ms each, and an amplitude of the order of several hundred and up to approximately 1000 A. These current pulses follow precisely control with a deviation of only a few mA or less in order to ensure the generation of MRI images with high quality and high spatial resolution and accuracy.

WO 2010/004492 discloses a digital amplifier with direct and feedback control for generating and controlling an electrical output power for operating such a gradient magnet system. Basically, such a digital amplifier comprises an input for receiving a digital input signal representing the desired current waveform of the analog output signal of the amplifier, and the gradient magnet system is driven by this analog output signal. The amplifier comprises a direct and feedback controller in order to detect and compensate for an error signal between the input signal and the output signal. A direct-coupled controller reads the input signal and predicts the output signal as accurately as possible based on the model of the system. Then, the predicted output signal is subtracted in the analog data region from the measured analog output signal, and the resulting analog power differential current is converted into a digital data region to provide a feedback signal, which is then combined with a direct signal. As a result, a digital representation of the determined output signal is obtained, which is then subtracted from the desired digital input signal so that the error signal, which is supplied to the digital controller to provide a suitable control signal to the modulator, leads to the opposite of the error signal. The modulator converts the control signal into a PWM signal, which is fed to a power converter that generates an analog output signal in order to excite the gradient magnet system.

The digital control scheme for PWM energy conversion was disclosed in a report at the conference "Advanced digital control scheme of two-paralleled bridge type current tracking power conversion amplifier for magnetic-resonance imaging" by S. Watanabe and M. Nakaoka on May 26, 1997.

SUMMARY OF THE INVENTION

In connection with the above requirements for the accuracy of current pulses for operating a gradient magnet system, the known gradient amplifier system is controlled more or less in the same way, since digital control inherently suffers from digital sampling noise and certain digitizing effects. In particular, it has been shown that the limited resolution of analog-to-digital converters and PWM modulators, as well as the limited processing speed of digital controllers, can cause serious deviations and reproducibility problems when the desired gradient magnet current is generated.

In general, one objective underlying the invention is to provide a digital amplifier method and system for operating an MRI gradient magnet system based on fully digital control without causing the above or other problems, but at the same time meeting the requirements for accuracy, as indicated above, is much more complete than when using known solutions.

Another objective underlying the invention is to provide a digital amplifier method and system for operating a gradient magnet MRI system with high accuracy and at a relatively low cost and based on the relatively cheap and standard serial existing low-power components.

At least one of these problems is solved by the method of claim 1 and the digital amplifier system of claim 2.

The control principle, as provided by the invention, has a very high suppression of interference so that the aforementioned effects of digitization are suppressed to the maximum and, therefore, standard digital components, such as serial, existing analog-to-digital converters, can be used to implement a digital controller.

In addition, the invention makes it possible to implement a digital amplifier method and system for operating an MRI gradient magnet system based on the principles of high-precision control and associated design algorithms such that the above requirements can also be fulfilled for fully digital control.

Another advantage of the solution in accordance with the invention is that for the proposed method of managing the state space, direct control is not required in order to increase the bandwidth. Therefore, the principle of the state space controller in accordance with the invention is more resistant to variations in system parameters and has higher interference suppression than the principle of a standard controller, such as a PID controller.

All this makes it possible to use a fully digital controller and, through this, a fully digital amplifier system to operate the MRI gradient magnet system using cost-effective standard digital components. In addition, the above leads to improved reproducibility of the output current, which leads to higher image quality in the MRI system.

The dependent claims disclose advantageous embodiments of the invention.

It should be appreciated that the features of the invention can be combined in any combination without departing from the scope of the invention as defined by the accompanying claims.

Additional details, features and advantages of the invention are apparent from the following description of preferred and exemplary embodiments of the invention, which is given with reference to the figures.

BRIEF DESCRIPTION OF THE DRAWINGS

In FIG. 1 shows the main components of an MR imaging system or scanner;

in FIG. 2 is a schematic view of a gradient magnet MRI system;

in FIG. 3 shows a general topology of an amplifier system for operating a gradient coil of an MRI system with a gradient magnet;

in FIG. 4 is a schematic architecture of one embodiment of a digitally controlled gradient amplifier system in accordance with the invention;

in FIG. 5 is a schematic functional structure of one embodiment of a digital controller for a gradient amplifier system in accordance with the invention;

in FIG. 6 shows the output of the controller of FIG. 5 in response to the first input signal;

in FIG. 7 shows the output of the controller of FIG. 5 in response to a second input signal;

in FIG. 8 is an enlarged section of the current waveforms of FIG. 6; and

in FIG. 9 shows the transient responses of all system states for the controller structure according to FIG. 5.

DESCRIPTION OF PREFERRED EMBODIMENTS

In FIG. 1 shows the essential components of a magnetic resonance imaging system or magnetic resonance scanner, including the test volume, in which the object to be visualized is placed. The system comprises a main magnet system 1 for generating a substantially uniform and constant main magnetic field B ₀ in the z direction to align the nuclear spins in the object to be visualized. An RF / MR transmitter / receiver antenna arrangement 2 is provided for transmitting RF signals for generating an RF magnetic variable field B ₁ for exciting nuclear magnetic resonance and for receiving subsequent MR relaxation signals from the associated nuclei of the object to be imaged. For spatial selection and spatial coding of the received MR relaxation signals emanating from excited nuclei, the system also comprises a gradient magnet system using a plurality of gradient magnetic field coils 3, 4, 5, with which gradient magnetic fields are generated in the orthogonal directions x, y and z, respectively, as described above.

Additionally, the magnetic resonance imaging system or magnetic resonance scanner contains a control unit 10, which during RF transmission controls the RF transmitter unit 12 to generate RF signals that are supplied via the RF transmitting / receiving circuit 13 to the arrangement 2 of the RF antenna. The received MR relaxation signals are fed through the RF transmitting / receiving circuit 13 to the amplifier and demodulator unit 14 and processed by the processing unit 15 to generate an MR image of the object under study on the display unit 16. In addition, while receiving the MR signal, the control unit 10 generates control signals for the amplifier system 11 for generating electrical output power, in particular in the form of output currents for operating each of the coils of the gradient magnetic field 3, 4, 5 in order to select the desired the slice or volume inside the test object to be visualized on the display unit 16, as is generally known.

The above and the following principles and circumstances are applicable both in the case of an axial (horizontal) MRI system in which a patient or other test object is guided axially through a cylindrical or tubular test space, and in the case of a vertical (open) system having a test space between the upper and lower end of the C-frame frame structure.

In FIG. 2 shows in more detail a schematic view of the arrangement of coils 3, 4, 5 of the gradient magnetic field of the MRI system with a gradient magnet for a magnetic resonance imaging system or magnetic resonance scanner.

To generate a position-dependent or gradient magnetic field in the x direction, the first coils 3 of the gradient magnetic field are provided, one along both sides of the test object 7. To generate a position-dependent or gradient magnetic field in the y direction, second coils 4 of the gradient magnetic field are provided, one above and below the test object 7. Finally, third gradient magnetic field coils 5 are provided that surround the test object and are displaced in the z direction to generate yaschego on the position or the gradient magnetic field in the z direction. The RF / MR arrangement 2 of the transmitting / receiving antenna is also schematically shown in FIG. 2.

In FIG. Figure 3 shows an exemplary general topology of the main stage of the amplifier of the gradient amplifier system for each of the gradient coils (GC) 3, 4, 5 to work from the first to the third of the MRI system with a gradient magnet by passing the current of the gradient coil i _GC through it (output filter of the gradient coil not shown in this FIG.). The amplifier stage preferably contains three (or more) full bridges that are connected in series, namely the first, second and third inverter A, B, C, to generate a first output voltage U _A , a second output voltage U _B and a third output voltage U _C, respectively . Each full bridge contains four ps power switches, each of which can be provided by means of several parallel-connected IGBTs, the preference of which is given due to economic reasons. In addition, it is preferable to use the known M modules with half or full IGBT bridge, containing from two to four IGBTs and even driver electronics for controlling the shutter in order to further reduce cost.

The power supply voltage U _S (for example, 600 V) of each full bridge has a floating potential. Alternatively, you can use the voltage of only one power source, which loads the capacitors C of all full bridges by using the intelligent control approach.

Each full bridge or inverter A, B, C can generate three levels of output voltage (+ U _S , 0, -U _S ) depending on the position of the ps power switch. Thus, the series connection of three full bridges can deliver seven different voltage levels. Due to this topology, two significant advantages are obtained. On the one hand, high output voltage can be generated by low voltage power switches ps so that the costs associated with ps power switches can be reduced. On the other hand, due to the high number of discrete voltage levels compared with an amplifier with one full bridge, the current fluctuations will be less, reducing the requirements on the output filter of the gradient coil and increasing current reproducibility.

In order to fulfill the above requirements of an amplifier system for operating a gradient magnet system, it was found that closed loop control for the main stages of the amplifier and the output current of the amplifier system for gradient coils is necessary and that, in particular, full digital control is preferred. In FIG. 4 is a schematic architecture of an embodiment of such a digitally controlled gradient amplifier system in accordance with the invention for each of one of the first to third gradient coils (GC) 3, 4, 5.

It contains an amplifier stage, for example, three (or more) inverters A, B, C, as shown in FIG. 3, which are reflected in FIG. 4 in a simplified representation in the form of an inverter Inv block. The output filter F of the inverter is shown schematically in the output filter F of the gradient coil, which has two series inductors Lf and one parallel capacitor Cf. The gradient coil GC itself (one of the first to third) is shown in the form of its equivalent circuit, namely the series connection of the inductor L and the ohmic resistance R. In addition, the gradient amplifier system contains a first current sensor for sensing the output current i _c of the inverter Inv block , a second current sensor for sensing the current of the gradient coil i _GC that flows through the gradient coil GC, and a voltage sensor for sensing the voltage of the gradient coil u _C on the gradient coil GC.

The output current i _C of the inverter Inv unit (inverter current), the voltage of the gradient coil u _C and the current of the gradient coil i _GC provide low noise levels to the inputs of the first, second and third operational amplifiers OP1, OP2, OP3, respectively, for amplification in the analog domain data. The output terminals of operational amplifiers OP1, OP2, OP3 are connected to the inputs of the high-resolution analog-to-digital converter A / D, which also receives the reference current I _ref from the reference current source Ref.

The digitized output values of the A / D analog-to-digital converter provide the input of the Con digital controller, which operates on the basis of digital hardware, such as a digital signal processor (DSP) or a user programmable gate array (FPGA), entirely in the field of digital data.

The output control signals u of the Con controller are provided to the PWM modulator PWM input to convert the output control signals u of the Con controller to suitable ss switch signals. Finally, these switch signals ss are provided to the inverter unit Inv, for example, in FIG. 3, to the power switches ps of the inverters A, B, C, for switching the voltage of the power supply U _S by the inverters A, B, C so as to generate PWM voltage at the output of the inverter Inv block so that the desired gradient current i _GC having a specific waveform and amplitude, as described above, pass through a connected gradient coil GC.

The advantages of such a digital control system and, in particular, the all-digital Con controller, are higher flexibility, shorter development time and, in particular, the fact that improved ways to control the new system architecture are possible. The latter advantage is especially important since it will be used to reduce the cost of the entire gradient amplifier system, which is very important for such new systems. Cost reduction can be realized due to the fact that cheaper electronic power management components can be used, the disadvantages of which can be compensated for by more advanced digital control algorithms.

The main problems encountered when implementing such a fully digital control system are the transition from the field of analog data to the field of digital data and the transition from the field of digital data to the field of analog data. The main causes of these problems are the limited resolution in value of the A / D A / D converter, the limited processing speed (time resolution) of the Con digital controller and the limited time resolution of the PWM PWM modulator, which together can lead to the desired current gradient coil i _GC cannot be reproduced with the required accuracy, as described above.

In general, two different controller principles can be used to implement such a control system in the field of digital data, namely, a PID controller and a state space controller. It has been revealed that the principle of the state space controller has many advantages over the principle of the PID controller in terms of reducing the above limited resolutions and reliability, in particular, in relation to changing the response to control if system parameters deviate from their nominal values. The control structural and functional blocks of the associated PI controller of the state space are schematically indicated in FIG. 5.

The main components are the integrating part G ₁ and the U-managed system G _PSYS . The _PSYS P-control system G contains a delay compensator / stabilizer G _STAB and a power plant G _PLANT .

The delay compensator / stabilizer GSTAB contains blocks z ^-1 , z ^-2 , z ^-3 delay, which reflect the computational delay for one, two or three synchronization cycles, respectively. In addition, the delay compensator / stabilizer GSTAB contains control coefficients K _U1 , K _U2 , K _U3 for the computational delay for one, two or three synchronization cycles of the output control signal u of the controller Con, respectively, and control coefficients K _UC and K _IC for the voltage of the gradient coil u _C and the output current i _C (see FIG. 4) of the inverter Inv block, respectively. Finally, the delay compensator / stabilizer GSTAB contains the transfer functions G _UC and G _{IC of the} output filter F of the gradient coil (see Fig. 4) for the voltage of the gradient coil u _C and the output current i _C of the inverter Inv block, respectively, and the control coefficient K _GC for the gradient current Coil i _GC .

The power plant G _PLANT contains the transfer function G _{GC in} series with a regular delay circuit with 20 µs of the measured value of the current of the gradient coil i _{GC of the} A / D analog-to-digital converter plus additional filtering.

The measured values of u _C and i _{C are} characterized by a delay that is equal to the computational delay. Thus, there is no additional delay circuit for these measured values.

In the integrating part G _I , a control deviation is integrated, which is the difference between the reference current I _ref and the current of the gradient coil i _GC . In addition, the current of the gradient coil ioc supplied through the control coefficient K _GC is subtracted from the output signal of the integrating part Gi. The result of the current of the gradient coil i _GC is the output signal of the power plant G _PLANT , into which the output value u of the controller Con, having a delay of one synchronization cycle z ^-1 , is input.

The digital output value u of the controller Con having a delay of one synchronization cycle z ^-1 is also fed back through the transfer function G _UC and the control coefficient K _UC and through the transfer function G _ic and the control coefficient K _ic . Finally, the output value u of the controller Con is delayed by one, two or three synchronization cycles z ^-1 , z ^-2 , z ^-3 , and each delayed output value u is also fed back through the associated control coefficients K _U1 , K _U2 , K _U3, respectively .

The parameters of such a state-space feedback controller shown in FIG. 5, in particular, the above control coefficients and transfer functions, can be developed using known methods, for example, a method for placing poles or others.

Preferably, the poles are selected so that they are real, to ensure that there is no overshoot of the response to closed loop control. Poles with absolute values that are less than one will guarantee the stability of the controller. The closer the poles to zero, the faster the response to control. For example, poles of 0.15 will usually lead to a sufficiently large controller bandwidth, while also guaranteeing maximum resistance to deviations of the controller parameters from their nominal values.

In FIG. Figure 6 shows the simulation of the time interval for the controller of the resulting state space for the step input reference current I _ref with an amplitude of 5 A. The response to the control is very fast. Reaching 63% of the final value of the current i _{GC of the} gradient coil amplifier takes approximately 40 μs (starting from the first control response). There is a slight overshoot / under-regulation, which is the result of the approximations performed to develop the controller. They represent the averaging of the measured value modeled with a net delay and an integer delay modeled with a weighted sum of two values with an integer delay.

However, this is not critical, since overshoot / under-regulation is even lower if a practically significant, ramp current I _ref is supplied as an input to the controller, as shown in FIG. 7, for a linearly varying current of 500 A. It is hardly possible to note a not ideal response to control (current i _{GC of the} gradient coil amplifier).

In FIG. 8 is an enlarged critical region of the control response of FIG. 6. It indicates that the maximum control deviation during overshoot / under-regulation is only approximately 1.4 A (0.3%). The corresponding steady state characteristic i _GC fully complies with the requirement S of the gradient amplifier for installation.

Finally, the transient response of all system states, including the current of the gradient coil i _GC and the states u _C and i _{C of the} output filter F, are shown in FIG. 9 for a linearly varying current of 500 A. In these examples, the voltage u _C on the capacitor C _f (see Fig. 4) is equivalent to the PWM voltage after the low-pass filter. The maximum voltage u _C is approximately ± 1300 V and is necessary in order to change the current of the i _GC gradient coil from 0 to approximately 500 A in approximately 0.2 ms. Due to the almost perfect attenuation of the LC filter controller F, the output of the Inv unit inverter cannot oscillate with a resonant frequency.

Although the invention is illustrated and described in detail in the drawings and in the above description, such illustrations and description should be considered as illustrative or exemplary, and not limiting, and the invention is not limited to the disclosed embodiments. Variations of embodiments of the invention described above, for example, with respect to the antenna elements themselves, their shapes and the number and locations relative to each other, to achieve the result explained above in the form of at least essentially identical fields of view for nuclei having essentially various gyromagnetic relationships are possible for a person skilled in the art without departing from the basic principle of the invention as defined by the accompanying claims.

Variations of the disclosed embodiments may be understood and carried out by those skilled in the art in practicing the invention described in the application, having studied the drawings, disclosure and appended claims. In the claims, the word “contains” does not exclude other elements or steps, and the singular form does not exclude the plural. One block can fulfill the functions of several elements listed in the claims. The mere fact that certain measures are listed in mutually different dependent claims does not indicate that a combination of these measures cannot be used to advantage. Any reference position in the claims should not be construed as limiting the scope of the claims.

Claims (4)

1. The method of operation of a magnetic resonance imaging (MRI) system with a gradient magnet to generate gradient magnetic fields by exciting the gradient coil currents in the gradient coils of a gradient magnet system, the gradient coil currents being controlled using a controller that is implemented in the form of a state space controller with feedback in the field of digital data and which generates a digital output value by which a pulse width modulated voltage is generated for gradient coil excitation currents through the gradient coils with a predetermined wave shape, and integrating the control deviation, which is generated in the form of the difference between the gradient coil current and a reference current,
generate a digital output value (u) by the integrated control deviation and
fed back and delayed the digital output value (u) for at least one, two and three synchronization cycles and then subtracted from the digital output value (u). 2. A digital amplifier system for operating a magnetic resonance imaging (MRI) system with a gradient magnet to generate gradient magnetic fields by exciting the gradient coil currents in gradient coils (GC) of a gradient magnet system, the amplifier system comprising a digital controller (Con) for controlling currents of the gradient coil, and the controller (Con) is implemented in the form of a controller
state spaces with feedback in the field of digital data to generate a digital output value (u), with which a pulse-width-modulated voltage is generated to excite the currents of the gradient coil in gradient coils with a predetermined waveform, and the controller (Con) contains an integrating part (G _I ) for integrating a control deviation that is generated in the form of a difference between the current of the gradient coil and the reference current,
moreover, the controller (Con) contains a delay equalizer (G _STAB ) for generating a digital output value (u) by the integrated control deviation and
moreover, the delay compensator (G _STAB ) contains a feedback loop through which the digital output value (u) is fed back and delayed by at least one, two and three synchronization cycles and then subtracted from the digital output value (u). 3. The digital amplifier system of claim 2, wherein the delay equalizer (G _STAB ) comprises a feedback loop through which the digital output value (u) is fed back through the transfer function (G _UC , G _IC ) of the output filter (F) of the gradient coil for the voltage of the gradient coil (u _C ) and the output current (i _C ) and then subtracted from the digital output value (u). 4. The digital amplifier system of claim 2, wherein the controller (Con) comprises a transfer function (G _GC ) serial with a delay circuit for a measured value of the current of the gradient coil (i _GC ) of an analog-to-digital converter (A / D), by which gradient coil current is generated by a digital output value (u).

Images