WO2012114217A1

WO2012114217A1 - Mri rf power amplifier with modulated power supply

Info

Publication number: WO2012114217A1
Application number: PCT/IB2012/050525
Authority: WO
Inventors: Ralf GOMMERS; Jan Simons; Christoph Leussler; Daniel Wirtz
Original assignee: Koninklijke Philips Electronics N.V.; Philips Intellectual Property & Standards Gmbh
Priority date: 2011-02-22
Filing date: 2012-02-06
Publication date: 2012-08-30

Abstract

The invention relates to a MR device comprising a RF pulse generator which generates a RF pulse signal, a RF power amplifier (302) which amplifies the RF pulse signal, and a power supply controller (202) which modulates the supply voltage of the RF power amplifier (302). According to one aspect of the invention, the RF power amplifier (302) is a class -E amplifier, which is operated according to a polar modulation scheme. Furthermore, the invention relates to a RF coil module (301) for a MR device comprising at least one RF power amplifier (302) which amplifies a RF pulse signal supplied to an input port of the RF coil module, a coil element (303) connected to the output of the RF power amplifier (302), and a power supply controller (202) which modulates the supply voltage of the RF power amplifier (302).

Description

MRI RF POWER AMPLIFIER WITH MODULATED POWER SUPPLY

Field of the Invention

The invention relates to the field of magnetic resonance (MR) imaging. It concerns a device for MR imaging of at least a portion of the body of a patient placed in an examination volume of the device. The invention also relates to a RF coil module for a MR device.

Background of the Invention

Image- forming MR methods which utilize the interaction between magnetic fields and nuclear spins in order to form two-dimensional or three-dimensional images are widely used nowadays, notably in the field of medical diagnostics, because for the imaging of soft tissue they are superior to other imaging methods in many respects, do not require ionizing radiation and are usually not invasive.

According to the MR method in general, the body of the patient to be examined is arranged in a strong, uniform magnetic field whose direction at the same time defines an axis (normally the z-axis) of the co-ordinate system on which the measurement is based. The magnetic field produces different energy levels for the individual nuclear spins in dependence on the magnetic field strength which can be excited (spin resonance) by application of an electromagnetic alternating field (RF field) of defined frequency (so-called Larmor frequency, or MR frequency). From a macroscopic point of view, the distribution of the individual nuclear spins produces an overall magnetization which can be deflected out of the state of equilibrium by application of an electromagnetic pulse of appropriate frequency (RF pulse) while the magnetic field of the RF pulse extends perpendicularly to the z-axis, so that the magnetization performs a precession about the z-axis. This motion of the magnetization describes a surface of a cone whose angle of aperture is referred to as flip angle. The magnitude of the flip angle is dependent on the strength and the duration of the applied electromagnetic pulse. In the case of a so-called 90° pulse, the spins are deflected from the z axis to the transverse plane (flip angle 90°). The RF pulse is radiated toward the body of the patient via a RF coil arrangement of the MR device. The RF coil arrangement typically surrounds the examination volume in which the body of the patient is placed.

After termination of the RF pulse, the magnetization relaxes back to the original state of equilibrium, in which the magnetization in the z direction is built up again with a first time constant Ti (spin lattice or longitudinal relaxation time), and the magnetization in the direction perpendicular to the z direction relaxes with a second time constant T₂ (spin-spin or transverse relaxation time). The variation of the magnetization can be detected by means of receiving RF coils which are arranged and oriented within the examination volume of the MR device in such a manner that the variation of the magnetization is measured in the direction perpendicular to the z-axis. The decay of the transverse magnetization is accompanied, after application of, for example, a 90° pulse, by a transition of the nuclear spins (induced by local magnetic field inhomogeneities) from an ordered state with the same phase to a state in which all phase angles are uniformly distributed (dephasing). The dephasing can be compensated by means of a refocusing pulse (for example a 180° pulse). This produces an echo signal (spin echo) in the receiving coils.

In order to realize spatial resolution in the body, linear magnetic field gradients extending along the three main axes are superposed on the uniform magnetic field, leading to a linear spatial dependency of the spin resonance frequency. The signal picked up in the receiving coils then contains components of different frequencies which can be associated with different locations in the body. The signal data obtained via the receiving coils corresponds to the spatial frequency domain and is called k-space data. The k-space data usually includes multiple lines acquired with different phase encoding. Each line is digitized by collecting a number of samples. A set of k-space data is converted to a MR image by means of Fourier transformation.

In a typical MR imaging environment, a RF pulse signal is generated by means of a RF pulse generator which can be an analog RF oscillator or a digital RF waveform generator (synthesizer). The RF pulse signal is of low voltage (in the region of 5 volts) and is fed to a

RF power amplifier which drives the RF coil arrangement of the MR device. In this way, the RF pulses radiated toward the body of the patient are generated. Typically, the RF amplifier of the MR device is a so-called class-AB amplifier. It is a drawback of these amplifiers that they are not very efficient and produce a significant amount of heat. The efficiency of a RF power amplifier is defined as the ratio of output power and input power. Poor efficiency results not only in wasted power but also in a large heat production. The heat load on the RF electronics makes the design more expensive, because provision has to be made for an appropriate cooling system, and negatively impacts reliability of the system.

Object and Summary of the Invention

From the foregoing it is readily appreciated that there is a need for an improved MR device. It is hence an object of the invention to provide a MR device having a RF power amplifier with high power efficiency.

In accordance with the invention, a MR device is disclosed which comprises:

a RF pulse generator which generates a RF pulse signal,

a RF power amplifier which amplifies the RF pulse signal, and a power supply controller which modulates the supply voltage of the RF power amplifier.

The modulation of the supply voltage of the RF power amplifier according to the invention significantly improves the power efficiency. The basic principle of the invention can be practically realized by superimposing the envelope signal of the RF pulse signal at the drain of the transistor of the RF power amplifier. In this way, the RF power amplifier is consistently operated in a high efficiency region. The approach of the invention is especially suitable for RF pulse signals having a high peak-to-average ratio like it is typically the case in MR imaging applications.

The power supply controller of the MR device according to the invention is capable of supplying a variable voltage to the transistor of the RF power amplifier at a bandwidth which is high enough to follow the envelope of a MR imaging pulse sequence. An important insight of the invention is that the modulation of the supply voltage of the RF power amplifier is a very suitable approach due to the nature of the waveforms of RF pulse signals in MR imaging applications. A typical RF pulse signal has a slowly varying envelope, which allows using a low-cost power supply controller having a bandwidth of the order of only 10 kHz. The modulation of the power supply of the RF power amplifier is further facilitated by the fact that the required waveform of the supply voltage is known in advance once the MR device is programmed according to the respective imaging task. The control unit of the MR device can therefore perfectly synchronize the supply voltage of the RF power amplifier and the RF pulse signal supplied to the RF power amplifier while keeping the requirements on bandwidth low.

According to a preferred embodiment of the invention, the MR device further comprises a RF pulse delay component which introduces a time delay into the RF pulse signal, such that the input signals of the RF pulse amplifier, namely the delayed RF pulse signal and the modulated supply voltage, are synchronized. It is of high importance that the supply voltage of the RF power amplifier is precisely adapted to the instantaneous power requirements. For a low output power requirement the supply voltage is reduced and for a high output power requirement the supply voltage is increased. In the ideal case, the RF power amplifier is always driven near its maximum output power where it has maximum efficiency. Hence, a good synchronization of the RF pulse signal supplied to the input of the RF pulse amplifier and the modulated supply voltage signal is essential in terms of efficiency.

According to another preferred embodiment of the invention, the RF pulse generator is adapted to pre-distort the RF pulse signal in such a manner that a non-linearity of the RF power amplifier is compensated for. Linearity is generally an important characteristic desired for the RF power amplifier of a MR device. In accordance with the invention, the drain voltage of the transistor of the RF power amplifier is dynamically biased for the purpose of obtaining the highest possible power efficiency. However, the resulting linearity performance of the RF power amplifier may be unacceptable for the stringent performance requirements of MR devices. For example, the spatial selectivity of the MR signal excitation sensitively depends on the linearity of the amplification of the RF pulse signal. According to the invention, the RF pulse generator, which typically comprises a RF oscillator and/or a RF digital synthesizer, is adapted to (digitally) pre-distort the RF pulse signal in order to improve the amplifier linearity without compromising power efficiency. According to yet another preferred embodiment of the invention, the MR device comprises two or more RF power amplifiers, wherein one or more power supply controllers control the gain of each RF power amplifier. It is generally desirable to have a relatively uniform homogeneity of the generated RF field for excitation of magnetic resonance throughout a cross section of the imaged portion of the patient's body. However, as the MR frequency increases with increasing main magnetic field strength, this becomes more difficult due to conductive losses and wavelength effects within the body of the patient. Multi-channel transmit imaging has been accepted as a standard method of operating the RF coils to achieve a relatively uniform RF field within the examination volume. In known multi- channel transmit systems, the amplified RF pulse signal is typically supplied to the RF coil arrangement including a plurality of individual coil elements. The RF coil arrangement may be a so-called birdcage resonator comprising a plurality of rungs as coil elements, which are arranged in parallel to a longitudinal axis of the main magnetic field, wherein the birdcage resonator surrounds the imaged body portion. Alternatively the coil elements may be local coils or array coils that are placed contiguous to the region selected for imaging. In accordance with the invention, the supply voltages of the RF power amplifiers driving the individual coil elements are controlled in order to optimize the total power efficiency.

The power supply controller of the MR device of the invention can be adapted to continuously vary the supply voltage of the at least one RF power amplifier. In this case, the supply voltage of the RF power amplifier closely follows the envelope of the respective RF pulse signal. This results in an optimum power efficiency. Alternatively, the power supply controller can be adapted to switch the supply voltage between two or more discrete voltages. This alternative results in a slightly reduced power efficiency but may simplify implementation.

The RF power amplifier of the MR device of the invention may be a class-A, class-B or class- AB amplifier. A class-A amplifier has very good linearity, but this type of amplifier has a very poor power efficiency. Hence, its power and heat production is relatively high. The modulation of the voltage supply of the RF power amplifier according to the invention is very well suited to improve the power efficiency in the case that a class-A RF power amplifier is employed. Class-B amplifiers have a linearity which is worse than that of class-A amplifiers. However, class-B amplifiers are much more efficient than class-A amplifiers. Class-B amplifiers are rarely used in MR environments because their linearity is reduced to a level that is not desirable for MR imaging applications. Class-AB amplifiers are often the RF power amplifiers of choice in MR devices. These amplifiers are in between class-A and class-B amplifiers. Both power efficiency and linearity are within acceptable limits. The power efficiency of class-AB amplifiers can be significantly improved by the approach of the invention.

Recently, a different kind of RF power amplifiers, known as switching amplifiers or class-E amplifiers has been developed. Class-E amplifiers have a high power efficiency due to the fact that principally a perfect switching operation does not dissipate power. An ideal switch has zero impedance when closed and infinite impedance when open, implying that there is a zero voltage across the switch when it conducts current (on state) and a non-zero voltage across it in the non- conductive state (off state). Consequently, the product of voltage and current (power loss) is zero at any time. Therefore, a class-E amplifier has a theoretical efficiency of 100 %, assuming ideal switching. Although those advantages of class-E amplifiers could be beneficial for a RF power amplifier in a MR device, their extremely high non-linearity makes class-E amplifiers usually undesirable for MR imaging applications. In accordance with a further aspect of the invention, however, linearity can be obtained even in combination with a class-E power amplifier by controlling the gain of the amplifier according to the instantaneous amplitude (envelope) of the RF pulse signal. The gain of the class-E amplifier can be controlled by modulating the supply voltage of its switching transistor. On the one hand, a very high efficiency can be obtained by employing a class-E amplifier according to the invention. On the other hand, the technical requirements with respect to the power supply controller are higher in comparison to the application of a class-A, class-B, or class-AB amplifier. In the case of a class-E amplifier, the bandwidth of the power supply controller should be in the region of 100-500 kHz. Moreover, while a dynamic range of about 10 dB is sufficient for controlling a class-A, class-B, or class-AB amplifier in accordance with the invention, a dynamic range in the region of 50-70 dB would be desirable in order to obtain sufficient linearity from a class-E amplifier.

Preferably, the class-E amplifier comprises a choke connected in series between a supply voltage source and the transistor of the class-E amplifier, wherein an output node of the amplifier is formed between the choke and the transistor. A shunt capacitor is connected in parallel with the transistor. Hence, the class-E amplifier usable in accordance with the invention has a very simple design and can be provided at low cost. Simultaneously, the power efficiency achieved according to the invention approaches 100 %. The RF coil arrangement of the MR device can be connected directly to the output node of the class-E amplifier or via an appropriate impedance matching or transmission-line circuit. Further advantages are good scalability for massive parallel transmission applications (for example RF body coil arrays comprising more than 64 individual coil elements) and good tolerance with respect to component and load variations.

In accordance with yet another preferred embodiment of the invention, the MR device comprises a polar conversion unit converting the RF pulse signal into a DC envelope signal and a phase-modulated RF signal of constant amplitude. Polar modulation is analog to quadrature modulation in the same way as polar coordinates are analogous to Cartesian coordinates. Quadrature modulation makes use of Cartesian coordinates, wherein one axis is called the I (in-phase) axis, while the other axis is called the Q (quadrature) axis. Polar modulation makes use of polar coordinates, i.e. amplitude and phase. The issue of linearity of the RF power amplifier of the MR system can be mitigated by requiring that the input signal of the RF power amplifier is of constant envelope, i.e. contains no amplitude variations. In a polar modulation system, the input signal of the RF power amplifier may vary only in phase.

Amplitude modulation is then accomplished by directly controlling the gain of the RF power amplifier through changing or modulating its supply voltage. Thus, the polar conversion unit of the MR device of the invention allows the use of a highly non-linear class-E amplifier as RF power amplifier.

According to yet another preferred embodiment of the invention, the at least one RF power amplifier of the MR device is integrated into at least one of a plurality of coil elements of the RF coil arrangement. The MR device may comprise individual RF coil modules, into which the RF power amplifier for driving the respective coil element is integrated. Because the supply voltage of the RF power amplifier is modulated according to the invention, its power efficiency is extremely high. Hence, the heat production of the RF power amplifier is low such that individual RF coil modules can be arranged directly on the body of the patient in a suitable configuration without the risk of injury due to heat production of the RF power amplifier. A further advantage is that a good tolerance to load variation can be achieved by using a high number of individual RF coil modules having integrated (class-E) RF power amplifiers.

The invention does not only relate to a MR device but also to a method of MR imaging comprising the steps of: □

subjecting an object to an imaging sequence comprising RF pulses, which are transmitted toward the object via a RF coil arrangement, wherein RF signals are supplied to the RF coil arrangement via a RF power amplifier which amplifies the RF pulse signals, wherein a supply voltage of the RF power amplifier is modulated by means of a power supply controller;□

acquiring MR signals from the object; and

reconstructing a MR image from the acquired MR signals.

Brief Description of the Drawings

The enclosed drawings disclose preferred embodiments of the present invention. It should be understood, however, that the drawings are designed for the purpose of illustration only and not as a definition of the limits of the invention. In the drawings:

Figure 1 schematically shows a MR device according to the invention;

Figure 2 schematically shows the control unit and the transmit unit of the MR device of Figure 1 in more detail;

Figure 3 shows a block diagram illustrating a class-E amplifier as RF power amplifier integrated into a local RF coil module according to the invention; Figure 4 schematically shows a body coil using an array of integrated polar modulated class-E amplifiers;

Figure 5 shows a circuit diagram of a push-pull class-E amplifier usable as RF power amplifier according to the invention.

Detailed Description

With reference to Figure 1, a MR device 101 is shown. The device comprises superconducting or resistive main magnet coils 102 such that a substantially uniform, temporally constant main magnetic field is created along a z-axis through an examination volume.

A magnetic resonance generation and manipulation system applies a series of RF pulses and switched magnetic field gradients to invert or excite nuclear magnetic spins, induce magnetic resonance, refocus magnetic resonance, manipulate magnetic resonance, spatially and otherwise encode the magnetic resonance, saturate spins, and the like to perform MR imaging.

More specifically, a gradient pulse amplifier 103 applies current pulses to selected ones of whole-body gradient coils 104, 105 and 106 along x, y and z-axes of the examination volume. A transmit unit 107 transmits RF pulses or pulse packets to a whole-body volume RF coil 109 to transmit RF pulses into the examination volume.

A typical MR imaging sequence is composed of a packet of RF pulse segments of short duration which taken together with each other and any applied magnetic field gradients achieve a selected manipulation of nuclear magnetic resonance. The RF pulses are used to saturate, excite resonance, invert magnetization, refocus resonance, or manipulate resonance and select a portion of a body 110 positioned in the examination volume. The MR signals are also picked up by the whole-body volume RF coil 109.

For the generation of MR images of limited regions of the body 110 by means of parallel imaging, a set of local array RF coils 1 1 1, 1 12, 113 are placed contiguous to the region selected for imaging. The array coils 111 , 112, 113 can be used to transmit RF pulses toward the body 110 and to receive MR signals induced, for example, by body-coil RF transmissions.

The resultant MR signals are picked up in the depicted embodiment by the array RF coils 111, 112, 113 and demodulated by a receiver 114 preferably including a preamplifier (not shown). The receiver 114 is connected to the RF coils 111, 112 and 113 via send/receive switch 108.

A control unit 115 controls the gradient pulse amplifier 103 and the transmitter 107 to generate any of a plurality of MR imaging sequences, such as echo planar imaging (EPI), echo volume imaging, gradient and spin echo imaging, fast spin echo imaging, and the like. For the selected sequence, the receiver 114 receives a single or a plurality of MR data lines in rapid succession following each RF excitation pulse. A data acquisition system 116 performs analog-to-digital conversion of the received signals and converts each MR data line to a digital format suitable for further processing. In modern MR devices, the data acquisition system 1 16 is a separate computer which is specialized in acquisition of raw image data.

Ultimately, the digital raw image data is reconstructed into an image representation by a reconstruction processor 117 which applies a Fourier transform and other appropriate reconstruction algorithms, such as SENSE or SMASH. The MR image may represent a planar slice through the patient, an array of parallel planar slices, a three-dimensional volume, or the like. The image is then stored in an image memory where it may be accessed for converting slices, projections, or other portions of the image representation into appropriate format for visualization, for example via a video monitor 1 18 which provides a man-readable display of the resultant MR image.

Figure 2 shows the control unit 115 and the transmit unit 107 of the MR device shown in Figure 1 in more detail. The control unit 115 comprises a RF pulse generator 201 which generates a low-voltage RF pulse signal corresponding to the respective MR imaging sequence. The RF pulse generator 201 is typically a RF oscillator or a digital RF synthesizer. A DC power supply controller 202 determines a suitable waveform envelope of the RF pulse signal coming from the RF pulse generator 201. The power supply controller 202 controls a DC power supply 203. The DC power supply 203 generates a supply voltage for a RF power amplifier 204. The power supply controller 202 is adapted to modulate the supply voltage of the RF power amplifier 204 according to the determined envelope of the RF pulse signal generated by the RF pulse generator 201. Moreover, provision is made for a RF pulse delay component 205 which introduces a time delay into the RF pulse signal, such that the RF input signal fed to the RF power amplifier 204 is synchronized with the modulated supply voltage generated by the DC power supply 203. The continuously varying supply voltage generated by the DC power supply 203 is supplied to the drain of the transistor (not depicted) of the RF power amplifier 204. The RF pulse envelope tracking approach illustrated in Figure 2 results in a high power efficiency at any output power, not only near the maximum power for which the RF power amplifier 204 is originally designed. The envelope tracking scheme is suitable in combination with linear amplifiers due to the nature of typical MR imaging waveforms that are characterized by slowly varying envelopes. Hence, a low-cost DC power supply 203 having a bandwidth in the region of 10 kHz is sufficient. In practice, the function of the RF pulse generator 201, the RF pulse delay component 205, and the power supply controller 202 can be integrated into a single digital circuit board. This circuit board has two synchronized outputs, namely the appropriately delayed RF pulse signal to be supplied to the input of the RF power amplifier 204 and the power supply control signal generated by the power supply controller 202. The resulting linearity performance of the system shown in Figure 2 may not be acceptable for the stringent performance requirements of MR imaging. To this end, the RF pulse generator 201 is adapted to pre-distort the RF pulse signal in such a manner that a non- linearity of the system is compensated for. In a section 206, the control unit 115 comprises digital circuitry which is used for other control purposes, such as, for example, the control of the magnetic field gradients.

Figure 3 shows a local RF coil module 301 incorporating a polar modulated class-E amplifier 302 as RF power amplifier. The output of the RF power amplifier 302 is connected to a coil element 303 which may be, for example, a conventional loop antenna. The RF pulse signal generated by the RF pulse signal generator (not depicted in Figure 3) is supplied to the input of a processor 304 which converts the RF pulse signal into quadrature signals I (in phase) and Q (quadrature). The output signals of the processor 304 are supplied to the inputs of a polar conversion unit 305 which converts the signals I and Q into signals p (amplitude) and Θ (phase). The amplitude signal is further supplied to the power supply controller 202 which comprises a D/A converter. The power supply controller 202 converts the digital amplitude signal p directly into a supply voltage of the switching transistor of the class-E amplifier 302. Further, provision is made for a direct digital synthesizer (DDS) 306 which converts the digital phase signal Θ into a constant amplitude phase modulated RF signal supplied to the input of the class-E amplifier 302. The RF coil module 301 is characterized by a good performance in terms of power efficiency and linearity. The heat dissipation is low so that the RF coil module 301 can be placed directly on the body of the examined patient without the risk of injury. The class-E amplifier 302 itself may have a simple construction and can be provided at low cost.

As illustrated in Figure 4, the system of the invention is advantageously scalable for parallel transmit applications (such as, for example, RF shimming or transmit SENSE). Figure 4 shows a body coil 109 comprising a plurality of coil elements 402 which may be, for example, the longitudinal rungs of a birdcage-type body coil. Each coil element 402 has its own class-E RF power amplifier in combination with the necessary driving circuitry as explained above with reference to Figure 3. A high number of coil elements 402 and the combination with individual RF power amplifiers results in a good tolerance with respect to load variations of the body coil 109.

Figure 5 shows a circuit diagram of a push-pull class-E amplifier usable as RF power amplifier according to the invention. Figure 5 illustrates the simple and, hence, low-cost amplifier design, which comprises two anti-parallel transistors 502 which operate as switches. Moreover, provision is made for chokes 503 via which the supply voltage at connectors 504 is supplied to the transistors 502. The input connector for the phase-modulated constant- amplitude RF signal to be amplified is designated by reference number 505 in Figure 5. The chokes 503 are connected in series between the supply voltage source and the respective transistor 502, wherein output nodes 506 are formed between the chokes 503 and the transistors 502. Shunt capacitors 507 are connected in parallel with the transistors 502. The output load 508 is the RF coil arrangement of the MR device. In the depicted embodiment, the RF coil arrangement 508 is connected to the amplifier via a transformer 509. In accordance with the invention, the modulated supply voltage is provided to connectors 504 in order to control the gain of the RF power amplifier according to the envelope of the RF pulse signal.

Claims

1. A MR device (101) comprising:

a RF pulse generator (201) which generates a RF pulse signal, a RF power amplifier (204, 302) which amplifies the RF pulse signal, and a power supply controller (202) which modulates the supply voltage of the RF power amplifier (204, 302).

2. The MR device of claim 1, wherein the power supply controller (202) is adapted to modulate the supply voltage of the RF power amplifier (204, 302) according to an envelope of the RF pulse signal.

3. The MR device of claim 2, further comprising a RF pulse delay component (205) which introduces a time delay into the RF pulse signal to obtain a delayed RF pulse signal synchronized with the modulated supply voltage.

4. The MR device of any one of claims 1 to 3, wherein the RF pulse generator (201) is adapted to pre-distort the RF pulse signal in such a manner that a non-linearity of the RF power amplifier (204, 302) is compensated for.

5. The MR device of any one of claims 1 to 4, comprising two or more RF power amplifiers (204, 302), wherein one or more power supply controllers (202) control the gain of each RF power amplifier (204, 302).

6. The MR device of any one of claims 1 to 5, wherein the power supply controller (202) is adapted to continuously vary the supply voltage of the at least one RF power amplifier (204, 302) or to switch the supply voltage between two or more discrete voltages.

7. The MR device of any one of claims 1 to 6, wherein the RF power amplifier (204, 302) is a class- A, class-B, or class-AB amplifier.

8. The MR device of any one of claims 1 to 6, wherein the RF power amplifier (204, 302) is a class-E amplifier comprising a transistor (502), wherein the power supply controller (202) controls the gain of the RF power amplifier (204, 302) through modulating the supply voltage of the transistor (502).

9. The MR device of claim 8, wherein the class-E amplifier comprises a choke (503) connected in series between a supply voltage source (504) and the transistor (502), an amplifier output node (506) being formed between the choke (503) and the transistor (502), and a shunt capacitor (507) being connected in parallel with the transistor (502).

10. The MR device of claim 8 or 9, comprising a polar conversion unit (305) converting the RF pulse signal into a DC envelope signal and a phase-modulated RF signal of constant amplitude.

1 1. The MR device of claim 10, wherein the power supply controller (202) controls the gain of the RF power amplifier (204, 302) according to the DC envelope signal, while the class-E amplifier amplifies the constant-amplitude phase-modulated RF signal.

12. The MR device of any one of claims 1 to 11, further comprising: □

a main magnet (102) for generating a uniform, steady magnetic field within an examination volume along a main field axis, □

- a number of gradient coils (104, 105, 106) for generating switched magnetic field gradients in different spatial directions within the examination volume, □

at least one RF coil arrangement (109) for generating RF pulses within the examination volume and/or for receiving MR signals from a body (110) of a patient positioned in the examination volume,

- a transmit unit (107) comprising one or more RF power amplifiers (204, 302) for transmitting the amplified RF signal to the RF coil arrangement (109), a control unit (115) for controlling the temporal succession of RF pulses and switched magnetic field gradients, □

a reconstruction unit (117), and □

a visualization unit (118).

13. The MR device of claim 12, wherein the transmit unit (107) has two or more transmit channels, each transmit channel comprising at least one RF power amplifier (204, 302), wherein each transmit channel is connected to one of a plurality of coil elements (303, 402) of the RF coil arrangement (109).

14. The MR device of claim 12 or 13, wherein at least one RF power amplifier (204, 302) is integrated into at least one of a plurality of coil elements of the RF coil arrangement (109).

15. A RF coil module (301) for a MR device (101) comprising:

at least one RF power amplifier (302) which amplifies a RF pulse signal supplied to an input port of the RF coil module (301),

a coil element (303) connected to the output of the RF power amplifier (302), and

a power supply controller (202) which modulates the supply voltage of the RF power amplifier (302).

16. A method of MR imaging comprising the steps of: □

subjecting an object (110) to an imaging sequence comprising RF pulses, which are transmitted toward the object via a RF coil arrangement (109), wherein RF signals are supplied to the RF coil arrangement (109) via a RF power amplifier (204, 302) which amplifies the RF pulse signals, wherein a supply voltage of the RF power amplifier (204, 302) is modulated by means of a power supply controller (202);□

acquiring MR signals from the object; and

reconstructing a MR image from the acquired MR signals.