Patent Description:
This section is intended to introduce the reader to certain aspects of art that may be related to aspects of the present disclosure, which are described and/or claimed below. Accordingly, it may be understood that these statements are to be read in this light, and not as admissions of prior art.

Magnetic resonance imaging (MRI) is an imaging modality that is often used to generate images (e.g., spatial maps) based on the distribution of molecules in a subject. Generally, an MRI image of a subject (e.g., a patient) is produced by measuring properties of the gyromagnetic materials of the subject, such as hydrogen nuclei. These properties are usually obtained by measuring emissions from the gyromagnetic materials within the subject in response to an excitation from applied magnetic fields. The magnetic fields used for excitation generally include a strong primary magnetic field, magnetic field gradients, and radiofrequency (RF) magnetic field excitation pulses. Of note, the magnetic field gradients may be used by the MRI system to provide spatial encoding to the acquired data. To that end, the gradient coils may be used to generate spatial gradients of magnetic fields such that spatial coordinates of the subject may be associated with a locally encoded magnetic field value. As a result of this magnetic field-based encoding, the emission from the gyromagnetic materials may contain information that may be used to inform the spatial origin of a particular emission during image reconstruction.

The magnetic field gradients may be obtained by driving specific currents to the gradient coils. More specifically, these magnetic field gradients may be controlled by adjustment of the currents in the gradient coils that are responsible for generating magnetic fields. Magnetic coils may be associated with each one of spatial axes, the x-axis, the y-axis, and the z-axis, and the current in each of the axis coils may be independently controlled during the course of the data acquisition, to obtain flexible three-dimensional slicing of the images. To generate the currents in the magnetic coils, magnetic coil drivers may be used. A magnetic driver may include a gradient power supply and a gradient amplifier (e.g., a current amplifier) that can induce currents in the magnetic coil. Magnetic coil drivers may demand very accurate currents for precise, high-resolution spatial encoding. Moreover, the currents and voltages required to drive the current in the magnetic coils can be very large and varying. As a result, the design for the power supplies in these systems may be particularly challenging due to the large current changes and the large voltages normally used by the amplifiers.

<CIT> relates to a power supply that may include a power transformer having a primary winding and a secondary winding, one end of the secondary winding connected to ground, and a shielded isolation transformer having a third winding, a fourth winding, and a shield, wherein the third winding is connected to the secondary winding and the shield is connected to ground. Primary side circuits may receive input power and generate a primary AC signal to drive the primary winding. Secondary side circuits may convert a secondary AC signal output from the fourth winding into a DC output.

<CIT> relates to the provision of a power amplifier, a power source device and a magnetic resonance imaging apparatus. The power amplifier comprises an amplifying circuit connected between a power source and a load and configured to provide amplified output power to the load according to a received input signal, and a bypass circuit connected between the amplifying circuit and the load in parallel and configured to be enabled during the periods of time sections between pulse durations of the output power. In this way, the bypass circuit and the load form a ring current loop. Therefore, the reliability of the power amplifier is improved.

"<NPL> proposes a modification in the existing power architecture of the MRI system to reduce the cost and footprint of a Magnetic Resonant Imaging (MRI) scanner setup in a hospital. A high frequency alternative for the frontend bulky low frequency transformer has been proposed. The implementation challenges and workaround solutions for the High Frequency Power Distribution Unit (HFPDU) has been explored and the experimental setup and results are discussed in the paper.

Certain embodiments commensurate in scope with the invention which is defined in the appended claims are summarized below. These embodiments are not intended to limit the scope of the claimed invention, but rather these embodiments are intended only to provide a brief summary of possible forms of the invention. Indeed, the invention may encompass a variety of forms that may be similar to or different from the embodiments set forth below. In any case, the invention is defined only by the appended claims.

To the extent that any example herein does not fall within the scope of the claims, that example is outside the scope of the claimed invention. In an embodiment, a magnetic resonance imaging (MRI) system is described. The MRI system may include multiple gradient coils and a gradient driver that drives the gradient coils. The gradient driver may include multiple gradient amplifiers and each gradient amplifier may independently control the current in the gradient coil by an electric
coupling. Moreover, each gradient amplifier employs a single semiconductor bridge to perform the control. The gradient driver also includes a power distribution unit (PDU). The PDU may receive an alternating current (AC) power signal from the main power source of the MRI system and provide a direct current (DC) power signal to the gradient amplifiers via a DC bus.

In an example , a High Frequency Power Distribution Unit (HFPDU) is described. The HFPDU may include a power distribution unit and a power supply. The power distribution unit may have a line filter and a first rectifier. The power distribution unit may receive a three-phase AC power signal from a power source and generates an intermediate DC signal. The power supply includes a semiconductor bridge, a high-frequency transformer, a high speed rectifier, and a filter. The semiconductor bridge may receive the intermediate DC signal of the rectifier and generate a high frequency AC power signal. The high frequency AC power signal may be provided to the high-frequency transformer, which may provide galvanic insulation. Moreover, a shield of the transformer may be coupled to a safety ground. The AC power from the transformer may be provided to the high-speed rectifier to produce an output DC power signal.

In a further example, a gradient amplifier is described. The gradient amplifier may drive a gradient coil in an MRI system. The gradient amplifier may have a single semiconductor bridge having a first branch that includes a first and a second switch. A first terminal of the gradient coil may be coupled to a midpoint of the first branch. The single semiconductor bridge may have a second branch that includes a third and a fourth switch. A second output terminal of the gradient coil may be coupled to a midpoint of the second branch. The first and the second branches are arranged in parallel with respect to an input DC bus. The first and the second terminals provide a signal that may be used to drive the gradient coil via, for example, a ripple filter.

Various features, aspects, and advantages of the present invention will become better understood when the following detailed description is read with reference to the accompanying drawings in which like characters represent like parts throughout the drawings, wherein:.

In an effort to provide a concise description of these embodiments, certain features of an actual implementation may be omitted in the specification. It should be appreciated that in the development of any such actual implementation, as in any engineering or design project, numerous implementation-specific decisions must be made to achieve the developers' specific goals, such as compliance with safety-related, system-related, and business-related constraints, which may vary from one implementation to another. Moreover, it should be appreciated that such a development effort might be complex and time consuming, but would nevertheless be a routine undertaking of design, fabrication and manufacture for those of ordinary skill having the benefit of this disclosure.

Magnetic resonance imaging (MRI) uses systems and devices that produce an image by mapping specific properties of gyromagnetic materials of the subject being imaged. More specifically, MRI systems excite gyromagnetic nuclei of the subject with strong magnetic fields and measure the radiofrequency (RF) emissions from the excited gyromagnetic nuclei in response to the excitation. The magnetic fields employed may be from multiple magnetic field sources in the MRI system. For example, the subject may be immersed by a strong constant magnetic field from a main coil, variable magnetic field gradients that provide spatial encoding, and RF excitation pulses that may have a frequency at or close to the resonance frequencies of the gyromagnetic nuclei. The variable magnetic field gradients may be used to add a spatial localization to the RF emissions from the gyromagnetic nuclei, which may be used to produce an image. The RF excitation pulses may be applied by RF excitation coils and may be used to excite the magnetization vector of the gyromagnetic materials through precession. A pulse sequence, a specific series of applications of these magnetic fields to the subject, may be employed to obtain suitable data from the MRI system to produce an image of an object.

The present application is generally related to systems and circuits that may be used to generate the variable magnetic field gradients. The magnetic field gradients can be generated by means of gradient coils, coils that may generate magnetization associated with spatial directions. For example, a two-dimensional MRI system may have two gradient coils (e.g., longitudinal, transversal) and a three-dimensional MRI system may have three gradient coils, with one gradient coil associated with each spatial direction (e.g., longitudinal, axial, transversal). Each gradient coil may be controlled independently from the other gradient coils to create strong magnetic fields. Moreover, the magnetization provided by the gradient coils may be adjusted over the course of the data acquisition procedure, to generate multidimensional images of the subject. As such, the circuitry driving the gradient coils, referred herein as the gradient coil driver circuitry, may have stringent performance specifications due to the high frequency switching, the large voltages and currencies, and the accuracy demands.

In order to satisfy the accuracy, high frequency, and large power specifications, traditional gradient power architectures often include gradient drivers with multilevel converters employing multiple cascaded semiconductor bridges. This type of design may allow high current accuracy while operating under the high powers employed when driving the gradient coils. However, the traditional gradient power architecture is bulky as it employs a large number of components. In the present disclosure, a more efficient design for a gradient power architecture is discussed. The present embodiments may use an integrated gradient power supply, or high-frequency power distribution unit (HFPDU), which may be capable of providing a reliable direct current (DC) output to the gradient amplifiers. The embodiments may also employ gradient amplifiers with a single bridge design to drive the coils, which may reduce power losses in the amplifier and the energy storage requirement at the input of the amplifier due to a shared DC bus across multiple axes. A method of grounding the system that reduces cross-talk between axes is also disclosed. The present embodiments may employ simplified amplifier design by reducing the number of bridges within the gradient amplifiers. To that end, the switches and/or semiconductor bridges employed
in the HFPDU and/or the gradient amplifiers may be implemented using high voltage wide band-gap devices (e.g., Silicon Carbide devices).

While the devices discussed herein are provided in the context of gradient amplifiers for MRI systems, it should be understood that the disclosed examples for the simplified high-power, high-frequency circuitry may be used to improve other systems that may employ multi-level converters by the simplification of the semi-conductor bridges, as understood in the art, provided that the resulting subject matter falls within the scope of the invention as defined by the appended claims.

With the foregoing in mind, <FIG> illustrates an example of an MRI system <NUM>. The MRI system <NUM> includes a scanner <NUM> and a scanner control system <NUM>. The scanner <NUM> may have a housing <NUM> around a bore <NUM>. A movable table <NUM> may be used to allow a patient <NUM> to be positioned within the bore <NUM>. The housing <NUM> of the scanner <NUM> may also include a primary magnet <NUM>, which may establish a primary magnetic field for data acquisition. Magnetic gradient coils <NUM>, <NUM> and <NUM> positioned in the scanner <NUM> may provide a magnetic field gradient that provides spatial encoding of the gyromagnetic nuclei of the patient <NUM> during the imaging process. An RF excitation coil <NUM> of the MRI system <NUM> may generate radiofrequency (RF) pulses for excitation of part of the gyromagnetic nuclei of the patient <NUM> during the imaging process. The MRI system <NUM> may also be provided with acquisition coils <NUM>, which may read out RF signals produced by gyromagnetic nuclei within the patient <NUM> as the nuclei go from an excited state to a relaxed state. In some embodiments, the RF excitation coil <NUM> and the acquisition coil <NUM> may be substantially the same. The various coils and magnets of the scanner <NUM> may be powered by a main power supply <NUM>.

The magnetic gradient coils <NUM>, <NUM>, <NUM>, that may be associated with an x-axis, a y-axis, and a z-axis, may be controlled by a gradient driver circuit <NUM>, which may adjust each spatial axis independently. The gradient driver circuit <NUM> may be powered by the main power supply <NUM>. The gradient driver circuit <NUM> may include gradient power supplies, and amplifiers that may control the gradients associated with the three axes independently, as detailed below. The RF excitation coil <NUM> may be controlled by an RF driver circuit <NUM>. Receive circuit <NUM> may acquire RF signals detected by the acquisition coil <NUM> and processed by a receive array switch <NUM>. The driver circuits <NUM> and <NUM>, and the receive circuit <NUM> may be coupled, through an interface <NUM>, to a controller <NUM>. Interface <NUM> may include memory banks and/or buffers that may be used to route the communication between the controller <NUM> and the driver circuits <NUM>, <NUM>, and receive circuit <NUM>. The controller <NUM> may include a general-purpose processor, an application-specific integrated circuit (ASIC) and/or a programmable logic device (PLD). The controller <NUM> may further communicate with a memory circuitry <NUM>, which may store data acquired through the receive circuit <NUM>. The memory circuitry <NUM> may also store instructions for the controller <NUM> in the form of imaging protocols. The imaging protocols may include instructions for the driver circuits <NUM> and <NUM> to control the gradient coils <NUM>, <NUM>, <NUM> and the RF excitation coil <NUM>, respectively, in a particular manner. Moreover, scanner control system <NUM> may have an interface <NUM> that allows a connection <NUM> between the MRI system <NUM> and other external equipment such as a computer cluster for image reconstruction or registration, a medical database, a diagnostic system, a PACS system, a display, a printer, a <NUM>-D visualization interface or any other device that may use MRI images or data. The memory circuitry <NUM> may also store instructions to control the gradient coils <NUM>, <NUM>, and <NUM>, by adjusting the currents in the coils via gradient driver circuitry <NUM>.

In order to control the gradient coils to generate the spatial encoding magnetic field gradients, the gradient driver may include a group of amplifiers. The MRI gradient power architecture <NUM>, illustrated in <FIG>, is a simplified power architecture that may be used to implement the gradient driver circuitry <NUM>. In the illustrated gradient driver circuitry <NUM>, the main power supply <NUM> provides the power that may be used by the gradient driver circuitry <NUM>. As an example, the main power supply may provide an alternating current (AC) signal. The voltage provided by the main power supply may have different rated voltages and/or frequency based on the geographical location or of the available power grid. For example, the voltages may be 380V, 415V, 480V, etc., and the signal frequency may be <NUM>/<NUM>. The main power supply <NUM> may supply power to the several single phase and three-phase MRI current drivers, including the gradient driver <NUM>.

The gradient driver circuit <NUM> may include the high-frequency power distribution unit (HFPDU) <NUM>, and the gradient amplifiers 106A, 106B, and 106C. The HFPDU <NUM> may be coupled to the power supply <NUM>, and may be receive three phases of the power supply <NUM>. The HFPDU <NUM> may provide a direct current (DC) signal to the gradient amplifiers 106A, 106B, and 106C via a shared DC bus <NUM>. In the illustrated example, the gradient amplifiers 106A, 106B, and 106C are responsible for driving the gradient coils <NUM>, <NUM>, and <NUM>. Each amplifier may be associated with a corresponding gradient coils. In the example, gradient amplifiers 106A, 106B, and 106C may drive the gradient coils <NUM>, <NUM>, and <NUM>, respectively. The controller <NUM> may control the provided currents and voltages and, ultimately, may control the magnetic gradients that provide spatial encoding during acquisition of data. To that end, the controller <NUM> may control switches (e.g., switches in the semiconductor bridges) in gradient amplifiers 106A, 106B, and 106C, to induce specific currents and/or voltages in the gradient coils <NUM>, <NUM>, and <NUM>, respectively.

As discussed above, the currents and voltages that drive gradient coils may be very large and subject to fast switching. Due to the stringent performance specifications, traditional power architectures may replicate several circuits and components, and include strong insulation structures in the power supplies, amplifiers, and gradient coils. The diagrams in <FIG>, <FIG>, <FIG>, and <FIG> are provided to illustrate the challenges present in traditional gradient power architecture systems. It should, however, be noted that teachings of the present application may be used within traditional power architectures, and as such, <FIG>, <FIG>, <FIG>, and <FIG> may also be examples of the present disclosure. In any case, the invention is defined only by the appended claims.

In view of this, the gradient power architecture <NUM> in <FIG> shows three driver circuits 201A, 201B, and 201C, whose outputs are electrically isolated from each other. Driver circuits 201A, 201B, and 20C drive the gradient coils <NUM>, <NUM>, and <NUM>, respectively. In the provided example, a power distribution unit <NUM> may be coupled to the power supply of the MRI system (e.g., main power supply <NUM>), and may provide power signals to driver circuits 201A, 201B, and 201C. In an example, the power distribution unit <NUM> may be a transformer that is coupled to the phases of the main power supply of the MRI system, and may provide an AC signal to the driver circuits 201A, 201B, and 201C.

Each driver circuit may receive the power signal from the power distribution unit <NUM>. In the example, a gradient power supply 204A receives the power signal from the power distribution unit <NUM>. The illustrated gradient power supply 204A generates multiple DC voltages, that are isolated from one another, which are provided to the gradient amplifier 206A. The multiple DC voltages in the gradient power supply 204A may facilitate generation of accurate currents to the gradient coil <NUM> by the multilevel converters used in the design of the gradient amplifier 206A. Similarly, a gradient power supply 204B may generate multiple isolated DC voltages to the gradient amplifier 206B, which may drive gradient coil <NUM>, and gradient power supply 204C may generate multiple isolated DC voltages to the gradient power amplifier 206C, which may drive gradient coil <NUM>. Note that, in this design, the gradient power supply 204A is coupled the gradient amplifier 206A via multiple dedicated DC buses 207A, gradient power supply 204B is coupled to the gradient amplifier 206B via dedicated DC buses 207B, and gradient power supply 204C is coupled to the gradient amplifier 206C via the dedicated DC buses 207C. The dedicated DC buses 207A, 207B, and 207C are decoupled from each other, to prevent cross-interference between the signals. Note, moreover, that each isolated circuity may have a ripple filter, between the gradient amplifier and the gradient coil. Ripple filters 208A, 208B, and 208C may be disposed in driver circuits 201A, 201B, and 201C to filter noise in the currents transmitted to the gradient coils <NUM>, <NUM>, and <NUM>, respectively.

The diagrams in <FIG>, <FIG>, and <FIG> are illustrations provided herein to depict challenges that may arise with the use of shared DC buses in a traditional power architectures, such as the gradient power architecture <NUM> illustrated in <FIG>. <FIG> illustrates, by means of an example, two gradient amplifiers, gradient amplifier 206A and gradient amplifier 206B. In this example, each gradient amplifier receives three DC power signals through a dedicated DC bus. The gradient amplifier 206A receives, via the dedicated DC bus 207A, two high voltage DC signals 220A and 222A, and a low voltage DC signal 224A. In the example, the gradient amplifier 206A is constructed using <NUM> semiconductor bridges 230A, 230B, and 230C. The semiconductor bridges are cascaded to form a multilevel converter. Each semiconductor bridge may be powered independently and, thus, receive a DC signal. As illustrated in this example, semiconductor bridge 230A is powered by the DC signal 220A, and may be coupled to output terminal <NUM> and with one leg of the semiconductor bridge 232A. In its turn semiconductor bridge 232A, which is powered by the DC signal 222A is coupled with one leg of the semiconductor bridge 234A. The semiconductor bridge 234A, which is powered by the DC signal 224A, is coupled to the output terminal <NUM>. As understood in the art, the switches in semiconductor bridges 234A, 234B, and 234C can be turned on and off at a high frequency with a variable duty ratio to control the currents provided between terminals <NUM> and <NUM>. The voltage output of the gradient amplifier 206A, across output terminals <NUM> and <NUM>, may be used to drive the current in the gradient coil <NUM>.

The gradient amplifier 206B, in <FIG> is arranged in a manner similar to the gradient amplifier 206A. The three semiconductor bridges 230B, 232B, and 234B, are powered by DC signals 220B, 222B, and 224B, respectively, via the dedicated DC bus 207B. The three semiconductor bridges are arranged in cascade to form the multilevel converter, and provide an output to terminals <NUM> and <NUM>. The voltage output between terminals <NUM> and <NUM> drives a current in the second gradient coil <NUM>. The current output in the gradient coil <NUM> can be controlled by turning on or off the switches in the semiconductor bridges 230B, 232B, and 234B at a high frequency. It should be noted that, in the system of <FIG>, dedicated DC buses 207A and 207B are separated, isolating the DC signals 220A and 220B, 222A and 222B, and 224A and 224B. <FIG> illustrates gradient amplifiers 206A and 206B similar to that of <FIG>, but with the contrast that the amplifiers are powered by a shared DC bus <NUM>. As a result of the use of the shared bus <NUM>, the gradient amplifiers 206A and 206B presents a coupling <NUM>.

In fact, due the coupling <NUM>, the system illustrated in <FIG> may suffer from DC voltage shoot-through fault during operation. Note that, both gradient amplifiers 206A and 206B receive, via the shared DC bus <NUM>, the same DC signals <NUM>, <NUM>, and <NUM>, coupled via the coupling <NUM>. Note, moreover, that the state of the switches in gradient amplifier 206A is independent from the state of switches in gradient amplifier 206B, as gradient coils <NUM> and <NUM> may be independently driven. As a result, short-circuits across the shared DC bus <NUM> may occur for certain states of the switches in the gradient amplifiers 206A and 206B. In fact, <FIG> illustrates a state of the switches in which a short circuit <NUM> on the shared DC bus <NUM> occurs. Due to the arrangement of switches, the short circuit <NUM> across the gradient amplifiers 206A and 206B may cause mal-function. Thus, as illustrated, gradient amplifiers implemented using a multilevel converter design with H-bridges, such as the gradient amplifiers 206A and 206B of <FIG>, may be vulnerable to short circuits, when powered with the shared DC bus <NUM>.

Traditional power architectures, such as the gradient power architecture <NUM>, may also suffer from high common-mode voltage stress in the gradient power supplies 204A, 204B, and 204C, and/or in the gradient amplifiers 206A, 206B, and 206C of <FIG>. Traditional power architecture <NUM> makes use of multilevel converters operating at high frequencies. In fact, the combined switching frequency in the gradient amplifiers 206A, 206B, and 206C may reach several tens of kHz. As a result of this high frequency switching, the circuits on the secondary side (i.e., the isolated DC voltage side) of the gradient power supplies 204A, 204B, and 204C, as well as the DC-link capacitor and bridges of the gradient amplifiers 206A, 206B, and 206C, may be subject to high frequency voltages relative to the ground (e.g., a safety ground, a system ground). These high frequency voltages may cause common-mode voltage stress in the cables, transformers, capacitors, and high-speed diodes. To prevent the common-mode voltage stress from causing damage, insulation requirements may be stringent. Further complicating the design, the common mode currents induced due to common-mode voltages in the system may also affect the fidelity of the amplifier, and may cause undesirable imaging artifacts. To reduce the high insulation requirement in the gradient coils <NUM>, <NUM>, and <NUM>, the output of the amplifiers can be center-point safety grounded in gradient filters <NUM>.

In view of the challenges presented by traditional architectures, such as the gradient power architecture <NUM>, <FIG>, <FIG> and <FIG> illustrate features of the gradient power architecture <NUM> discussed in <FIG>. The features discussed herein may benefit by, among other things, facilitate the use of shared DC buses and the addition of points for suitable grounding of the power circuitry. As a result, the design of the gradient power architecture <NUM> may decrease the replication of large and/or expensive components, reduce of the insulation requirements, reduce electrical stress due to common mode current leakage, and improve of the spatial encoding provided by the gradient coils <NUM>, <NUM>, and <NUM>. As illustrated in <FIG> the gradient power architecture <NUM> may employ an integrated HFPDU <NUM> that provides DC signals to the gradient amplifiers 106A, 106B, and 106C. These DC signals may be provided by the shared DC bus <NUM>. In some embodiments, the output DC signal in the shared DC bus <NUM> may have a voltage of in a range from 350V-2kV. The HFPDU <NUM> may include a power distribution unit <NUM> and a single power supply <NUM>.

<FIG> illustrates an electrical diagram for an embodiment for the HFPDU <NUM>. The HFPDU <NUM> includes the power distribution unit <NUM> and the power supply <NUM>. The power distribution unit <NUM> is coupled to the three phase terminals of the main power supply of the MRI system (e.g., main power supply <NUM>), and produce an intermediate DC signal provided to the power supply <NUM>. The power distribution unit <NUM> may include a line-side filter <NUM> and a rectifier circuit <NUM>. The power supply <NUM> may receive the unregulated intermediate DC signal from the power distribution unit <NUM> and create a regulated output DC signal provided to the gradient amplifiers via DC bus <NUM>. The power supply unit <NUM> includes H-bridge <NUM>, a high frequency transformer <NUM>, high-speed rectifiers <NUM>, a filter circuit <NUM>, the grounding Y-capacitor circuit <NUM>, and a capacitor bank <NUM>. As a result of this architecture, the HFPDU <NUM> may obviate the use of a bulky transformer. In fact, the power distribution unit <NUM> may be powered using different voltage levels and/or signal frequencies based on the electrical specification of the different locations.

As discussed above, the power supply <NUM> may receive the unregulated DC signal the power distribution unit <NUM> and create a regulated DC signal via the shared DC bus <NUM>. In the example, the power supply <NUM> may include an H-bridge <NUM>, which provides, via the high frequency transformer <NUM>, an internal high-frequency AC signal. In its turn, the rectifier <NUM> produces a DC signal, which may be provided to the DC bus <NUM>. Filter circuitry <NUM> may be used to filter noise produced in the rectifier <NUM>. In the gradient power circuitry, capacitor bank <NUM> be a particularly expensive component due to its dimensions and high capacitance requirement to meet peak gradient power demands. As the HFPDU <NUM> employs a single power supply <NUM>, the single capacitor bank <NUM> used to meet the peak power demand of the three gradient amplifiers 106A, 106B, and 106C. The reduction in the number of capacitors may decrease constraints associated with dimensions and thermal dissipation of the gradient amplifiers 106A, 106B, and 106C. Moreover, since the power for the three gradient amplifiers 106A, 106B, and 106C is provided from the single power supply <NUM>, a safety ground may be added to a midpoint in the capacitor bank <NUM>, as illustrated. Such a grounding of the capacitor bank <NUM> in the shared DC bus <NUM> does not lead to increased insulation requirement in the gradient coils. The output of the HFPDU <NUM> may be taken between terminals <NUM> and <NUM> of the shared DC bus <NUM>.

The HFPDU <NUM> may be coupled to the gradient amplifiers 106A, 106B, and 106C illustrated in <FIG>. The electrical coupling illustrated is between the gradient amplifiers 106A, 106B, and 106C and the terminals <NUM> and <NUM> of the shared DC bus <NUM>. The gradient amplifiers 106A, 106B, and 106C may be designed using single semiconductor bridges <NUM>, <NUM>, and <NUM>, respectively. The single bridge design may be used in place of the H-bridge multilevel converters illustrated, for example, in <FIG>. The single semiconductor H-bridges <NUM>, <NUM>, and <NUM> maybe implemented using switches designed with high voltage, wide band-gap devices (e.g. Silicon Carbide (SiC) devices). The high voltage, wide band-gap devices may allow accurate high-frequency control of very large currents and, thus, obviate the use of multiple H-bridges and multiple DC inputs as used in the traditional power architectures (e.g., power architecture <NUM> of <FIG>).

The illustrated gradient amplifier 106A includes a single H-bridge <NUM>. The single bridge <NUM> may include a first branch having switches 352A and 352B, and a second branch having switches 354A and 354B. The first branch controls the voltage at the midpoint of the first branch. Similarly, the second branch controls the voltage at the midpoint of the second branch. Terminals from the midpoints of the H-bridge <NUM> may be coupled to the input terminals of the ripple filter <NUM> and the output terminals 358A and 358B of the ripple filter <NUM> are coupled to the gradient coil <NUM>. The gradient ripple filter <NUM> may filter the high frequency noise from the single bridge <NUM> and reduce the ripple content in the current flowing into the gradient coil <NUM>.

Similarly, the illustrated gradient amplifier 106B includes a single bridge <NUM>. The single H-bridge <NUM> may include a first branch having switches 362A and 362B, and a second branch having switches 364A and 364B. Terminals from the midpoint of first and the second branches of the H-bridge <NUM> may be coupled to the input terminals of the ripple filter <NUM> and the output terminals 368A and 368B of the tipple filter <NUM> are coupled to the gradient coil <NUM>. The gradient ripple filter <NUM> may block high frequency noise from the single bridge <NUM> and reduce the ripple content in the current flowing into the gradient coil <NUM> from affecting the magnetic fields produced by gradient coil <NUM>.

The gradient amplifier 106C employs a similar design, with the single H-bridge <NUM>. The single H-bridge <NUM> may include a first branch having switches 372A and 372B, and a second branch having switches 374A and 374B. Terminals from the midpoint of the first and second branches of the H-bridge <NUM> may be coupled to the input terminals of the ripple filter <NUM> and the output terminals 378A and 378B of the ripple filter <NUM> are coupled to the gradient coil <NUM>. The gradient ripple filter <NUM> may block high frequency noise from the single bridge <NUM> and reduce the ripple content in the current flowing into the gradient coil <NUM>.

Technical effects of the embodiments presented herein include, generally, the reduction of the complexity of the power circuitry used in MRI systems. The use of an integrated HFPDU having a single power supply may allow a more compact power circuitry for the gradient subsystem of an MRI device. The HFPDU may also provide galvanic isolation and insulation, which facilitates meeting specifications for recommended means of patient protection (MOPPs). Moreover, the absence of cross-bridge connections in the gradient amplifiers may allow for an integrated DC bus, which further reduced the amount of wiring and the amount of insulation employed in the system. The shared DC bus further allows the use of a single capacitor bank, instead of multiple capacitor banks, as used in a traditional design.

It should be noted that the reduction and simplification in the size, the thermal stress, and the common mode voltage stress in the gradient amplifiers, due to the advantages discussed above, may allow the amplifier to be placed within the scan room. In fact, a single cable may be used to connect the HFPDU, located in an equipment room outside the scan room, to the gradient amplifiers, that may be placed within the scan room. The architecture may also allow flexible packaging of the parts, due to the reduction in the dimensions. For example, the gradient amplifiers may be deployed, each, in an individual package (e.g., box). Alternatively, the gradient amplifiers may share a single box, which may be in the scan room. In some systems, the gradient amplifiers may be disposed in the same box or packaging as the HFPDU.

Furthermore, the midpoint of a capacitor bank placed at the output of HFPDU may be connected to the safety ground. Such placement of the safety ground may reduce the above-discussed switching stress of the secondary side diode rectifiers, as the switching of the amplifier may not impact the common mode voltage at the terminals. This modification may lead to potential elimination of insulation failures in the secondary side diode and high frequency transformer. Finally, it should be noted that, by employing the methods and systems here described, the gradient ripple filter may become more effective and becomes primarily differential in nature as, in this design, it is subject to reduced common mode currents disturbances.

Claim 1:
A magnetic resonance imaging, MRI, system (<NUM>) configured to be powered by a main power source (<NUM>), the MRI system comprising: a plurality of
gradient coils; and
a gradient driver (<NUM>) configured to drive the plurality of gradient coils (<NUM>, <NUM>, <NUM>), the gradient driver comprising:
a plurality of gradient amplifiers (<NUM>), wherein each gradient amplifier of the plurality of gradient amplifiers is electrically coupled to a gradient coil of the plurality of gradient coils, and wherein each gradient amplifier comprises a respective single semiconductor bridge (<NUM>, <NUM>, <NUM>); and
a high frequency power distribution unit, HFPDU, (<NUM>) configured to receive an alternating current, AC, power signal from the main power source (<NUM>) and to generate a direct current, DC power signal and supply it to the plurality of gradient amplifiers via two terminals of a DC bus (<NUM>) that is shared among the plurality of gradient amplifiers to provide each of the plurality of gradient amplifiers is provided with the same DC voltage, wherein the HFPDU comprises:
a power distribution unit configured to receive the AC power signal from the main power source and produce an intermediate DC power signal; and
a power supply configured to receive the intermediate DC power signal from the power distribution and to provide the DC power signal to the DC bus, wherein the power supply comprises a transformer configured to isolate the DC power signal from the intermediate power signal, wherein the power supply comprises a capacitor bank comprising a capacitor connected between the two terminals of the DC bus and further comprising a grounding Y-capacitor circuit coupled to a safety ground of the MRI system, the grounding Y-capacitor circuit comprising a first capacitor connected between one terminal of the DC bus and the safety ground and a second capacitor connected between the other terminal of the DC bus and the safety ground.