METHOD FOR ACQUIRING MAGNETIC RESONANCE IMAGE DATA AND MAGNETIC RESONANCE DEVICE

Description

[0001] The present invention relates to a method for acquiring magnetic resonance image data of an examination subject using a magnetic resonance device. The invention is further directed to a corresponding magnetic resonance device.

[0002] Magnetic resonance imaging devices use a strong external magnetic field to align the nuclear spins of an object under investigation and, by applying an RF excitation pulse, excite them to precess around the corresponding alignment. The precession, or transition of the spins from this excited state to a lower-energy state, generates an alternating electromagnetic field in response, which can be detected as an MR signal via receiving antennas.

[0003] Using magnetic gradient fields, a position coding can be applied to the signals, which then allows the received signal to be assigned to a volume element of the object under examination. The received signal can then be evaluated, for example, to provide an image representation of the object under examination.

[0004] Objects, such as paramagnetic material, located within the object under investigation can alter the magnetic field surrounding the object due to a magnetic susceptibility that differs from that of the surrounding material. This can result in the magnetic field surrounding the object exhibiting local magnetic field gradients caused by the object. Due to these field gradients, different local phase changes can occur in the area surrounding the object during a repetition time. The locally induced field inhomogeneities can lead to protons within a volume element of a voxel being magnetized in antiphase, which can lead to a signal attenuation of the signal assigned to the voxel.In order to at least reduce signal attenuation, it is common practice to provide an additional gradient during the sequence to at least partially compensate for the dephasing caused by the object altering the magnetic field in the examination subject. This gradient is also referred to as a so-called white marker gradient. By providing the white marker gradient, objects altering the magnetic field can be made visible. The tissue in the immediate vicinity of the objects generates signals that cause high-intensity image areas, so-called artifacts. The artifacts stand out from the surrounding area due to their intensity. The surrounding area, which is characterized by a homogeneous magnetic field, is dephased by the white marker gradient. This provision of a white marker gradient is also referred to as dephased magnetic resonance imaging.White marker gradients can be used, for example, in gradient-echo sequences. A white marker contrast can be created by providing an extra gradient, or by omitting, reducing, and / or varying rephasing gradients typical for the sequence. The rephasing gradient can, for example, be a rephaser of a slice selection gradient, as described in Seppenwoolde et al. (2003), (SEPPENWOOLDE, Jan-Henry; VIERGEVER, Max A.; BAKKER, Chris JG. Passive tracking exploiting local signal conservation: the white marker phenomenon. Magnetic Resonance in Medicine: An Official Journal of the International Society for Magnetic Resonance in Medicine, 2003, 50th vol., no. 4, pp. 784-790.). In spin-echo sequences, it is not possible to increase signal strengths using the white marker gradient.The reason for this is that, due to the excitation of the object under investigation, the phase change of 180 degrees would reverse the dephasing introduced by the white marker gradient and the local field inhomogeneities. With spin-echo sequences, there is no artifact that could be rephased, since the dephasing caused by the field inhomogeneities around the object is already compensated for by the 180-degree pulse.

[0005] A special type of sequence is a balanced steady-state free precession sequence, or bSSFP for short. As with gradient echo sequences (GRE), magnetic field gradients are applied to generate echo signals. However, these are balanced during a repetition time (TR). As with gradient echo sequences, signal cancellation can occur near the object with a changing magnetic field due to local field inhomogeneities. The special signal generation in bSSFP can lead to an alternating phase at the echo time TE in a voxel, causing destructive spin interference. The special signal generation in bSSFP imaging results in a rephasing of the spins at time TE=TR / 2, which compensates for the dephasing caused by local field inhomogeneities. The bSSFP sequence can therefore also be described as a "spin-echo-like" sequence.The crucial difference, however, is that the phase to which the spins rephase depends on their phase advance in the time period TR, i.e., the time period between two excitation pulses. If this lies between (2n-1)*pi and (2n)*pi, where n is a natural number, the spins rephase with a phase advance opposite to the spins with a phase advance between (2n)*pi and (2n+1)*pi, where n is a natural number including zero. Due to the strong local field gradients in the immediate vicinity of an object subject to a magnetic field change, which leads to a locally varying accumulated spin phase during TR, the spins within a voxel can rephase with alternating phase at time TE, causing destructive signal interference. This signal attenuation can be partially compensated for by a white marker gradient, as is the case with gradient echo sequences.

[0006] Using a white marker gradient can unintentionally cause increased signal intensities at interfaces between different materials. This occurs, for example, at interfaces between fat and other tissue. The increased signal intensities are due to different resonance frequencies of the protons in the water of the tissue and the protons in the fat. The frequency difference between the resonance frequency of the protons in water and the resonance frequency of the protons in fat is 225 Hz at an external field strength of 1.5 Tesla. If the water protons and the fat protons are located in the same volume element, the deviation between the resonance frequencies leads to a phase difference between the water protons and the fat protons after the protons have been deflected.

[0007] At the beginning of the deflection of the magnetization from the longitudinal to the transverse and the transverse relaxation, the water protons and the fat protons are in phase. Due to the different resonance frequencies, the phase of the water protons changes more significantly than the phase of the fat protons. As a result, the two proton groups are out of phase and, at a magnetic field strength of, for example, 1.5 Tesla, are oppositely aligned after an initial echo time of 2.2 msec after excitation. Due to the opposite alignment of the protons in the volume element, a destructive superposition of the respective signals occurs, whereby the resulting volume element signal has a weak intensity at the first echo time. After a second echo time of 4.4 msec, the phase difference between the water protons and the fat protons is 0 degrees, so that both proton groups are in phase.This results in constructive superposition, which increases the intensity of the volume element signal. In gradient-echo imaging, this leads to the signal at the first echo time being canceled or attenuated, while the signal at the second echo time is constructively superimposed, resulting in the intensities adding up. Cancellation at interfaces can also occur in bSSFP sequences. After the steady state in bSSFP has been reached after several TRs, the water and fat protons are no longer in phase at the beginning of TR. Depending on the sequence design (choice of repetition time), these protons can have opposite phases at time TE=TR / 2, since, as described above, this phase is either 0 or 180 degrees depending on the phase accumulated during TR.If the spins of two materials at interfaces are out of phase at the time of the echo, the use of a white marker gradient can partially bring them back into phase and thus generate signal intensity.

[0008] The additional white marker gradient causes magnetization in tissue exposed to a homogeneous magnetic field to be dephased. However, the dephasing in inhomogeneous areas is reduced by the white marker gradient. This results in tissue areas exposed to a homogeneous magnetic field being reduced in intensity, while tissue areas exposed to an inhomogeneous magnetic field are simultaneously increased in intensity.

[0009] As described above, providing a white marker gradient results in a further change in signal intensity at the interfaces. When providing the white marker gradient in the sequence, the presence of fatty tissue can lead to an increase in signal intensity at volume elements at the interface between fatty tissue and other tissue. This not only increases the signal intensity in the vicinity of the object altering the magnetic field, but also the signal intensity along the interfaces. However, this can be undesirable if a prominent representation of the object altering the magnetic field is intended in the imaging, and increased signal intensity at interfaces is undesirable.

[0010] An exemplary bSSFP method is known from the article by Koktzoglou et al., "Dephased FLAPS for improved visualization of susceptibility-shifted passive devices for real-time interventional MRI", Physics in Medicine and Biology, Vol. 52, No. 13, 2007, pages N277-N286.

[0011] It is therefore an object of the invention to reduce signal intensity at interfaces in dephased balanced steady-state free precession sequences.

[0012] This object is achieved by the respective subject matter of the independent claims. Advantageous further developments and preferred embodiments are the subject matter of the dependent claims.

[0013] A first aspect of the invention relates to a method for acquiring magnetic resonance image data of an examination subject using a magnetic resonance device. In a first step of the method, it is provided that a balanced steady-state free precession sequence is performed by the magnetic resonance device. The steady-state free precession sequence is a magnetic resonance imaging sequence designed such that phase coherence of a transverse magnetization is partially maintained between the excitation times. In the balanced SSFP, the sequence is designed such that all gradient dephasing moments are compensated before the next excitation pulse. It is provided that the sequence, by means of an excitation pulse, usually causes a phase advance of 180 degrees per repetition time through the magnetic resonance device.It is intended that a white marker gradient be provided in the bSSFP sequence to at least partially compensate for local dephasing caused by an object that alters the magnetic field in the examination subject. In other words, the magnetic resonance device generates the white marker gradient, which attenuates local dephasing caused by the object that alters the magnetic field.

[0014] The magnetic resonance device acquires image data of the examination subject at a predetermined echo time. The echo time can be, for example, half the repetition time. In other words, the magnetic resonance device can acquire the image data of the examination subject at a time midpoint within the repetition time. The echo time can also deviate from half the repetition time by, for example, a maximum of 20%, 15%, 10%, or 5% of the repetition time.

[0015] It is provided that the magnetic resonance device performs an adaptation of a phase development between a phase of magnetization of a first substance and a phase of magnetization of a second substance. The first substance and the second substance can form an interface in the examination object. In other words, the substances can have different resonance frequencies, whereby the phases of the substances can change differently over time. Due to the different changes in the phases, a phase development between the two phases can occur. The phase development can be such that the phase of the first substance and the phase of the second substance can have a phase difference of 180 degrees at a time of the predetermined echo time.As a result, signals resulting from the phase of the protons of the first substance and signals resulting from protons of the second substance can destructively overlap, whereby a signal intensity of a signal from a volume element containing both protons can be reduced or canceled out.

[0016] The white marker gradient can lead to a change in the phases of the two substances in the volume element. This can result in a phase difference of 180 degrees at the predetermined echo time. The extent of destructive signal superposition is thus reduced for that time. This leads to an increase in signal intensity in the volume element containing both the first and second substances.

[0017] However, this increase in signal intensity due to the white marker gradient may be undesirable at interfaces. In this case, it is necessary to reduce the signal intensity in these volume elements.

[0018] The invention provides for changing the phase development between the phase of the magnetization of the first substance and the phase of the magnetization of the second substance such that the phases, without the effect of the white marker gradient, are in phase with the predetermined echo time. For this purpose, the magnetic resonance device can adapt the phase development such that the phase development before the effect of the white marker gradient exhibits an in-phase alignment of the magnetization of the first substance and the magnetization of the second substance at the interface. In general, without the presence of the white marker gradient, with an in-phase alignment of the substances, an addition of the signal intensities occurs in the volume element that contains both substances and thus images an interface between the two substances.However, due to the presence of the white marker gradient, in the described situation the signal is weakened in the said volume element of the interface.

[0019] Due to the phase adjustment between the phases of the two substances, the provision of the white marker gradient results in a decrease in signal intensity. Without the phase adjustment, however, the signal intensity would increase.

[0020] The invention provides the advantage that the white marker gradient enables an increase in the signal intensity in an area of ​​an object that changes the magnetic field without simultaneously increasing the signal intensity at an interface between the two substances.

[0021] The invention also includes further developments which result in further advantages.

[0022] A further development of the invention provides that adapting the phase development comprises setting a scanner frequency of the BSSFP sequence. The scanner frequency is selected by the magnetic resonance device such that it is midway between a Larmor frequency of the first substance and a Larmor frequency of the second substance. The scanner frequency fScanner is obtained according to the formula fScanner = 0.5* (fLarmor,m1 + fLarmor,m2) from the Larmor frequency of the first substance fLarmor,m1 and the Larmor frequency of the second substance fLarmor,m2. In other words, the scanner frequency has a value that is obtained by adding the Larmor frequency of the first substance to the Larmor frequency of the second substance and then dividing the sum by 2. The scanner frequency is thus the arithmetic mean between the Larmor frequency of the first substance and the Larmor frequency of the second substance.By adjusting the scanner frequency, the object under investigation is not excited with the usual Larmor frequency of water, but with a different scanner frequency. The change in the scanner frequency results in the temporal phase change of the magnetization of the first substance being identical to the temporal phase change of the magnetization of the second substance. However, the directions of the temporal phase changes of the two substances are opposite. The magnetization along the interface is thus refocused in phase. Due to the in-phase refocusing of the protons of both substances, the applied white marker gradient at the interface leads to a dephasing instead of a rephasing of the magnetization, whereby the signal intensity along the interface is reduced by the white marker gradient.

[0023] A further development of the invention provides that the adaptation of the phase development comprises providing an additional phase advance per repetition time of the bSSFP sequence, wherein the following applies to the additional phase advance of the bSSFP sequence: ((fLarmor,m1 - fLarmor,m2) / 2) x TR x 360 degrees. In other words, it is provided that the Larmor frequency of water is used as the scan frequency for excitation during the sequence. The adaptation of the phase development takes place in that, in addition to the provision of the usual phase advance per repetition time, a further phase advance is provided during the respective repetition time, which results from the aforementioned formula. In other words, the value of the usual phase advance is supplemented by an additional value of the further phase advance.The additional phase advance depends on the Larmor frequency of the first substance and the Larmor frequency of the second substance. Thus, the Larmor frequency of the first substance is still used as the scanning frequency, and instead of adjusting the scanning frequency, the phase cycle of the excitation pulses is changed to shift the band structure of the sequence relative to the resonant frequencies of the first substance and the second substance. Typically, the sequence is designed to provide a phase advance of 180 degrees per repetition time so that the first substance resonates at a center of the central band, assuming that the scanning frequency matches the Larmor frequency of the first substance. By setting the scanning frequency to the resonant frequency of the first substance and providing the additional phase advance between the pulses, the resonant frequencies are placed symmetrically in the band structure.

[0024] A further development of the invention provides that the first substance is fat and the second substance is water. In other words, the method is intended to reduce signal intensity at an interface between fat and water.

[0025] This combination is particularly relevant for tissue interfaces, because interfaces between adipose tissue and other tissue can produce undesirably high signal intensities due to the protons of the fat and the protons of the water of the other tissue in a common volume element. Interfaces between adipose tissue and other tissue are particularly relevant in biological studies.

[0026] A further development of the invention provides that the first substance is silicone and the second substance is water. In other words, the method is intended to reduce signal intensity at an interface between water and silicone. This combination is particularly relevant for interfaces between tissue and implants.

[0027] For use cases or application situations that may arise during the method and which are not explicitly described here, it may be provided that, in accordance with the method, an error message and / or a request to enter user feedback is issued and / or a default setting and / or a predetermined initial state is set.

[0028] A second aspect of the invention relates to a magnetic resonance device for acquiring magnetic resonance image data of an examination subject. The magnetic resonance device is configured to provide a balanced steady-state free precession sequence with a usual phase progression of 180 degrees per repetition time. The magnetic resonance device is configured to provide a white marker gradient in the sequence in order to at least partially compensate for a dephasing caused by an object in the examination subject that alters the magnetic field. The magnetic resonance device is configured to acquire image data of the examination subject at a predetermined echo time.It is provided that the magnetic resonance apparatus is configured to adapt a phase evolution between a phase of a magnetization of a first substance and a phase of a magnetization of a second substance, which form an interface in the examination object, in the balanced steady-state free precession sequence. It is provided that the adaptation of the phase evolution before the effect of the white marker gradient brings about an in-phase alignment of a magnetization of the first substance and the second substance at the interface at the echo time.

[0029] Further embodiments of the magnetic resonance device according to the invention follow from the various embodiments of the method according to the invention.

[0030] According to a third aspect of the invention, a computer program is provided which can be loaded directly into a memory of a control device of an above-mentioned magnetic resonance apparatus, with program means for carrying out the steps of the above-mentioned method when the program is executed in the control device of the magnetic resonance apparatus. Likewise, an electronically readable data carrier is provided with electronically readable control information stored thereon, which comprises at least one computer program of the above type and is designed such that it carries out a method of the above type when the data carrier is used in a control device of the above-mentioned magnetic resonance apparatus. The system can in particular comprise a computer, a microcontroller, or an integrated circuit. Alternatively, the magnetic resonance apparatus can comprise a real or virtual network (i.e., cluster) of computers.In this exemplary embodiment, the magnetic resonance apparatus comprises a processor or a computing unit. The computing unit can comprise hardware elements and / or software elements, for example, a microprocessor or a so-called FPGA (Field Programmable Gate Array). A memory unit can be implemented as a temporary random access memory (RAM) or as a permanent mass storage device (hard drive, USB stick, SD card (SSD)).

[0031] A computing unit can be understood, in particular, as a data processing device that contains a processing circuit. The computing unit can therefore, in particular, process data to perform computing operations. This may also include operations for performing indexed access to a data structure, for example, a look-up table (LUT).

[0032] The computing unit may, in particular, contain one or more computers, one or more microcontrollers, and / or one or more integrated circuits, for example, one or more application-specific integrated circuits (ASICs), one or more field-programmable gate arrays (FPGAs), and / or one or more single-chip systems (SoCs). The computing unit may also contain one or more processors, for example, one or more microprocessors, one or more central processing units (CPUs), one or more graphics processing units (GPUs), and / or one or more signal processors, in particular one or more digital signal processors (DSPs). The computing unit may also include a physical or virtual network of computers or other of the aforementioned units.

[0033] In various embodiments, the computing unit includes one or more hardware and / or software interfaces and / or one or more memory units.

[0034] A memory unit can be a volatile data storage device, such as dynamic random access memory (DRAM) or static random access memory (SRAM), or a non-volatile data storage device, such as read-only memory (ROM), programmable read-only memory (PROM), erasable read-only memory (EPROM), electrically erasable read-only memory (EEPROM), flash memory (FRAM), ferroelectric random access memory (FRAM), magnetoresistive random access memory (MRAM), or phase-change random access memory (PROM).be configured as a PCRAM (English: "phase-change random access memory").

[0035] The at least one evaluation unit, the control device and / or the measuring unit of the magnetic resonance device according to the invention may include one or more computing units according to this understanding, or one or more computing units of the magnetic resonance device may include the at least one evaluation unit, the control device and / or the measuring unit.

[0036] If, within the scope of the present disclosure, it is stated that a component of the magnetic resonance device according to the invention, in particular the control device, the measuring unit or at least one evaluation unit of the magnetic resonance device, is set up, designed, configured or the like to carry out or implement a specific function, to achieve a specific effect or to serve a specific purpose, this can be understood to mean that the component, beyond the fundamental or theoretical usability or suitability of the component for this function, effect or purpose, is concretely and actually capable of carrying out or implementing the function, achieving the effect or serving the purpose through appropriate adaptation, programming, physical design and so on.

[0037] A connection between two electrical or electronic components can, unless expressly stated otherwise, be understood to mean that an electrical connection exists between the components or can be established by actuating one or more switching elements. In particular, the components can be connected to one another directly or indirectly, unless stated otherwise. A direct connection can be understood to mean that, apart from the optional one or more switching elements, no further electrical or electronic components are arranged between the components, whereas an indirect connection can be understood to mean that, in addition to the optional one or more switching elements, one or more further electrical or electronic components, such as resistors, capacitors, coils, and so on, are arranged between the components.

[0038] Further embodiments of the computer program according to the invention follow from the various embodiments of the method according to the invention and the magnetic resonance device according to the invention.

[0039] A fourth aspect of the invention relates to an electronically readable data carrier having electronically readable control information stored thereon, which comprises at least one computer program according to the third aspect of the invention and is designed such that, when the data carrier is used in a control device of a magnetic resonance apparatus according to the second aspect of the invention, it carries out a method according to the first aspect of the invention. Further embodiments of the electronically readable data carrier according to the invention follow from the various embodiments of the method according to the invention, the magnetic resonance apparatus according to the invention, and the computer program according to the invention.

[0040] Further features of the invention emerge from the claims, the figures and the description of the figures. The features and combinations of features mentioned above in the description as well as the features and combinations of features mentioned below in the description of the figures and / or shown in the figures can be encompassed by the invention not only in the respectively specified combination, but also in other combinations. In particular, the invention can also encompass embodiments and combinations of features that do not have all the features of an originally formulated claim. Furthermore, the invention can encompass embodiments and combinations of features that go beyond the combinations of features set out in the backreferences to the claims or deviate from them.

[0041] The invention is explained in more detail below using specific embodiments and associated schematic drawings. In the figures, identical or functionally equivalent elements may be provided with the same reference numerals. The description of identical or functionally equivalent elements may not necessarily be repeated for different figures. The figures show: FIG. 1 shows a schematic representation of an exemplary embodiment of a magnetic resonance apparatus according to the invention; FIG. 2 shows a schematic representation of an exemplary embodiment of a sequence of a method according to the invention; FIG. 3 shows a schematic representation of an exemplary readout gradient of a dephased bSSFP sequence; FIG. 4 shows a schematic representation of an exemplary bSSFP signal formation in the vicinity of a magnetic field-altering object; FIG. 5 shows a schematic representation of an exemplary bSSFP signal formation around a paramagnetic needle for bSSFP-based sequences without a WM gradient. FIG 6 shows a schematic representation of an exemplary bSSFP signal formation around a paramagnetic needle for bSSFP-based sequences with WM gradient. FIG 7 shows a schematic representation of an exemplary analysis of signal formation in the dephased bSSFP at fatty tissue interfaces. FIG 8 shows a schematic representation of an exemplary analysis of signal formation in the dephased bSSFP at fatty tissue interfaces. FIG 9 shows a schematic representation of an exemplary simulation of a virtual human phantom containing a virtual paramagnetic needle. FIG 10 shows a schematic representation of an exemplary dephased bSSFP sequence.

[0042] In FIG 1 An exemplary embodiment of a magnetic resonance device 1 according to the invention is shown schematically.

[0043] The magnetic resonance apparatus 1 comprises a magnetic resonance imaging system 5 with a field magnet 6 that generates a static magnetic field for aligning nuclear spins of an examination subject 4, for example, a patient, in an imaging region. The imaging region is characterized by an extremely homogeneous static magnetic field, wherein the homogeneity relates in particular to the magnetic field strength or its amplitude. The imaging region is located in a patient tunnel that extends in a longitudinal direction Z through the magnetic resonance imaging system 5. The field magnet 6 can, for example, be a superconducting magnet that can generate magnetic fields with a magnetic flux density of up to 3 Tesla or more. However, for lower field strengths, permanent magnets or electromagnets with normally conducting coils can also be used. A patient table 3 can be movable within the patient tunnel.

[0044] Furthermore, the magnetic resonance imaging system 5 comprises a gradient coil arrangement 7 with a plurality of gradient coils, which serve to superimpose gradient fields, i.e., location-dependent magnetic fields, on the static magnetic field in the three spatial directions for spatial differentiation of the scanned image areas in the imaging area. The gradient coils of the gradient coil arrangement 7 can be configured, for example, as coils made of normally conducting wires, which can generate, for example, mutually orthogonal fields or field gradients in the imaging area.

[0045] The magnetic resonance imaging scanner 5 comprises a transmit coil arrangement, which may, for example, comprise a body coil 8 as a transmit antenna, which is configured to emit a radio-frequency signal into the imaging region. The body coil 8 can therefore be understood as an RF transmit coil arrangement of the magnetic resonance apparatus 1 or as part of the RF transmit coil arrangement. In some embodiments, the body coil 8 can also be used to receive resonant MR signals emitted by the examination subject 4. In this case, the body coil 8 can also be considered part of a signal acquisition device of the magnetic resonance apparatus 1. Optionally, the signal acquisition device comprises a local coil 9, which can be arranged in the immediate vicinity of the examination subject 4, for example, on the examination subject 4 or in the patient table 3.The local coil 9 can serve as a receiving coil or receiving antenna alternatively or in addition to the body coil 8.

[0046] The magnetic resonance apparatus 1 also includes a control device 2. The control device 2 may include a reception control unit, which can be considered part of the signal acquisition device and is connected to the body coil 8 and / or the local coil 9. Depending on the acquired MR signals, the reception control unit, which may include an analog-to-digital converter (ADC), can generate corresponding MR data, particularly in k-space. A computing unit of the control and computing system can evaluate the MR data and, for example, perform image reconstruction to generate an MRI image 10 of the examination subject 4.

[0047] The control device 2 comprises a transmission control unit connected to the body coil 8 and controlling it to generate RF pulses, such as excitation pulses and / or refocusing pulses. The control device 2 comprises a gradient control unit connected to the gradient coil arrangement 7 and controlling it to switch slice selection gradients, gradients for frequency and / or phase encoding, and / or readout gradients.

[0048] It should be noted that the described structure of the control device 2 represents only a non-limiting example. The various required tasks and functions can also be different and / or distributed among different control units and / or other computing units. The magnetic resonance apparatus 1 is particularly capable of performing an MRI method according to the invention. FIG 2 shows a schematic representation of an exemplary embodiment of a sequence of a method according to the invention; the method can be provided for generating an MRI image 10 of the examination subject 4 by a magnetic resonance device 1.

[0049] In a first step S1, a control device 2 of the magnetic resonance device 1 can transmit control data for controlling the magnetic resonance device 1 in order to control the magnetic resonance device 1 to carry out a balanced steady-state free precession sequence.

[0050] The magnetic resonance apparatus 1 can perform the balanced steady-state free precession sequence with a usual phase advance Φ of 180 degrees per repetition time TR, wherein in the balanced steady-state free precession sequence a white marker gradient GWM is provided by the magnetic resonance apparatus 1 in order to at least partially compensate for a dephasing caused by a magnetic field-altering object in the examination object 4.

[0051] A second step S2 may include adjusting a phase development between a phase of a magnetization of a first substance and a phase of a magnetization of a second substance, which form an interface in the examination object 4, in the balanced steady-state free precession sequence by the magnetic resonance device 1,whereby the adaptation of the phase development before the effect of the white marker gradient GWM causes an in-phase alignment of a magnetization of the first substance m1 and the second substance m2 at the interface to the echo time TE.

[0052] A third step S3 of the method may comprise acquiring image data of the examination subject 4 by the magnetic resonance device 1 at an echo time TE. The echo time TE may be half a repetition time TR or deviate from it by a predetermined value. The image data may be evaluated by the control device 2 to generate an MRI image 10 of the examination subject 4.

[0053] FIG 3 shows a schematic representation of an exemplary readout gradient Gread of a dephased bSSFP sequence.

[0054] Balanced Steady State Free Precession (bSSFP) imaging describes a specific signal generation mechanism in MRI. This signal generation mechanism is characterized by the fact that all magnetic field gradients Gx, Gy, and Gz are balanced at the end of each repetition time TR, and a coherent phase relationship exists between successive excitation pulses RF. White marker gradients (GWM) can be introduced into gradient-echo sequences, but not into spin-echo sequences, since the usual phase advance Φ of 180 degrees would reverse the introduced dephasing. Although the bSSFP sequence can be classified as a spin-echo sequence, a combination of the sequence with white marker gradients (GWM) is possible, since the magnetization rephases at half a repetition time either at 0 or 180 degrees, depending on the phase advance of the same during a repetition time.Field inhomogeneities can thus lead to antiphase magnetization within a voxel and thus to signal cancellation, which can be counteracted using a white marker gradient.

[0055] Shown is the course of the readout gradient Gread in the dephased bSSFP sequence. All gradient moments are equalized before the end of the repetition time TR to achieve bSSFP signal formation. An additional white marker gradient GWM introduces an additional phase. In areas of homogeneous susceptibility, the white marker gradient GWM causes a signal loss due to the introduced dephasing. In the vicinity of objects with changing magnetic fields, the white marker gradient GWM can counteract a signal reduction induced by local interference gradients GS caused by a susceptibility difference, thus increasing or even restoring the signal intensity of a signal.

[0056] FIG 4 shows a schematic representation of an exemplary bSSFP signal formation in the vicinity of a magnetic field-altering object;

[0057] Under FIG 4 a schematically shows a section through an examination object 4. A paramagnetic needle 11 is placed in a tissue 12 of the examination object 4 as an object that changes the magnetic field on the left side of the MRI image 10. The needle 11 induces a local disturbance gradient GS in the magnetic field, which causes a magnetic field in the tissue 12 to change around the needle 11. This leads to a spatially varying phase development during the repetition time TR. Due to the bSSFP signal formation, the magnetization at TR / 2 is refocused with a phase of either 0° or 180°, depending on the local phase development. The circles with the pointers represent the phase of the local magnetization of the protons of the tissue 12 at the echo time TE. The diagram shows the phase at TE versus the phase progression per TR.

[0058] (b) shows an MRI image 10 of an examination subject 4 around the needle 11, which exhibits the contrast generated by dephased bSSFP. The needle 11 may have been inserted into the surrounding tissue 12 of the examination subject 4. The subject is characterized by a dark artifact resulting from a signal loss caused by the refocusing of the intravoxel magnetization during TE with changing phase, as shown in (a), resulting in a destructive superposition of signal components.

[0059] Due to the magnetic field disturbance induced around needle 11 and the resulting magnetization dephasing, the inherent signal formation behavior of the sequence leads to an alternating phase at the echo time TE in the magnetization of tissue 12 around needle 11. The susceptibility-induced disturbance gradient GS causes the magnetization to have a spatially dependent resonance frequency (Larmor frequency), which consequently leads to the magnetization accumulating a spatially dependent phase during the repetition time TR. Depending on this accumulated phase, the echo is formed either at 0 degrees or 180 degrees. The characteristic band structure is depicted below (a). If this occurs in the same voxel, the antiphase magnetization cancels out and the signal is lost, resulting in a dark image artifact near needle 11 in MRI image 10.

[0060] FIG 5 shows a schematic representation of an exemplary bSSFP signal formation around a paramagnetic needle 11 for bSSFP-based sequences without white marker gradient GWM.

[0061] Shown is a signal formation around a paramagnetic needle 11 for bSSFP-based sequences without white marker gradient GWM. Similar to FIG 4 A section through tissue 12 with a needle 11 is schematically depicted. The needle 11 induces a locally alternating phase at the echo time TE in the tissue 12, represented by the circles with the pointers. This manifests as a signal loss around the needle 11 when the alternating phase of the magnetization cancels out within the same voxel V1, V2. The image contrast for this case is shown in the MRI image 10 of the examination subject 4.

[0062] FIG 6 shows a schematic representation of an exemplary bSSFP signal formation around a paramagnetic needle 11 for bSSFP-based sequences with white marker gradient GWM.

[0063] Shown is signal formation around a paramagnetic needle 11 for bSSFP-based sequences with white marker gradient GWM. Applying a white marker gradient GWM can partially refocus the phase of the protons within voxels V1 and V2, thus restoring the signal. A typical WM contrast is formed in MRI image 10. The image contrast in this case is shown in the image of the examination subject 4.

[0064] Dephased MRI makes it possible to generate image contrast that highlights magnetic field-altering objects such as paramagnetic interventional objects, e.g., biopsy needles 11 or catheters. This is achieved, for example, by applying additional so-called "white marker (WM)" gradients or white marker gradients GWM. Objects with a different susceptibility than the tissue 12 locally disrupt the magnetic field and cause local disturbance gradients GS in the magnetic field. These lead to signal loss around the object due to the dephasing of the local magnetization. Applying the additional white marker gradients GWM causes a dephasing of the magnetization in tissue 12 with a homogeneous magnetic field B0, but can partially counteract the dephasing induced by the magnetic field-altering object.This leads to the formation of a contrast characterized by signal-intensive areas around the object altering the magnetic field and a dark dephased background at a greater distance from the interferer.

[0065] Signal loss near objects that alter the magnetic field can be counteracted with a white marker gradient. The introduction of the additional white marker gradient (GWM) results in at least partial rephasing of the magnetization around the object that alters the magnetic field. Dephased bSSFP has been used in the past to localize objects in interventional MRI. Since bSSFP generally has a higher signal-to-noise ratio than GRE and magnetization rephasing around objects that alter the magnetic field does not depend on gradient polarity due to the symmetric alternating phase behavior, it can be considered advantageous over dephased GRE in these aspects.

[0066] FIG 7 shows a schematic representation of an exemplary analysis of signal generation in dephased bSSFP at interfaces.

[0067] FIG 7 b,c shows an interface at which a first substance m1 borders on a second substance. A voxel can be assigned to a volume element in which both substances m1, m2 can be located. The substances m1, m2 can be, for example, fat and adjacent tissue 12. If a phase difference phi between fat and adjacent tissue 12 (e.g. B. muscle) develops, d. h. in the range (2*n-1)*pi < phi < 2*n*pi with n natural, the magnetization is refocused at the echo time TE with a phase difference of 180 degrees across the tissue interface ( FIG 7 a) . The graph shows the phase at TE versus the phase advance per TR. This can lead to a signal drop due to phase cancellation in the voxel containing the tissue boundary ( FIG 7 b ). The application of the white marker gradient GWM can partially rephase the magnetization at the tissue interface, resulting in an increase in the signal ( FIG 7 c) Since the resonance frequencies of fat and water are separated by a chemical shift of approximately 3.5 ppm, it is possible that the magnetization in the different tissue types forms an echo with opposite phase, as the relative phase shift between the two tissue types causes them to fall into different phase bands. This occurs for all values ​​of the repetition times TR that result in a phase difference between fat and adjacent tissue 12 (e.g., B.Muscle) of (2*n-1)*pi < phi < 2*n*pi, where n is natural. This effect leads to increased signal intensity at the interfaces of the adipose tissue 12, since the white marker gradient GWM has a rephasing effect at these interfaces, comparable to the effect near the metal device.

[0068] FIG 8 shows a schematic representation of an exemplary analysis of signal generation in dephased bSSFP at adipose tissue 12 interfaces.

[0069] By adjusting the sequence, the phase evolution for the magnetization of the adipose tissue 12 and the adjacent tissue 12 is intentionally changed so that the magnetization at the interfaces of the adipose tissue 12 is refocused with the same phase.

[0070] To achieve this, for example, a scanner frequency d. h. The frequency to which the receiver / transmitter is tuned must be deliberately adjusted so that it is fScanner = 0.5*(fLarmor,m1 + fLarmor,m2). The position of the Larmor frequencies of the substances m1 and m2 in the band structure is in FIG 8 a. The graph shows the phase at TE versus the phase advance per TR. By choosing the scanner frequency in this way, instead of the standard choice of the water frequency as the scanner frequency, the phase shift for the magnetization in the first material has the same value but opposite sign as for the magnetization from the second material, and the magnetization at the interfaces is refocused with the same phase. Consequently, the white marker gradient GWM has a dephasing rather than a rephasing effect on the magnetization, and the signal strength is suppressed.

[0071] By shifting the scanner frequency to fScanner = 0.5* (fLarmor,m1 + fLarmor,m2), for example, the phase shift for the magnetization in the fatty tissue 13 has the same value with the opposite sign as for the magnetization from the other tissue 12. This refocuses the magnetization of the protons across the interface in the voxel with the same phase ( FIG 8 b) . Consequently, the white marker gradient GWM has a dephasing effect instead of a rephasing effect on the magnetization and the signal strength of the signal of the shown voxel is suppressed ( FIG 8 c) .

[0072] FIG 9 shows a schematic representation of an exemplary simulation of a virtual human phantom containing a virtual paramagnetic needle 11.

[0073] Using a virtual human phantom containing a virtual paramagnetic needle 11, Bloch simulations were performed to generate a phase-shifted bSSFP contrast as used in the literature for needle tracking.

[0074] With a sequence based on the dephased bSSFP readout scheme from FIG 3 based, two images were created. FIG 9 b shows the contrast with the scanner frequency fScanner centered on the Larmor frequency of the water. FIG 9 c shows the contrast with fScanner = 0.5* (fLarmor,Fat + fLarmor,Water). This results in the magnetization at the interfaces of adipose tissue 12 being refocused with the same phase at TE. The white marker gradient GWM results in a homogeneous dephasing across voxels V1 and V2 at the adipose tissue 12 interface, and no additional signal intensity is caused.

[0075] Under FIG 9 a shows the virtual human phantom. A transverse section from the simulated dataset containing the needle 11 is shown in FIG 9 b It can be seen how the contrast technique highlights the area surrounding the needle 11, which is desirable. However, a high signal is also visible at the interfaces to the fatty tissue 13, which is undesirable. Another transverse section from the simulated dataset, which contains the needle 11, is shown in FIG 9 c. The scanner frequency was intentionally changed from fScanner = fLarmor,m1 to fScanner = 0.5*(fLarmor,m1 + fLarmor,m2). This results in the magnetization in the fat and adjacent tissue 12 being refocused with the same phase at TE, so that the white marker gradient GWM cancels the signal at the interfaces and suppresses the intensity. Any remaining signal at the interfaces may be due to a partial volume effect. Consequently, the needle 11 is more clearly visible.

[0076] FIG 10 shows a schematic representation of an exemplary dephased bSSFP sequence with an additional phase advance Φ. In order to provide a dephased bSSFP, an additional GWM moment can be introduced, in contrast to a normal bSSFP. This can be achieved, for example, by selecting a rephasing lobe of the slice selection gradient Gz that is not exactly 1 / 2x as large as the slice selection gradient applied during the rf pulse. For example, if a larger moment is selected, as in Fig. 10 drawn, this must be compensated again at the end of TR by a layer selection prephaser lobe.

[0077] In the prior art, it is common to change the NCO (Numerically Controlled Oscillator) phase by 180 degrees, so that it alternates, for example, between +90 degrees and -90 degrees. In the described method, the change in the NCO phase per repetition time can deviate from 180 degrees, so that it does not alternate between +90 degrees and -90 degrees.

[0078] Instead of changing the scanner frequency, it may be intended to change the phase cycle of the bSSFP RF pulses instead to achieve the same goal of shifting the bSSFP band structure relative to the fat and water resonances.

[0079] A standard bSSFP sequence uses a phase shift of 180° / TR to place the water resonance in the middle of the central band, assuming that the scanner frequency is tuned to the water frequency.

[0080] If the scanner frequency is set to the water resonance, which is the default for a standard frequency calibration, and an additional phase advance of ((fLarmor,fat - fLarmor,water) / 2) * TR * 360° is applied from pulse to pulse, the resonances of the two substances are inserted symmetrically into the band structure.

[0081] In dephased bSSFP, unwanted signal may originate from adipose tissue interfaces 13, which is due to the signal generation mechanism in bSSFP imaging described above. This additional signal magnitude can be suppressed with the presented approach.

[0082] This offers several advantages. It allows for a clearer image contrast, which can improve the visibility of the object that is changing the magnetic field. z. B.This is advantageous for the object localization use case, especially when either only degraded image data is used for fast subsampled image acquisition or when machine learning is to be used for automatic needle localization. The additional size introduced by the adipose tissue interfaces increases the complexity of the image and consequently the complexity of the localization task, and therefore negatively impacts localization accuracy if not suppressed.

[0083] The solution allows full flexibility in the choice of the repetition time TR, since the accumulated phase difference between adipose tissue 13 and other tissue 12 becomes irrelevant by the provided approaches.

[0084] In a fully balanced sequence, the repetition time TR would be chosen for robust signal generation. Thus, a repetition time TR would have to be selected that removes the fat and water resonances from the dark bSSFP stopbands. However, this is not necessary for "dephased bSSFP." In regions distant from the magnetic field-altering object, the signal is supposed to be suppressed using the white marker gradients GWM anyway, so it doesn't matter whether the stopbands are hit. Near the magnetic field-altering objects, due to the large B0 inhomogeneity, many stopbands and passbands will fall into one voxel V1, V2. The repetition time TR is irrelevant, as this always creates an intra-voxel V1, V2 averaging effect across the band structure in the areas around the object relevant for the WM signal.

[0085] This allows flexibility in choosing the repetition time TR. Using a fat suppression technique could result in the phase-shifted signal disappearing at the water-fat interfaces. However, this would be counterproductive if the subject traverses adipose tissue, since no WM contrast can be generated if the fat signal is completely suppressed.

[0086] The proposed solutions enable a much more elegant solution.

[0087] It is likely that the effect of unwanted signal rephasing may also play a role in dephased GRE imaging. However, unlike dephased bSSFP with the proposed method, this unwanted signal cannot be suppressed, so dephased bSSFP with the proposed method is theoretically the better technique in this respect.

[0088] This invention achieves a significant improvement in the contrast produced by dephased bSSFP at interfaces between adipose tissue 13 and adjacent tissue 12, which normally oscillate near the water resonance frequency.

Claims

1. Method for capturing magnetic resonance image data of an object under examination (4) using a magnetic resonance apparatus (1), comprising the steps: - performance of a balanced steady-state free precession sequence with a nominal phase progress (Φ) of 180 degrees per repetition time (TR) using the magnetic resonance apparatus (1), wherein in the balanced steady-state free precession sequence a white-marker gradient (GWM) is provided in order at least partially to balance a dephasing caused by a magnetic-field-changing object in the object under examination (4), - capture of image data of the object under examination (4) using the magnetic resonance apparatus (1) at an echo time (TE), characterised in that the method comprises the step: - adjustment of a phase development between a phase of a magnetisation of a first material (m1) and a phase of a magnetisation of a second material (m2), which form an interface in the object under examination (4), in the balanced steady-state free precession sequence using the magnetic resonance apparatus (1), wherein thanks to the adjustment of the phase development prior to the effect of the white-marker gradient (GWM) a co-phasal alignment of a magnetisation of the first material (m1) and of the second material (m2) at the interface is effected at the echo time (TE).

2. Method according to claim 1, characterised in that the adjustment of the phase development comprises setting a scanner frequency of the balanced steady-state free precession sequence, wherein the scanner frequency lies in a centre between a Larmor frequency of the first material (m1) (fLarmor,m1) and a Larmor frequency of the second material (m2) (fLarmor,m2).

3. Method according to claim 1, characterised in that the adjustment of the phase development comprises a provision of an additional phase progress (Φ) per repetition time (TR) of the balanced steady-state free precession sequence, wherein the following applies for the additional phase progress (Φ1) of the balanced steady-state free precession sequence: fLarmor , m 1 − fLarmor , m 2 / 2 * TR * 360 ∘ 4. Method according to one of the preceding claims, characterised in that the first material (m1) is fat and the second material (m2) is water.

5. Method according to one of claims 1 to 3, characterised in that the first material (m1) is silicone and the second material (m2) is water.

6. Magnetic resonance apparatus (1) for capturing magnetic resonance image data of an object under examination (4), wherein the magnetic resonance apparatus (1) is designed - to provide a balanced steady-state free precession sequence with a nominal phase progress (φ) of 180 degrees per repetition time (TR), - in the balanced steady-state free precession sequence to provide a white-marker gradient (GWM), in order at least partially to balance a dephasing caused by a magnetic-field-changing object in the object under examination (4), - to capture image data of the object under examination (4) at an echo time (TE), characterised in that the magnetic resonance apparatus (1) is designed - to adjust a phase development between a phase of a magnetisation of a first material (m1) and a phase of a magnetisation of a second material (m2), which form an interface in the object under examination (4), in the balanced steady-state free precession sequence, wherein the adjustment of the phase development prior to the effect of the white-marker gradient (GWM) effects a co-phasal alignment of a magnetisation of the first material (m1) and of the second material (m2) at the interface at the echo time (TE).

7. Computer program, which can be loaded directly into a memory of a control facility (2) of a magnetic resonance apparatus (1) according to claim 5, with program means in order to execute the steps of the method according to one of claims 1 to 4, if the program is executed in the control facility (2) of the magnetic resonance apparatus (1).

8. Electronically readable data storage medium with electronically readable control information stored thereon, which comprises at least one computer program according to claim 7 and is configured so that when the data storage medium is used in a control facility (2) of a magnetic resonance apparatus according to claim 6 it performs a method according to one of claims 1 to 5.