• 专利附录
  • 专利说明
  • 权利要求
  • 类似技术
  • 同族专利
  • 引用文献
  • 法律状态
  • 优先权
    • 专利摘要:
    Method and system for generating magnetic resonance signals by rapid rotation with magic angle of fields with spatial coding. The present invention preferably refers to a system for generating magnetic resonance signals on a sample under study, comprising: i) a first set of n magnetic elements (1) configured to generate an activatable magnetic field, and the They are arranged in such a way that the axis formed by the poles of each of said magnetic elements (1) is oriented towards the field of view (3) forming an angle {ze]m with an axis (5) substantially perpendicular to the plane in which the magnetic elements (1) are arranged; and ii) activation means configured to sequentially activate the magnetic fields (4) of the n magnetic elements (1), generating an effective magnetic field (4) rotating around the axis (5). The invention also relates to a method of generating magnetic resonance signals comprising the use of said system. (Machine-translation by Google Translate, not legally binding)
    公开号:ES2773333A1
    申请号:ES201930015
    申请日:2019-01-10
    公开日:2020-07-10
    发明作者:Otamendi Joseba Alonso；Conde Fernando Galve；Guisado José Miguel Algarin；Baviera José María Benlloch
    申请人:Consejo Superior de Investigaciones Cientificas CSIC；Universidad Politecnica de Valencia；
    IPC主号:
    专利说明:

    [0002] METHOD AND SYSTEM FOR THE GENERATION OF MAGNETIC RESONANCE SIGNALS BY RAPID ROTATION WITH MAGIC ANGLE OF FIELDS WITH SPACE CODING
    [0004] FIELD OF THE INVENTION
    [0006] The present invention relates, in general, to the field of nuclear magnetic resonance signal generation and imaging technologies obtained with said signals (known as MRI, from its English term "Magnetic Resonance Imaging") and, more particularly, to MRI techniques for hard and soft tissue imaging.
    [0008] BACKGROUND OF THE INVENTION
    [0010] Since its invention in the early 1970s, the impact of MRI techniques has been crucial in the healthcare sector, where they outperform any other known technique for soft tissue imaging. The main reason is that MRI is the only known technique capable of obtaining in vivo images of deep tissues with high spatial resolution, without the need in addition to the use of harmful ionizing radiation on said tissues to generate the images.
    [0012] Despite this unquestionable success, MRI of hard tissues (such as bones, cartilage or teeth) remains a problem that remains unsolved today. This is mainly due to the short useful life of magnetic resonance (MRI) signals emitted by solid bodies, in contrast to the case of signals from soft or non-solid tissues.
    [0014] MRI protocols are based on excitation and on the detection of the degree of freedom of spin of the nuclei of the samples on which the images are to be obtained. When these nuclei are subjected to an external magnetic field, they have a magnetic energy proportional to the field intensity and a dipole moment that tends to align with the lines of the external magnetic field. The temperature of the samples leads to magnetic dipole fluctuations, which constitute a noise environment for the surrounding spin vectors. As a result, the coherence of the spin value decays with a time constant (commonly referred to as T2 *), which it is an upper limit to the time during which information can be obtained from the sample, before being excited again by radio frequency (RF) radiation.
    [0016] These spin-spin interactions are predominantly of the dipole-dipole type, for which the intensity of the interaction largely depends on the angle Q formed between the line connecting the dipoles and the direction they point, becoming zero in known as " magic angle "( Q m = arccos (1 / V3)« 55 °). In non-solid samples, Q is continually changing, being an isotropic distribution on average, suppressing coupling between neighboring nuclei and leading to strong signals that MRI scanners use for image reconstruction. This averaging effect does not occur in solids, since both the nuclei and the magnetic field are static in the laboratory reference frame. Although the effects of the magic angle are commonly observed in crystalline and quasi-crystalline structures (such as tendons), they only occur in specific privileged orientations with respect to the main magnetic field and cannot be used for imaging of hard tissues, with character general.
    [0018] Magic Angle Spinning (MAS) can be used to suppress dipole-dipole interactions and improve the intensity and lifespan of MR signals emitted by solids. MAS is based on mechanically rotating the sample around an axis that is at the magic angle with respect to the main magnetic field, leading to an average of the Q angle, similar to the case of non-solid samples. For strong suppression, the sample must rotate at frequencies at least as fast as the dipole-bare dipole interaction rate (1 / T2 *). Typical T2 * values range from tens (eg, in teeth) to hundreds of microseconds (eg, in bone and cartilage), thus requiring impractical mechanical rotation frequencies from a few to 100 kHz.
    [0020] An alternative to conventional MAS is the “Magic Angle Rotating Frame” technique (or MARF, from its English term “Magic Angle Rotating Frame”), where the sample is static in the laboratory reference system, and the field magnetic is rotated at an angle Qm with respect to a fixed axis z. Currently, this type of technique is carried out through two variants: one is to rotate the main magnet that creates the required field for MRI and another is to dynamically change the direction of the field with alternating current (AC) fields. The former leads to the same mechanical limitations found in MAS (large objects rotating at tens of kHz) and Current approaches to the second are based on combinations of magnetic elements, where one of them produces a static field along the z-axis, and the others are controlled with time-modulated currents to tilt and rotate the overall field. While this technique is used routinely for solid-state nuclear magnetic resonance (NMR), the additional magnetic field heterogeneities required for NMR constitute a notable technical barrier. This, along with the high power required to install the main field (Q m is a large angle), makes current MARF approaches impractical for MRI as well.
    [0022] The anterior background explains the absence of MRI scanners capable of obtaining high-quality simultaneous images of soft and hard tissues. The present invention presents a solution to the problems of the state of the art mentioned above.
    [0024] BRIEF DESCRIPTION OF THE INVENTION
    [0026] The present invention addresses the aforementioned drawbacks, providing a system and a method to achieve rapid control of the spatial distribution of the magnetic field, generally based on the magic angle rotation of spatially non-homogeneous magnetic fields, or MASSIF (from its term in English "Magic Angle Spinning of Spacially Inhomogeneous Fields") although, in particular embodiments of the invention, said fields can also be configured locally homogeneously. The invention is therefore capable of obtaining images of soft and hard tissues simultaneously using a MARF technique and, advantageously, allows a fast time control (reaching frequencies above 100 kHz) of the direction of the magnetic field and the spatial distribution of the force over the field of view (or FOV, from its term in English "Field of View", corresponding to the region or volume over which the MR signals are generated and the images are obtained is corresponding).
    [0028] The invention is based mainly on a plurality of magnetic elements that are arranged around a perimeter, and are configured in such a way that the magnetic field generated by each of said elements in the field of view (FOV) points at an angle ~ Qm with respect to a static reference axis z, in the laboratory's reference frame. These magnetic elements can comprise, for example, coils arranged in order to directly generate the desired magnetic field, or they can be wound around electro-permanent magnets to control their magnetization, and which in turn generate the desired magnetic fields.
    [0029] By connecting said magnetic elements to one or more electronic modules, it is possible to generate the time-varying currents that pass through the coils. In this way, control over time of the intensity and direction of the magnetic field generated in the FOV is achieved, which is the main requirement for the MASSIF technique.
    [0031] More specifically, a first object of the invention refers to a system for generating magnetic resonance signals on a sample under study, which advantageously comprises:
    [0032] - a first set of n magnetic elements, configured to generate an activatable magnetic field on the sample under study in a field of view;
    [0033] - activation means configured to sequentially activate the magnetic fields of the n magnetic elements, generating an effective magnetic field of revolution around an axis;
    [0034] and where the magnetic elements are arranged so that the axis formed by the poles of each of said magnetic elements is oriented towards the field of view at an angle dm with the axis of revolution of the effective magnetic field on the sample.
    [0036] The invention can be configured to generate both homogeneous instantaneous magnetic fields (allowing the system to be used for MR / MRS spectroscopy), and to generate homogeneous magnetic fields (suitable for MRI techniques), by installing additional magnetic elements (typically called "coils gradient "in the MRI community) to generate inhomogeneous fields that encode spatial information at the spin precession frequency. If the intensity of the rotating magnetic field is not homogeneous, it makes dedicated gradient fields (and their corresponding hardware) unnecessary. This enables the reconstruction of three-dimensional images without additional requirements, qualitatively extending the performance of state-of-the-art rotating spatial encoding magnetic field (rSEM) configurations.
    [0038] In a preferred embodiment of the invention, as mentioned, the angle dm has a value substantially of arcs (1 / V3), corresponding to the value of the magic angle.
    [0039] In another preferred embodiment of the invention, the magnetic elements are located at a distance / or from the center of the field of vision of said system, on a circumference whose plane is substantially perpendicular to the axis of revolution of the effective magnetic field.
    [0041] In another preferred embodiment of the invention, the n magnetic elements are arranged, angularly, in a substantially equispaced manner in the path of revolution of the effective magnetic field.
    [0043] In another preferred embodiment of the invention, the magnetic elements comprise solenoids, conductive flat coils or electromagnets subjected to electric currents of magnetic induction.
    [0045] In another preferred embodiment of the invention, the activation means comprise pulsed electric current generators, connected to the magnetic elements.
    [0047] In another preferred embodiment of the invention, the system comprises a second set of additional magnetic elements, which can be activated by the activation means and arranged in the system in addition to the arrangement of the first set of magnetic elements, so that both sets generate a magnetic field homogeneous over the field of view when activated by the activation means.
    [0049] In another preferred embodiment of the invention, the system additionally comprises one or more means of excitation by radio frequency of the sample under study.
    [0051] In another preferred embodiment of the invention, the means for activating the magnetic elements comprise a control computer integrated with a low-power electronics module, connected to the magnetic elements through a high-power electronics module.
    [0053] In another preferred embodiment of the invention, the high power electronics module comprises a power supply and / or an analog power amplifier, configured to communicate with the control computer in cooperation with the low power electronics module.
    [0055] A second object of the invention refers to a method of generating magnetic resonance signals on a sample under study, which comprises the use of a system according to any of the preceding claims to generate an activation sequence of the magnetic fields of the first set of n magnetic elements, generating an effective rotating magnetic field around the axis.
    [0057] In a preferred embodiment of the method of the invention, a single magnetic element is activated at a time from among the n that make up the first set, for a period of time of duration t = Tr / n = 2 n / (nwr), where Tr is the total period of revolution of the global magnetic field of the system, and wr is the corresponding angular frequency.
    [0059] In another preferred embodiment of the method of the invention, a plurality of magnetic elements are activated at the same time, by means of alternating current generators.
    [0061] In another preferred embodiment of the method of the invention, it also comprises the detection of the magnetic resonance signals at time s ( t) generated in the sample under study in the field of view, and the analysis of the spin density distribution p (r) and the magnetic field generated on said sample, when it is subjected to radio frequency excitation.
    [0063] In another preferred embodiment of the method of the invention, the detected time signal s (t) is related to the spin density distribution p (r) through a coding function ® ( r, t) that calculates the phase of acquired magnetic resonance for each point in space r, for each instant in time t:
    [0067] with:
    [0072] where wl is the Larmor precession frequency, and is the gyromagnetic factor and B ( r, t ') the magnetic field at a given position r and time t'.
    [0074] The foregoing and other aspects and advantages of the invention will be described in greater detail in subsequent sections herein. In said description, reference will be made to a series of drawings which are shown by way of example of preferred embodiments of the invention. However, these embodiments do not necessarily represent the full scope of the invention, which will be delimited by the claims and their interpretation by one of ordinary skill in the art in light of the detailed description herein.
    [0075] Throughout the present description, the expressions "substantially" "substantially equal" or "substantially of" shall be understood to refer to the value of a given quantity or a certain property, as equal to said value or property, or included in a variation range of ± 10% with respect to them.
    [0077] DESCRIPTION OF THE DRAWINGS
    [0079] For a more complete understanding of the invention, reference is made to the following description and accompanying drawings, in which:
    [0081] Figure 1 shows a diagram with a general view of two possible embodiments of the invention, depending on the distribution of their magnetic elements.
    [0083] Figure 2 shows the magnetic field and the proton Larmor frequency spatial distributions in three Cartesian planes, for a preferred embodiment of the invention at t = 0.
    [0085] Figure 3 shows the magnetic field and the proton Larmor frequency distributions in the lines along the magic angle for a preferred embodiment of the invention at the instants when the coils used as magnetic elements of the system are pulsed. said invention.
    [0087] Figure 4 shows the angle Q between the magnetic field and the z axis in the z = 0 plane for a preferred embodiment of the invention at t = 0.
    [0089] Figure 5 shows a block diagram with the main components of the hardware of the system of the invention, in a preferred embodiment thereof, and a possible sequence of current pulses through the individual coils for the MASSIF technique.
    [0091] Figure 6 shows a homogeneous spherical proton distribution (Fig. 6a), a discrete version of the original 512 voxel distribution (Fig. 6b) and the reconstruction of the resulting image by the method of the invention in a preferred embodiment of the same, after 100 ps (Fig. 6c).
    [0092] Figure 7 shows the induced MR signal for a homogeneous spherical proton distribution and a preferred embodiment of the system of the invention, including the resulting image reconstructions after acquisition times of 10, 30 and 100 ps.
    [0094] Figure 8 shows a chain of three homogeneous spherical proton distributions (Fig. 8a), a discrete version of the original 512 voxel distribution (Fig. 8b) and the reconstruction of the resulting image after 100 ps, using the system of the invention, in a preferred embodiment thereof (Fig. 8c).
    [0096] Figure 9 shows the MR signal induced for a chain of three homogeneous spherical proton distributions with the system of the invention, in a preferred embodiment thereof.
    [0098] Figure 10 shows the reconstruction of the image resulting from the sample in FIG. 6 after 100 ps with the system of the invention, in a preferred embodiment thereof and at half the rotation frequency ( wr = 2n50 kHz).
    [0100] Figure 11 shows the MR signal induced for a homogeneous spherical proton distribution with the system of the invention, in a preferred embodiment thereof and at half the rotation frequency ( wr = 2n50 kHz).
    [0102] Figure 12 shows the comparison between the reconstruction of the image resulting from the sample in Figure 6, after 100 ps with the system of the invention, in a preferred embodiment of the same and at half the rotation frequency ( wr = 2n50 kHz).
    [0104] DETAILED DESCRIPTION OF THE INVENTION
    [0106] Different preferred embodiments of the invention are described below, based on Figures 1-12 herein.
    [0108] Figure 1 shows an arrangement of a preferred embodiment of the system of the invention, which is presented by way of example and, therefore, for non-limiting purposes thereof. In said Figure 1, a plurality of n magnetic elements (1) can be seen, preferably arranged around a circumference (although other configurations are also possible within the scope of the invention). When a Electric current is pulsed through a magnetic element (1), it becomes an active magnetic element (2). In this configuration, each magnetic element (1) is preferably positioned at a distance / or from the center of the field of view (3) (FOV) and is oriented in such a way as to generate a magnetic field (4) whose direction in the FOV (3) forms an angle dm with a z axis (5), which is static in the laboratory frame of reference. The coordinates of each i-th magnetic element are: / o {sin (n-0m) ■ cos (0 /), sin (nQ m) sin (0 /), cos (n-Qm)}, with 0 / = (/ -1) -2n / n. An optional array of additional magnetic elements (6) can be added when spatial homogeneity of the magnetic field over the FOV (3) is required. In most of the embodiments considered below, this will be omitted to avoid the need for the use of dedicated gradient coils for imaging purposes.
    [0110] There are different options for the realization of the magnetic elements (1). For example, they can take the form of a solenoid, a flat coil or an electro-permanent magnet (or EPM, from its English term “Electro-permanent Magnet”), which can be pulsed by passing an electric current through a coil wound around it.
    [0112] Next, a particular configuration of the invention is described, which will be referred to hereinafter as "Embodiment 1", which is provided for illustrative purposes and on which most of the calculations and simulations included in the present description are based. Said Embodiment 1 comprises n = 4 magnetic elements (1) that take the form of solenoids with 10 windings each, which are wound around cylinders with radii r = 3.47 cm. In said Embodiment 1, the length l of the solenoid is much less than r and ro, so it can be approximated to l = 0 in the calculations that follow. The solenoids are placed with their centers at a distance / 0 = 5 cm from the center of the FOV (3), with their axes of symmetry forming angles Q m with the z axis. The FOV (3) is cubic, with sides I fov = 1 cm. Each active magnetic element (2) carries an electric current I 0 = 500 A.
    [0114] Figures 2-3 show the distribution of the intensity of the spatial magnetic field for Embodiment 1 of the invention, when the only active magnetic element (2) is i = 1. The proton Larmor frequencies are also represented in the figures . The intensity of the magnetic field (4) generated by a single magnetic element (1) takes values of 13-21 mT on the FOV (3), corresponding to proton Larmor frequencies of 500-900 kHz. The distribution of angles between magnetic field (4) and the z axis (5) in the z = 0 plane are shown in Figure 4 and are always within a 3 ° variation margin with respect to Qm.
    [0116] The elements shown in Figure 1 thus form the essence of the disclosed invention. However, in order to apply it to MR techniques, spin excitation and detection capabilities must be added. This can be done, for example, with standard radio frequency (RF) elements, tuned to the Larmor precession resonance frequency of the nuclei under study. For imaging samples, a typical pulse sequence begins with radiofrequency excitation of nuclei, transforming longitudinal magnetization to detectable transverse magnetization.
    [0118] In the system of Figure 1, RF excitation can be carried out with a single active magnetic element (2), or with multiple activated magnetic elements (1), which produces a larger magnetic field (4), which pre -polarize the sample to obtain an improved Signal to Noise Ratio (SNR).
    [0120] Instead of using a conventional MRI sequence after the RF excitation pulse, the MASSIF technique of the invention relies on the rapid control of electrical currents to rotate the global magnetic field (4) around the z- axis (5). Figure 5 shows a block diagram with the main components of the hardware of the system of the invention, together with a possible sequence of current pulses through a plurality of individual coils to implement the imaging method of the invention. The hardware components include, for example, a control computer (7), preferably integrated with a low-power electronics module (8) and a high-power electronics module (9). Figure 5 shows an embodiment in which the control computer (7) and the low-power electronics (8) are combined. Typically the high power electronics (9) will be a separate module. If it is a power supply, preferably it will support digital communication, in which case the control computer (7) sends digital information about the sequence of pulses that must be programmed in an internal memory of the high-power electronics (9) , and the low power electronics (8) send the triggers to generate the high power pulses. If the high power electronics (9) is an analog power amplifier, the control computer (7) sends digital information about the sequence of pulses that will be programmed into an internal memory of the low power electronics (8), and the low power electronics (8) sends the equations of the low-power pulses that are amplified in high-power electronics (9).
    [0122] Multiple sequences of pulses can be devised for rapid control of the magnetic field (4). One possibility is to have a single active magnetic element (2) at a time from among the n that make up the total set, during a time window of duration ¿coil = Tr / n = 2 n / (nwr), where Tr is the total period of revolution of the system, and Wr is the corresponding angular frequency. This is depicted in Figure 5 and is used for all calculations below. Another possibility is to use a more complex time control, for example, activating all the magnetic elements (1) simultaneously with alternating current (AC) generators, in the form lo • cos (Wrt%), with % = ( i - 1 ) n / n, for magnetic elements (1) that run from i = 1 to n.
    [0124] In addition to the system described in the preceding paragraphs, the present invention also refers to a method for implementing MASSIF with said system, for the generation of MR images of the objects under study. Unlike typical MRI embodiments, magnetic field inhomogeneities in MASSIF can vary on timescales significantly shorter than the signal acquisition window. Likewise, as in the case of spatial coding magnetic field (rSEM) MRI techniques, the Fourier transformations directly obtained from the time-dependent signals acquired in the RF detector do not produce an image of the sample, since the Magnetic field homogeneities are not described by simple linear gradients. Instead, the detected time signal s (t) is related to the spin density distribution p (r) through a coding function ® ( r, t) that calculates the acquired MR phase for each point in space , for each instant in time:
    [0125] s ( t) = f e A>[F'1'p ( r) dr.
    [0126] 7 fo V (Eq. 1) with:
    [0128]
    [0130] where wl is the Larmor precession frequency, and the gyromagnetic ratio ( ~ 2n ■ 42 MHz / T for protons), and B ( r, t ') the magnetic field at a given position r and time t'.
    [0131] When discretized, s (t) becomes a vector S of length equal to the number of time steps tn, p ( r) becomes a vector p of length equal to the number of voxels Vn, and ® ( r, t ) is converted to an OR matrix with tn rows and Vn columns, referred to as a "coding matrix". After discretization, Equation 1 becomes:
    [0132] S = Op, (Eq. 3) and p can be obtained as O _1 S , from where an image can be unequivocally retrieved.
    [0134] All image reconstructions below are based on Embodiment 1, with the following parameters, unless otherwise specified:
    [0135] • Wr = 2 rn-100 kHz (Tr = 10 us, enough for the narrowing of magic angle lines for the hardest biological tissues).
    [0136] • Total acquisition time tacq = 100 ys.
    [0137] • tn = 1000 time steps (resolution time 5t = 100 ns).
    [0138] • Vn = 512 voxels (spatial resolution 5r = 1.25 mm).
    [0139] • The pulse sequence represented in Figure 5.
    [0141] It will also be assumed that the detection of the signal is homogeneous throughout the FOV (3), regardless of the instantaneous local orientation of the magnetic field (4). The small variations of Q visible in Figure 4 for these conditions indicate that this approximation is reasonable if a single detector coaxial with the z- axis is used (5).
    [0143] A sample whose spins are homogeneously distributed over the volume of a sphere is considered below. This corresponds to the graph in Figure 6a. The second graph (Figure 6b) is the discretized version of the original distribution (512 voxels). The discretized time signal S induced by the sample in an ideal detector, as described above, is represented in the upper left graph of Figure 7. In order to reconstruct an image from it, the matrix O is calculated with Equation 2. Subsequently, using Equation 3, the p values are solved, from which the reconstruction corresponding to Figure 6c is obtained.
    [0145] The information obtained from the sampled object depends, to a large extent, on the total acquisition time tacq. The 3D graphs in the lower part of Figure 7 show the reconstructions following the procedure described in the previous paragraph, but for tacq = 10 and 30 ys. As you can see, the difference between the two is little. This is related to the scarce information (that is, the amplitude of the signal) in the time interval between 10 and 30 ys, which in turn is related to the frequency of rotation of the magnetic field Wr. This already suggests an important aspect of MASSIF on which we detail below: the fast rotation frequencies of MASSIF allow a much faster image reconstruction than other MRI techniques.
    [0147] The results of the above calculations applied to a second less symmetrical object are presented in Figures 8-9. The object photographed in this case is a chain of three small spherical spin distributions centered at the center and two at the corners of the FOV (3).
    [0149] High quality MR images can be obtained from MRI scanners in which the evolution of the phase changes significantly within the dimensions of the sample. For this reason, to obtain high resolution MR images, conventional scanners use strong magnetic gradients. In MASSIF, however, there is an additional degree of freedom that can be used to control spatial resolution: the rotation frequency wr. The MR phase acquired at each point in space depends not only on its position r (as is the case with conventional NMR), but also on Wr. This is evidenced in Equation 2, where the explicit time dependence of B ( r, t) is solely due to the fact that the magnetic field is rotating.
    [0151] The influence of Wr on the reconstructed image is clearly seen in Figure 10, which shows the reconstruction of the image of the sample obtained in Figure 6 after 100 ys with Run 1 at half the rotation frequency (Wr = 2n 50 kHz). The time signal S is represented in Figure 11, while Figure 12 makes a direct comparison of the results of the images at 50 and 100 kHz.
    权利要求:
    Claims (15)
    [1]
    1 System for generating magnetic resonance signals on a sample under study,
    characterized by comprising:
    - a first set of n magnetic elements (1), configured to generate an activatable magnetic field on the sample under study in a field of view (3);
    - activation means configured to activate sequentially the magnetic fields (4) of the n magnetic elements (1), generating an effective magnetic field of revolution around an axis (5);
    and where the magnetic elements (1) are arranged so that the axis formed by the poles of each of said magnetic elements (1) is oriented towards the field of view (3) forming an angle dm with the axis (5) of revolution of the effective magnetic field on the sample.
    [2]
    two.
    [3]
    3. - System according to any of the preceding claims, where the magnetic elements (1) are located at a distance / or from the center of the field of vision (3) of said system, on a circumference whose plane is substantially perpendicular to the axis (5 ) of revolution of the effective magnetic field.
    [4]
    Four.
    [5]
    5. - System according to any of the preceding claims, wherein the magnetic elements (1) comprise solenoids, flat conductive coils and / or electromagnets subjected to magnetic induction electric currents.
    [6]
    6. - System according to any of the preceding claims, wherein the activation means comprise pulsed electric current generators, connected to the magnetic elements (1).
    [7]
    7. - System according to any of the preceding claims, comprising a second set of additional magnetic elements (6), activatable by the activation means and arranged in the system complementary to the arrangement of the first set of magnetic elements (1), of so that both sets (1, 6) generate a homogeneous magnetic field over the field of view (3) when activated by the activation means.
    [8]
    8.
    [9]
    9.
    [10]
    10.
    [11]
    eleven.
    [12]
    12. - Method according to the preceding claim, where a single magnetic element (2) is activated at a time from among the n that make up the first set, for a period of time of duration t = Tr / n = 2 n / (nwr ), where Tr is the total period of revolution of the global magnetic field (4) of the system, and wr is the corresponding angular frequency.
    [13]
    13.
    [14]
    14. - Method according to any of claims 11-13, which also comprises the detection of the magnetic resonance signals at time s ( t) generated in the sample under study in the field of view (3), and the analysis of the spin density distribution p (r) and the magnetic field (4) generated on said sample, when it is subjected to radiofrequency excitation.
    [15]
    fifteen. Method according to the preceding claim, where the detected time signal s (t) is related to the spin density distribution p (r) through a coding function ® ( r, t) that calculates the acquired magnetic resonance phase for each point in space r, for each instant in time t:
    s ( t ) = í e A>[F'1'p ( r) dr.
    J F or V
    with:

    where wl is the Larmor precession frequency, and the gyromagnetic relationship and B ( r, t ') the magnetic field at a given position r and time t'.
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    引用文献:
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    EP2762071A1|2011-09-28|2014-08-06|National University Corporation Kumamoto University|Image analysis device, image analysis method, and image analysis programme|
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