WO2020144389A2

WO2020144389A2 - Method and system for the generation of magnetic resonance signals using rapid rotation with magic-angle fields and spatial encoding

Info

Publication number: WO2020144389A2
Application number: PCT/ES2020/070010
Authority: WO
Inventors: Joseba ALONSO OTAMENDI; Fernando GALVE CONDE; José Miguel ALGARÍN GUISADO; José María BENLLOCH BAVIERA
Original assignee: Universitat Politècnica De València; Consejo Superior De Investigaciones Científicas (Csic)
Priority date: 2019-01-10
Filing date: 2020-01-09
Publication date: 2020-07-16
Also published as: WO2020144389A3; ES2773333A1; ES2773333B2

Abstract

The present invention relates, preferably, to a system for the generation of magnetic resonance signals on a sample under study, comprising: i) a first array of n magnetic elements (1) configured to generate an activatable magnetic field, the same being disposed in such a way that the axis formed by the poles of each of said magnetic elements (1) is oriented toward the field of vision (3) forming an angle θ_m with an axis (5) substantially perpendicular to the plane whereon the magnetic elements (1) are disposed; and ii) activation means configured to sequentially activate the magnetic fields (4) of the n magnetic elements (1), generating an effective rotating magnetic field (4) around the axis (5). The invention likewise relates to a method for the generation of magnetic resonance signals comprising the use of the aforementioned system.

Description

DESCRIPTION

METHOD AND SYSTEM OF GENERATION OF MAGNETIC RESONANCE SIGNALS BY QUICK ROTATION WITH MAGIC ANGLE OF FIELDS WITH SPACE CODING

FIELD OF THE INVENTION

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 obtaining images of hard and soft tissues.

BACKGROUND OF THE INVENTION

Since their invention in the early 1970s, the impact of MRI techniques has been crucial in the healthcare sector, where they outperform any other known soft tissue imaging technique. 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 for the use of harmful ionizing radiation on said tissues to generate the images.

Despite this unquestionable success, MRI of hard tissues (such as bone, cartilage, or teeth) remains a problem that remains unresolved today. This is mainly due to the short lifespan of magnetic resonance (MR) signals emitted by solid bodies, unlike the case of soft or non-solid tissue signals.

MRI protocols are based on the excitation and 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 strength and a dipole moment that tends to align with the lines of the external magnetic field. The temperature of the samples leads to fluctuations of the magnetic dipole type, 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 re-excited by radio frequency (RF) radiation.

These spin-spin interactions are predominantly of the dipole-dipole type, for which the intensity of the interaction depends largely on the angle Q formed between the line connecting the dipoles and the direction in which they point, becoming zero in known as " magic angle "(QM = arccos (1 / V3)« 55 °). In non-solid samples, Q is continuously changing, on average being an isotropic distribution, suppressing coupling between neighboring nuclei, and leading to strong signals that MRI scanners use for image reconstruction. This averaging effect does not take place in solids, since both the nuclei and the magnetic field are static in the laboratory reference system. Although the effects of the magic angle are commonly observed in crystalline and quasi-crystalline structures (such as tendons), they only take place in specific privileged orientations with respect to the main magnetic field and cannot be used for imaging of hard tissues, with character general.

Magic Angle Spinning (MAS) can be used to suppress dipole-dipole interactions and improve the intensity and lifespan of RM 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 naked dipole-dipole interaction rate (1 / T2 ^* ). Typical T2 ^* values range from tens (for example, in the teeth) to hundreds of microseconds (for example, in bone and cartilage), so they require impractical mechanical rotation frequencies from a few to 100 kHz.

An alternative to conventional MAS is the “magic angle rotation frame” (or MARF) technique, 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 performed 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 first 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 general field. While this technique is routinely employed for solid state nuclear magnetic resonance imaging (NMR), the additional magnetic field heterogeneities required for NMR constitute a notable technical barrier. This, coupled 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.

The background above explains the absence of MRI scanners capable of obtaining high-quality simultaneous images of soft and hard tissues. The present invention proposes a solution to the problems of the state of the art mentioned above.

BRIEF DESCRIPTION OF THE INVENTION

The present invention addresses the aforementioned drawbacks by providing a system and method for achieving rapid control of the spatial distribution of the magnetic field, generally based on the magical angle rotation of spatially non-homogeneous magnetic fields, or MASSIF (of its English term "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 rapid time control (reaching frequencies above 100 kHz) of the direction of the magnetic field and the distribution spatial force on the field of view (or FOV, from its English term "Field of View", corresponding to the region or volume on which the MR signals are generated and the corresponding images are obtained).

The invention is based, mainly, on a plurality of magnetic elements that are arranged around a perimeter, and are configured so that the magnetic field generated by each of these elements in the field of view (FOV) points at an angle ~ Q _M with respect to a static reference axis z, in the laboratory reference frame. These magnetic elements may comprise, for example, coils arranged to directly generate the desired magnetic field, or they may be wound around electro-permanent magnets to control the magnetization of them, and which in turn generate the desired magnetic fields. By connecting said magnetic elements to one or more electronic modules, it is possible to generate the variable currents in the time they 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.

More specifically, a first object of the invention relates to a system for generating magnetic resonance signals on a sample under study, which advantageously comprises:

- a first set of n magnetic elements, configured to generate an activatable magnetic field on the sample under study in a field of view;

- activation means configured to sequentially activate the magnetic fields of the n magnetic elements, generating an effective magnetic field of revolution about an axis;

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 forming an angle 0 _m with the axis of revolution of the effective magnetic field on the sample.

Within the scope of the invention, there are different options for the realization of the magnetic elements. For example, they can take the form of a solenoid, a flat coil, a permanent magnet, or an electro-permanent magnet (or EPM), which can be pulsed by passing an electric current. through a coil wound around it. Likewise, although in this document the generic expression “poles” will be used to refer to the orientation of said magnetic elements in space, this expression is to be understood, without limitation, as relative to the main axis of polarity of the magnetic field of said elements, regardless of the specific implementation they adopt.

The invention can be configured to generate both spatially homogeneous magnetic fields (allowing the system to be used for MR / MRS spectroscopy), and to generate inhomogeneous magnetic fields (suitable for MRI techniques), for example, by installing additional magnetic elements (typically called "gradient coils" in the MRI community) to generate inhomogeneous fields that encode spatial information at the frequency and precession phase spin. If the intensity of the rotating magnetic field is not homogeneous, it makes the dedicated gradient fields (and their corresponding hardware) unnecessary. This allows the reconstruction of three-dimensional images without additional requirements, which qualitatively expands the performance of the state-of-the-art rotary spatial encoding magnetic field (rSEM) configurations.

In a preferred embodiment of the invention, as mentioned, the angle 0 _m has a value of substantially arccos (1 / V3), corresponding to the value of the magic angle.

In another preferred embodiment of the invention, the magnetic elements are located at a distance / ¾ 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.

In another preferred embodiment of the invention, the n magnetic elements are arranged, angularly, substantially equally spaced in the effective magnetic field revolution path.

In another preferred embodiment of the invention, the magnetic elements comprise solenoids, flat conductive coils, or electromagnets subjected to electric currents of magnetic induction.

In another preferred embodiment of the invention, the activation means comprise pulsed electric current generators, connected to the magnetic elements.

In another preferred embodiment of the invention, the system comprises a second set of additional magnetic elements, activatable 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.

In another preferred embodiment of the invention, the system additionally comprises one or more means of radiofrequency excitation of the sample under study.

In another preferred embodiment of the invention, the activation means of the magnetic elements comprise a control computer integrated with an electronics module. Low power, connected to the magnetic elements through a high power electronics module.

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.

In another preferred embodiment of the invention, the generation system is part of an MR kit and, more preferably, of an MR medical dental imaging kit. The invention is especially suitable for this type of application, given its capacity for implementation in devices of reduced dimensions.

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 previous claims to generate a sequential activation of the magnetic fields of the first set of n magnetic elements, generating an effective magnetic field rotating around the axis.

In a preferred embodiment of the method of the invention, a single magnetic element is activated at a time among the n that make up the first set, for a period of time of duration t = T _r / n = 2 p / (hw), where T _r is the total period of revolution of the global magnetic field of the system, and w _G is the corresponding angular frequency.

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.

In another preferred embodiment of the method of the invention, it also comprises detecting the magnetic resonance signals at time s (t) generated in the sample under study in the field of view, and analyzing the spin density distribution p (r) and the magnetic field generated on said sample, when it is subjected to radiofrequency excitation.

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 an encoding function F (r, t) that calculates the resonance phase acquired magnetic for each point in space r, for each instant in time t

with:

where O _L is the Larmor precession frequency, and is the gyro-magnetic factor, B (r, t ') the magnetic field at a given position r and an instant /', and m (r) is a unit vector in the direction of magnetization in r.

The foregoing and other aspects and advantages of the invention will be described in greater detail in subsequent sections of this document. 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 by their interpretation by a person skilled in the art in light of the detailed description of the present document.

Throughout the present description, the expressions, “substantially” “substantially equal” or “substantially of”, referring to the value of a given quantity or to a certain property, will be understood as equal to said value or property, or included in a variation range of ± 10% with respect to them.

DESCRIPTION OF THE DRAWINGS

For a more complete understanding of the invention, reference is made to the following description and the accompanying drawings, in which:

Figure 1 shows a diagram with an overview of two possible embodiments of the invention, depending on the distribution of its magnetic elements.

Figure 2 shows the magnetic field and the spatial frequency distributions of the proton Larmor in three Cartesian planes, for a preferred embodiment of the invention at / = 0.

Figure 3 shows the magnetic field and the proton Larmor frequency distributions on the lines along the magic angle for a preferred embodiment of the invention in the instants in which the coils that are used as magnetic elements of the system of said invention are pulsed.

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.

Figure 5 shows a block diagram with the main hardware components 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.

Figure 6 shows a homogeneous spherical distribution of protons (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).

Figure 7 shows the induced MR signal for a homogeneous spherical distribution of protons and a preferred embodiment of the system of the invention, including the resulting image reconstructions after acquisition times of 10, 30 and 100 ps.

Figure 8 shows a chain of three homogeneous spherical distributions of protons (Fig. 8a), a discrete version of the original 512 voxel distribution (Fig. 8b), and the resulting image reconstruction after 100 ps, using the the invention, in a preferred embodiment thereof (Fig. 8c).

Figure 9 shows the induced MR signal for a chain of three homogeneous spherical distributions of protons with the system of the invention, in a preferred embodiment thereof.

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 (w _G = 2tt · 50 kHz).

Figure 11 shows the MR signal induced for a homogeneous spherical distribution of protons with the system of the invention, in a preferred embodiment thereof and at half the rotation frequency (w _G = 2tt · 50 kHz). 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 thereof and at half the rotation frequency (w _G = 2tt · 50 kHz).

DETAILED DESCRIPTION OF THE INVENTION

Different preferred embodiments of the invention are described below, based on Figures 1-12 herein.

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) is appreciated, preferably arranged around a circumference (although other configurations are equally possible within the scope of the invention). When an 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 placed at a distance / ¾ from the center of the field of view (3) (FOV) and is oriented in such a way that it generates a magnetic field (4) whose direction in the FOV (3) forms an angle 0 _m with a z-axis (5), which is static in the laboratory reference frame. The coordinates of each i-th magnetic element are: / ¾ {sin (TT- @ _m ) eos ( _/ ), sin (TT-0 _m ) sin ( _/ ), cos (TT- @ _m )}, with, = (/ - 1) · 2tt / h. An optional array of additional magnetic elements (6) can be added when spatial homogeneity of the magnetic field on the FOV (3) is required. In most of the embodiments considered below, this will be omitted to avoid the need to use dedicated gradient coils for imaging purposes.

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 electrical current through a coil wound around it.

In the following, 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 the majority of the calculations and simulations included herein are based description. 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 / of the solenoid is much less than r and / ¾, so it can be approximated to / = 0 in the calculations that follow. The solenoids are placed with their centers at a distance / ¾ = 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 / F _OV = 1 cm. Each active magnetic element (2) carries an electric current k = 500 A.

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 the magnetic field (4) and the z-axis (5) in the z = 0 plane are shown in Figure 4 and are always within a variation range of 3 ^or with respect to 0 _m .

The elements shown in Figure 1 thus form the essence of the disclosed invention. However, in order to be applied 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 of samples, a typical pulse sequence begins with radio frequency excitation of the nuclei, transforming the longitudinal magnetization as a detectable transverse magnetization.

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), than pre -polarize the sample to obtain an improved Signal to Noise Ratio (SNR). Optionally, a dedicated pre-bias configuration can also be integrated. In both cases, the sample is thermalized with the pre-polarization field and the magnetization grows accordingly, after which the field is turned off and the RF pulses are executed according to a programmed sequence. In this way, it is possible to increase the low SNR inherent in weak magnetic fields, provided that the The sequence is carried out before the magnetization rethermalizes with the low field, which occurs with a typically long time constant.

Rather than using a conventional MR sequence after the RF excitation pulse, the MASSIF technique of the invention relies on 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 hardware components 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. 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, it will preferably support digital communication, in which case the control computer (7) sends digital information on the pulse sequence 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 on the pulse sequence to be programmed in 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 the high power electronics (9).

Multiple pulse sequences can be devised for rapid control of the magnetic field (4). One possibility is to have a single active magnetic element (2) at the same time among the n that make up the total set, during a time window of duration / _coil = T _r / n = 2 p / (hw _G ), where T _r is the total period of revolution of the system, and ü) _r is the corresponding angular frequency. This is represented in Figure 5 and 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 k eos (or _r t + fί), with f, = ( / - 1) p / h, for magnetic elements (1) that run from i = 1 to n.

In addition to the system described in the preceding paragraphs, the present invention also relates to a method for implementing MASSIF with said system, for the generation of MR images of the objects under study. Unlike typical MRI embodiments, inhomogeneous magnetic fields employed in MASSIF can vary on significantly shorter time scales than the signal acquisition window. Likewise, as in spatially encoded magnetic field MRI (rSEM) techniques, Fourier transformations directly obtained from the time-dependent signals acquired in the RF detector do not necessarily produce an image of the sample, since the inhomogeneities of the magnetic field are not described by simple linear gradients. Instead, the detected time signal s (f) is related to the spin density distribution p (r) through a coding function F (r, t) that calculates the acquired MR phase for each point in space , for each instant in time:

(Eq. 1) with:

where w _\ _ is the Larmor precession frequency, and the gyro-magnetic ratio (~ 2p 42 MHz / T for protons), B (r, t ') the magnetic field at a given position r and an instant t', and m (r ) is a unit vector in the direction of magnetization at r.

When discretized, s (f) becomes a vector S of length equal to the number of time steps t _n , p (r) becomes a vector p of length equal to the number of voxels v _n, and F (r , t) becomes an F matrix with t _n rows and v _n columns, referred to as the "coding matrix". After discretization, Equation 1 becomes:

S = O p, (Eq. 3) and p can be obtained as F ¹ S, from where an image can be unequivocally retrieved.

All image reconstructions below are based on Embodiment 1, with the following parameters, unless otherwise specified:

• w _G = 2 77-100 kHz (T _r = 10 ps, sufficient for the narrowing of magic angle lines for the hardest biological tissues).

• Total acquisition time f _aCq = 100 ps.

• t _n = 1000 steps of time (resolution time 5t = 100 ns).

• v _n = 512 voxels (spatial resolution <5r = 1.25 mm). The pulse sequence represented in Figure 5.

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 in Q visible in Figure 4 for these conditions indicate that this approximation is reasonable if a single coaxial detector with the z axis is used (5).

Next, we consider a sample whose spins are homogeneously distributed over the volume of a sphere. 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 F is calculated with Equation 2. Subsequently, using Equation 3, the p-values are solved, from which the reconstruction corresponding to Figure 6c is obtained.

The information obtained from the sampled object depends, to a large extent, on the total acquisition time f _acq . The 3D graphics at the bottom of Figure 7 show the reconstructions following the procedure described in the previous paragraph, but for f _acq = 10 and 30 ps. As you can see, the difference between the two is small. This is related to sparse information (i.e. signal amplitude) in the time interval between 10 and 30 ps, which in turn is related to the frequency of rotation of the magnetic field w _G. This already suggests an important aspect of MASSIF about which we detail below: the fast rotation frequencies of MASSIF allow a much faster image reconstruction than other MRI techniques.

The results of the above calculations applied to a second less symmetrical object are presented in Figures 8-9. The object in this case is a chain of three small spherical spin distributions centered in the center and two of the corners of the FOV (3).

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, scanners Conventionals 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 w _G. The phase of MR acquired at each point in space depends not only on its position r (as is the case with conventional NMR), but also on w _G. This is evidenced in Equation 2, where the explicit dependence in time of B (r, t is solely due to the fact that the magnetic field is rotating.

The influence of w _G on the reconstructed image is clearly seen in Figure 10, which shows the reconstruction of the sample image obtained in Figure 6 after 100 me with Embodiment 1 at half the rotation frequency (w _G = 2tt · 50 kHz). The time signal S is represented in Figure 11, while Figure 12 makes a direct comparison of the image results at 50 and 100 kHz.

Claims

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 sequentially activate the magnetic fields (4) of the n magnetic elements (1), generating an effective magnetic field of revolution about 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 0 _m with the axis (5) of revolution of the effective magnetic field on the sample.

2. System according to the preceding claim, where 0 _m has a value of substantially arccos (1 / V3).

3. System according to any of the preceding claims, wherein the magnetic elements (1) are located at a distance ro from the center of the field of view (3) of said system, on a circumference whose plane is substantially perpendicular to the axis (5) of revolution of the effective magnetic field.

4 - System according to any of the preceding claims, wherein the n magnetic elements (1) are arranged, angularly, in a substantially equispaced way in the path of revolution of the effective magnetic field.

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. System according to any of the preceding claims, wherein the activation means comprise pulsed electric current generators, connected to the magnetic elements (1).

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 in addition 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. System according to any of the preceding claims, further comprising one or more means of radio-frequency excitation of the sample under study.

9.- System according to any of the previous claims, where:

- the activation means of the magnetic elements comprise a control computer (7) integrated with a low power electronics module (8), connected to the magnetic elements (1) through a high power electronics module (9 ); and

- where, optionally, said high power electronics module (9) comprises a power supply and / or an analog power amplifier, configured to communicate with the control computer (7) in cooperation with the low power electronics module (8).

10. System according to any of the preceding claims, wherein said system is part of an MR equipment and / or a dental medical imaging equipment.

11.- Method of generating magnetic resonance signals on a sample under study, comprising the use of a system according to any of the preceding claims to generate a sequential activation of the magnetic fields (4) of the first set of n magnetic elements ( 1), generating an effective magnetic field (4) rotating around the axis (5).

12.- Method according to the preceding claim, where a single magnetic element (2) is activated at the same time among the n that make up the first set, for a period of time of duration t = T _r / n = 2 p / ( hw _G ), where T _r is the total period of revolution of the global magnetic field (4) of the system, and w _G is the corresponding angular frequency.

13. Method according to claim 11, where a plurality of magnetic elements (1) are activated at the same time, by means of alternating current generators.

14.- Method according to any of claims 11-13, which also comprises detecting 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. 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 F (r, t) that calculates the phase of acquired magnetic resonance for each point in space r, for each instant in time t

with:

where wi_ is the Larmor precession frequency, and the gyro-magnetic ratio, B (r, t ') the magnetic field at a given position r and an instant /', and m (r) is a unit vector in the direction of magnetization at r.