CN114209300A

CN114209300A - Pulse magnetic particle imaging method and system

Info

Publication number: CN114209300A
Application number: CN202111528372.8A
Authority: CN
Inventors: 贾广; 黄力宇; 田捷; 惠辉; 苗启广; 李檀平; 席力; 王颖; 梁小凤; 胡凯
Original assignee: Xidian University
Current assignee: Xidian University
Priority date: 2021-12-14
Filing date: 2021-12-14
Publication date: 2022-03-22

Abstract

The invention discloses a method and a system for imaging pulsed magnetic particles, wherein the method comprises the following steps: generating a pulsed, uniformly alternating main magnetic field; selecting a pulsed non-uniform alternating gradient magnetic field which generates at least one direction of X, Y and Z according to imaging requirements, changing the size of the pulsed non-uniform alternating gradient magnetic field, enabling a total spatial gradient magnetic field to traverse at least one preset direction in a space where a target to be detected which is injected with magnetic nanoparticles is located, and enabling the size of the pulsed magnetic field to obtain preset value changes in each preset direction by changing the size of a main magnetic field; continuously acquiring a voltage signal generated by the magnetic nano particles; obtaining the time domain attenuation area of the voltage signal as an imaging parameter in each half pulse oscillation period; and carrying out reconstruction imaging on the concentration distribution of the magnetic nanoparticles based on the system matrix by using the obtained multiple imaging parameters. Compared with the traditional MPI, the invention can reduce the power consumption of equipment, improve the signal-to-noise ratio of an image, the spatial resolution and the scanning efficiency and enlarge the imaging visual field.

Description

Pulse magnetic particle imaging method and system

Technical Field

The invention belongs to the field of medical imaging, and particularly relates to a pulse magnetic particle imaging method and system.

Background

Currently, clinical medical imaging is largely divided into two categories: one is structural imaging and one is functional imaging. Structural imaging is mainly to show the structures of organs and tissues inside the human body, and imaging methods include B-mode ultrasound, magnetic resonance, CT (Computed Tomography), and the like. Functional imaging is to show the function of blood vessels, organs, tissues and cells, and imaging methods include DSA (Digital subtraction angiography), PET (positron Emission Computed Tomography), SPECT (Single-Photon Emission Computed Tomography), CTA (CT angiography), and other techniques. Functional imaging typically requires the injection of a tracer into the body. If the tracer itself is radioactive, imaging can be performed directly with detectors, such as PET and SPECT techniques. If the tracer itself does not have radioactivity, such as iodine-containing contrast agents, imaging by X-ray scanning equipment, such as CTA and DSA techniques, is necessary. However, the radioactive tracer and X-ray in the functional imaging technique can cause certain ionizing radiation harm to the patient and the operating doctor.

In 2001, a completely new tracer-based Imaging modality, Magnetic Particle Imaging (MPI), was proposed. In 2005, Gleich and Weizenecker developed the first MPI static scanner in philips laboratories, which imaged using the non-linear magnetic response of superparamagnetic nanoparticles. It uses clinically certified superparamagnetic iron oxide nanoparticles (SPIONs) as tracers. The magnetic nano-particle has a magnetic core size in the range of 10-60nm, and can generate a high-frequency harmonic signal along with the change of an excitation magnetic field. MPI imaging mainly utilizes a selection Field to generate a magnetic Field Free Region (FFR), utilizes a focusing Field to rapidly move the magnetic Field Free Region, utilizes an excitation Field (driving Field) to excite the magnetic orientation of magnetic nanoparticles in the magnetic Field Free Region to generate a high-frequency harmonic signal, utilizes a receiving coil to receive the high-frequency harmonic signal, and obtains a spatial distribution image of the interior of a concentration living body of the magnetic nanoparticles through image reconstruction. Because the magnetic nanoparticles used in MPI do not have radioactivity, the imaging process does not need to use X-rays, so that no ionizing radiation exists, and the safety of doctors and patients is higher.

MPI can be used as an auxiliary treatment of a blood vessel imaging technology, for example, in the diagnosis and treatment process of cardiovascular and cerebrovascular diseases, the operation of implanting a stent and the like needs to be referred to blood vessel imaging. Conventional vascular imaging, however, requires the injection of iodine or gadolinium contrast agents into the patient, which require metabolism through the kidneys, and can be a significant burden and hazard for patients with reduced renal function. The magnetic nanoparticles used for magnetic particle imaging are metabolized through the liver, so that the kidney is not burdened, and the magnetic particle imaging method is safer for patients. Further, MPI does not require digital subtraction processing in DSA, and has fewer motion artifacts.

MPI to obtain signals at specific points or lines, it is necessary to use gradient coils to generate a small free region of magnetic field, which can be either a point region (free point of magnetic field) or a line region (free line of magnetic field). MPI adopts a point-by-point scanning or line-by-line scanning mode, a magnetic field free area is continuously moved for imaging, signals acquired each time only come from the magnetic field free area at a specific position, and the signal intensity depends on the concentration of magnetic particles in the magnetic field free area.

Since MPI usually uses one or more pairs of anti-helmholtz coils to construct the selection field, and a free magnetic field region (point or line) is formed in the middle of the selection field, in order to improve the image resolution, the free magnetic field point needs to be small enough, and the free magnetic field line needs to be thin enough, so that a large power consumption device is needed to generate a large enough current, so as to generate a large gradient magnetic field to meet the above requirements, which may result in large power consumption of the device. The spatial resolution of MPI is determined by the strength of a gradient magnetic field, the larger the gradient magnetic field is, the smaller the range of a magnetic field free region is, the fewer magnetic nanoparticles are used for generating signals, the smaller the signal strength is, the lower the signal-to-noise ratio is, and the poorer the image quality is, and the image resolution can only reach 5mm under a view field of 20 cm in MPI. The smaller the extent of the free region of the magnetic field, the more acquisition points are required, which results in longer scanning time and lower time resolution. Meanwhile, the relaxation effect of the magnetic nanoparticles can cause the movement of the free region of the magnetic field to lag and delay, so that the image becomes blurred, the spatial resolution of the image can be further reduced, and the scanning speed is reduced. And the MPI imaging field of view size is determined by a composite magnetic field formed by superposition of the selection field and the excitation field. Currently, MPI is mainly applied to mouse imaging, the imaging field of view is 1-3 centimeters, and the required excitation field strength is 10-30 mT. The scanning field of view of the human body usually needs 20-50 cm, which requires high excitation field strength and is therefore difficult to realize.

Therefore, in summary, MPI is difficult to meet the requirement of clinical human body scanning imaging.

Disclosure of Invention

In order to solve the above problems in the prior art, embodiments of the present invention provide a pulsed magnetic particle imaging method and system. The technical problem to be solved by the invention is realized by the following technical scheme:

in a first aspect, an embodiment of the present invention provides a pulsed magnetic particle imaging method, including:

generating a pulsed, uniformly alternating main magnetic field;

according to the imaging requirement, selecting a gradient magnetic field which generates pulse non-uniform alternation in at least one of the X direction, the Y direction and the Z direction, changing the size of the selected gradient magnetic field, traversing at least one preset direction in a space where a target to be detected which is injected with magnetic nanoparticles is located by the superposed total spatial gradient magnetic field, and changing the size of the main magnetic field in each preset direction to ensure that the size of the pulse magnetic field obtained by superposing the total spatial gradient magnetic field and the main magnetic field obtains preset value changes;

continuously acquiring a voltage signal generated by exciting the magnetic nanoparticles by the pulsed magnetic field;

aiming at each half pulse oscillation period, obtaining the time domain attenuation area of the voltage signal as an imaging parameter of the half pulse oscillation period;

and reconstructing and imaging the concentration distribution of the magnetic nanoparticles in the target to be detected based on the system matrix by using the obtained multiple imaging parameters.

In a second aspect, an embodiment of the present invention provides a pulsed magnetic particle imaging system, including:

the pulse excitation magnetic field module comprises a main magnetic field coil pair and gradient coil pairs in the X direction, the Y direction and the Z direction; wherein the main magnetic field coil pair is used for providing a pulse alternating main magnetic field in a Z direction under a controlled state; each direction gradient coil pair is used for providing a gradient magnetic field which is pulsed in the direction and is non-uniformly alternated under a controlled state; the target to be measured, into which the magnetic nanoparticles are injected, is placed in the spatial central region of the pulse excitation magnetic field module, and the long axis of the target is parallel to the Z axis; two coils of each coil pair are respectively arranged in parallel and oppositely at intervals;

the control module is used for controlling the gradient coil pairs in all directions to selectively generate a pulse non-uniform alternating gradient magnetic field in at least one of an X direction, a Y direction and a Z direction according to imaging requirements, changing the size of the selected gradient magnetic field, enabling a total spatial gradient magnetic field generated by superposition to traverse at least one preset direction in a space where the target to be detected is located, and changing the size of the main magnetic field by controlling the main magnetic field coil pairs in each preset direction, so that the size of the pulse magnetic field obtained by superposing the total spatial gradient magnetic field and the main magnetic field obtains a preset value change;

the receiving coil pair is used for generating induced voltage under the excitation of the pulse magnetic field;

the signal processing module is used for carrying out signal processing on the voltage signal obtained from the receiving coil and obtaining the time domain attenuation area of the voltage signal as an imaging parameter of each half pulse oscillation period aiming at each half pulse oscillation period;

and the image reconstruction module is used for reconstructing and imaging the concentration distribution of the magnetic nano particles in the target to be detected based on the system matrix by using the obtained multiple imaging parameters.

In the scheme provided by the embodiment of the invention, the field intensity of the pulsed non-uniform alternating gradient magnetic field in at least one direction is selectively changed according to the imaging requirement, the total spatial gradient magnetic field traverses at least one preset direction in the space where the target to be detected injected with the magnetic nanoparticles is located, and in each preset direction, the size of the pulsed magnetic field obtained by superposing the total spatial gradient magnetic field and the main magnetic field is changed for a preset value time, so that the field intensity spatial distribution of the pulsed magnetic field is different under each change of the main magnetic field intensity. Therefore, it is possible to deenergize the magnetic nanoparticles by a pulsed excitation field of different magnetic field strength provided in each preset direction, thereby achieving one-dimensional spatial encoding of magnetic nanoparticle concentration distribution information of a predetermined number of layers in the preset direction. On the basis, through the continuous change of the direction of the total spatial gradient magnetic field in one or more planes, the two-dimensional or three-dimensional spatial coding of the magnetic nanoparticle concentration can be realized by utilizing the field intensity spatial distribution change of the pulsed magnetic field, and voltage signals generated by exciting the magnetic nanoparticles along a plurality of directions and a plurality of gradient sizes are obtained. By extracting the time domain attenuation area of the voltage signal from the obtained voltage signal, based on the system matrix, the distribution image of the magnetic nanoparticle concentration after different dimensionalities are reconstructed can be obtained.

Compared with the traditional MPI imaging method, the embodiment of the invention carries out non-uniform pulse excitation on the magnetic nano particles in the whole space, the contribution of the voltage signal comes from all the magnetic nano particles in the space where the target to be detected is located, and a magnetic field free area is not required to be generated by using a selection field, so that high-power-consumption selection field hardware equipment can be avoided. The embodiment of the invention does not use a free area of a focusing field moving magnetic field, thereby avoiding the defects of sparse sampling and low spatial resolution caused by artifacts caused by non-uniform moving speed of the focusing field, non-uniform spatial sampling caused by irregular moving track and the like. Compared with the traditional MPI imaging method in which the voltage signal of a single magnetic field free area is weak due to the excitation voltage signal of the magnetic field free area, the embodiment of the invention has the advantages that the signal intensity is greatly enhanced, the signal-to-noise ratio is high, the image quality can be improved, and the requirement of clinical diagnosis is met. Because the embodiment of the invention does not adopt a scanning mode of a free area of a moving magnetic field, but carries out non-uniform pulse magnetic field excitation and space coding on the whole space, and a scanning area is not determined by the selection field gradient and the driving field strength together any more, the scanning area and the scanning range are easily expanded, the imaging visual field is not limited to small animals any more, but the size of a human body can be matched, and the clinical application of the human body can be realized. In addition, the embodiment of the invention does not adopt a movement mode of a magnetic field free zone of a Lissajous curve, so that the scanning time of a large area is obviously shortened, and the clinical scanning efficiency can be improved.

Drawings

FIG. 1 is a schematic flow chart of a pulsed magnetic particle imaging method according to an embodiment of the present invention;

FIG. 2 is a field intensity diagram of a pulsed, uniformly alternating magnetic field provided by an embodiment of the present invention;

FIG. 3 is a field intensity diagram of a pulsed, non-uniformly alternating gradient magnetic field provided by an embodiment of the present invention;

FIG. 4 is a schematic view of a spherical coordinate system;

FIG. 5 is a schematic diagram of a pulsed alternating magnetic field strength distribution, a pulsed current signal, and a magnetic nanoparticle responsive voltage relaxation decay signal provided by an embodiment of the present invention;

FIG. 6 is a graph of AUC versus pulsed alternating field strength H provided by an embodiment of the present invention;

FIG. 7 is a graph illustrating the variation of the AUC with the strength of a pulsed magnetic field provided by an embodiment of the present invention;

FIG. 8 is a schematic diagram of the structure and spatial orientation of a gradient coil pair in the X direction according to an embodiment of the present invention;

FIG. 9(a) is a schematic diagram of the coil shape and current flow direction in a Y-direction gradient coil pair provided by an embodiment of the present invention;

FIG. 9(b) is a schematic diagram of the structure and spatial orientation of a Y-direction gradient coil pair provided by an embodiment of the present invention;

FIG. 10(a) is a schematic diagram of coil shape and current flow direction in a Z-direction gradient coil pair provided by an embodiment of the present invention;

FIG. 10(b) is a schematic diagram of the structure and spatial orientation of a Z-direction gradient coil pair provided by an embodiment of the present invention;

fig. 11(a) and fig. 11(b) are a schematic diagram of a combined structure and a spatial orientation of gradient coil pairs in three directions and a schematic diagram of a position relationship between an object to be measured and each gradient coil, respectively, according to an embodiment of the present invention;

fig. 12 is a schematic structural diagram of a shielding coil assembly according to an embodiment of the present invention;

FIG. 13(a) is an original image of a one-dimensional reconstruction simulation experiment according to an embodiment of the present invention;

FIG. 13(b) is a one-dimensional projection view of a one-dimensional reconstruction simulation experiment according to an embodiment of the present invention reconstructed using a method according to an embodiment of the present invention;

FIG. 14(a) is an original image of a two-dimensional reconstruction simulation experiment according to an embodiment of the present invention;

FIG. 14(b) is a two-dimensional projection diagram of a two-dimensional reconstruction simulation experiment according to an embodiment of the present invention reconstructed by a method according to an embodiment of the present invention;

fig. 15 is a schematic structural diagram of a pulsed magnetic particle imaging system according to an embodiment of the present invention.

Detailed Description

The technical solutions in the embodiments of the present invention will be clearly and completely described below with reference to the drawings in the embodiments of the present invention, and it is obvious that the described embodiments are only a part of the embodiments of the present invention, and not all of the embodiments. All other embodiments, which can be derived by a person skilled in the art from the embodiments given herein without making any creative effort, shall fall within the protection scope of the present invention.

In order to solve the problems of the existing MPI magnetic particle imaging technology and meet the requirements of human clinical application, the embodiment of the invention provides a pulse magnetic particle imaging method and a pulse magnetic particle imaging system. In the embodiment of the invention, the magnetic nanoparticles are injected into the target to be detected in advance, and the target to be detected is scanned and imaged, so that a distribution image of the magnetic nanoparticle concentration in the space corresponding to the target to be detected can be obtained. The target to be detected in the embodiment of the invention can be a human body, an animal body and the like which can be injected with the magnetic nano particles, and the magnetic nano particles can be excited by utilizing the spatial non-uniform pulse magnetic field to generate an object reflecting the concentration signal of the magnetic nano particles.

In a first aspect, an embodiment of the present invention provides a pulsed magnetic particle imaging method. As shown in fig. 1, the pulsed magnetic particle imaging method proposed by the embodiment of the present invention may include the following steps:

s1, generating a pulsed, uniformly alternating main magnetic field.

In the embodiment of the invention, the main magnetic field generated by the step is an alternating magnetic field generated by using pulse current, and the strength of the alternating magnetic field is constant and uniform in the magnetic field direction. In the embodiment of the invention, the direction of the pulse current is alternately changed positively and negatively according to a half oscillation period. The pulse current of the embodiment of the invention can comprise: rectangular pulses, peaked pulses, trapezoidal pulses, etc.

The following description will be given by taking the magnetic field direction of the main magnetic field as the Z direction as an example.

In an alternative embodiment, the means for generating a pulsed, uniformly alternating main magnetic field comprises:

and a main magnetic field coil is loaded with equidirectional pulse alternating current with constant current.

The two main magnetic field coils of the embodiment of the invention are axially overlapped and face to the Z direction, and the two main magnetic field coils have a certain distance. The main magnetic field coil may be implemented by any coil in the prior art, such as a normal conducting coil or a superconducting coil, and the shape of the coil may be rectangular, circular, and the like, which is not limited herein.

To facilitate understanding of the form of the pulsed uniform alternating magnetic field, please refer to fig. 2, and fig. 2 is a field strength diagram of the pulsed uniform alternating magnetic field according to the embodiment of the present invention. Each trapezoid in fig. 2 is a field strength curve for half a pulse oscillation period. The dotted line frame shows a pulse oscillation period, the magnetic field intensity directions of two half pulse oscillation periods of the pulse oscillation period are opposite, and the pulse oscillation period is respectively in a regular trapezoid distribution and an inverse trapezoid distribution, but the amplitude values of the field intensities are the same. The process of generating the trapezoidal uniform field intensity distribution is described by taking a pulse oscillation period as an example, specifically: for example, for the first half of the pulse oscillation period, the main magnetic field coil pair is loaded with trapezoidal pulse current in the same direction, so that the current rapidly rises to a preset current value, and after the current is stable for a period of time, the current rapidly falls to a 0 value. And in the second half pulse oscillation period, a pulse current which is opposite to the previous pulse current is loaded to the main magnetic field coil pair, so that the current is quickly increased to a preset current value, and after the current is stable for a period of time, the current is quickly decreased to a 0 value.

Through the mode, the two main magnetic field coils can generate a pulse uniform alternating magnetic field with a certain frequency as the main magnetic field in the central imaging area corresponding to the space where the target to be measured is located.

The superposition of the gradient magnetic fields in all directions can generate a spatial total gradient magnetic field, the superposition of the main magnetic field and the spatial total gradient magnetic field can obtain a pulse magnetic field, and the pulse magnetic field is the sum of all final magnetic fields of the target to be detected in the space. On one hand, the embodiment of the invention introduces the main magnetic field with certain field intensity amplitude to change the position of the field intensity of the magnetic field 0 in the middle of the gradient magnetic field, so as to avoid the situation that the position of the field intensity of the magnetic field 0 is always at a fixed position, which can cause that the magnetic nanoparticles at the position can not be excited, no voltage signal is generated, the magnetic nanoparticles in the reconstructed image are in a black area, and the concentration of the magnetic nanoparticles at the position can not be correctly detected. On the other hand, in the magnetic nanoparticle imaging process, the primary magnetic field which is initially uniform and alternating can change the spatial distribution of the pulse magnetic field strength by changing the size, so that the field strength spatial distribution of the pulse magnetic field is different after the primary magnetic field is changed, independent spatial encoding of the magnetic nanoparticle concentration distribution can be performed, and then reconstruction imaging of the magnetic nanoparticle concentration distribution is realized. This part will be described in detail later.

S2, according to imaging requirements, selecting and generating a pulse non-uniform alternating gradient magnetic field in at least one of the X direction, the Y direction and the Z direction, and changing the size of the selected gradient magnetic field to enable the superposed total spatial gradient magnetic field to traverse at least one preset direction in the space where the target to be detected with the injected magnetic nanoparticles is located, and changing the size of the main magnetic field in each preset direction to enable the size of the pulse magnetic field obtained by superposing the total spatial gradient magnetic field and the main magnetic field to obtain a preset value change.

In an alternative embodiment, the means for generating a pulsed, non-uniformly alternating gradient magnetic field in either direction comprises:

the gradient coil pair in the direction is loaded with reverse pulse alternating current with the same current magnitude.

In the embodiment of the invention, the X direction, the Y direction and the Z direction are respectively provided with a gradient coil pair, and the types of the coils adopted by the gradient coil pairs in the three directions can be completely consistent, such as Maxwell coils and the like; of course, the gradient coils in the three directions may also have differences in the types of coils used, and are not particularly limited herein. The two gradient coils of the gradient coil pair in each direction have a spacing, the direction is taken as an axial direction, and the axial directions of the two gradient coils are coincident.

Regarding to the application of reverse pulse alternating currents with the same magnitude to the gradient coil pair in one direction, taking a pulse oscillation period as an example for explanation, regarding the first half of the pulse oscillation period, trapezoidal pulse currents in opposite directions are applied to the gradient coil pair in the direction, so that the currents of the two gradient coils both rapidly rise to a preset current value in the respective directions, and after the currents are stable for a period of time, the currents are rapidly reduced to a value of 0. And in the second half pulse oscillation period, loading pulse current in the direction opposite to the respective direction of the two gradient coils in the previous direction, so that the current of the two gradient coils rapidly rises to a preset current value in the respective direction, and rapidly falls to a value of 0 after the current is stable for a period of time.

By applying a reverse pulse alternating current with the same magnitude to the two gradient coils of the gradient coil pair in any direction, a gradient magnetic field in that direction can be generated in the central imaging region. The gradient magnetic field in any direction is a pulse non-uniform alternating magnetic field with a certain frequency, the magnetic field intensity is in linear gradient distribution in the direction, but is in uniform distribution in the other two directions, and the gradient magnetic field in any direction is the same as the magnetic field direction of the main magnetic field. Referring to fig. 3, fig. 3 is a field intensity diagram of a pulsed non-uniform alternating gradient magnetic field according to an embodiment of the present invention. It can be seen that the field intensity amplitude of the gradient magnetic field of fig. 3 becomes non-uniform linear distribution, exhibiting trapezoidal distribution of different heights, compared to the uniformly distributed field intensity of fig. 2.

In an embodiment of the present invention, the manner of changing the magnitude of any selected gradient magnetic field includes:

the currents of the gradient coil pairs of this direction are simultaneously increased or simultaneously decreased.

In the embodiment of the invention, for the gradient magnetic field in any direction, the magnitude of the current of the two gradient coils in the direction is increased or decreased simultaneously, so that the magnitude of the field strength of the gradient magnetic field in the direction can be changed.

Specifically, for example, for the Z direction, two gradient coils of a gradient coil pair in the Z direction are loaded with opposite pulse alternating currents of the same value, and in the central imaging region, a pulsed non-uniform alternating gradient magnetic field of a certain frequency can be generated. The current values of the two gradient coils in the Z direction are increased synchronously by a preset step after each half of the cosine oscillation period, so that the current values of the two gradient coils change synchronously for many times, and the magnetic field size of the original gradient magnetic field along the Z direction is increased linearly.

Due to the principle of vector superposition, a spatial total gradient magnetic field with any direction and magnitude can be generated by changing the magnitude of the gradient magnetic fields in three directions. Therefore, by changing the currents carried by the gradient coil pairs in each direction, the direction of the total gradient magnetic field in space can be changed. The total gradient magnetic field in space and the main magnetic field jointly form a pulse magnetic field, so that the direction of the pulse magnetic field can be changed by changing the direction of the total gradient magnetic field in space.

During the imaging process, the gradient magnetic field and the main magnetic field in which directions are used are superposed and are specifically selected according to the imaging requirements.

In an optional implementation manner, according to an imaging requirement, selecting a gradient magnetic field that generates non-uniform pulse alternation in at least one of an X direction, a Y direction, and a Z direction, and changing a magnitude of the selected gradient magnetic field to superpose a generated total spatial gradient magnetic field, traversing at least one preset direction in a space where a target to be detected into which magnetic nanoparticles have been injected is located, and changing a magnitude of a main magnetic field in each preset direction to obtain a predetermined value change in a magnitude of a pulse magnetic field obtained by superposing the total spatial gradient magnetic field and the main magnetic field, includes:

in the corresponding relation between the imaging and the magnetic field, selecting a gradient coil pair in at least one direction matched with the imaging requirement, obtaining a voltage sequence of the selected gradient coil pair in each direction, and obtaining a voltage sequence of a main magnetic field coil pair in each preset direction.

According to the mode of generating a current sequence by driving a voltage sequence, carrying out current loading on a gradient coil pair in at least one direction to generate a gradient magnetic field in a selected direction and change the size of the gradient magnetic field, so that a total spatial gradient magnetic field generated by superposition traverses at least one preset direction in a space where a target to be detected is located; and in each preset direction, a current sequence is generated by utilizing the driving of a voltage sequence of the main magnetic field coil pair, so that the current of the main magnetic field coil pair is changed according to the sequence of the current sequence after each half pulse oscillation period, the size of the main magnetic field is changed, and the size of the pulse magnetic field is changed for a preset value.

The imaging requirements comprise a target imaging dimension, a target imaging direction when the target imaging dimension is one-dimensional, and a target imaging plane when the target imaging dimension is two-dimensional. Wherein the target imaging dimension is one-dimensional, two-dimensional or three-dimensional. And when the target imaging dimension is one-dimensional, the target imaging direction comprises an X direction, a Y direction, a Z direction or any other space direction. The target imaging plane when the target imaging dimension is two-dimensional is an XY plane, an XZ plane, a YZ plane, or any other plane.

The gradient magnetic fields selected to be generated and changed are different for different imaging requirements, and the currents supplied to the several gradient magnetic fields selected are also different. But the respective current sequences of the main magnetic field, which cause the magnitude of the pulsed magnetic field to vary by a predetermined number of times, are identical in different predetermined directions of the total gradient magnetic field in space.

The corresponding relation between the imaging and the magnetic field is determined in advance by carrying out experiments on the current change and the field intensity of the gradient coil pair and the main magnetic field coil pair according to the relation between the gradient magnetic fields in the three directions and the total spatial gradient magnetic field.

Specifically, the correspondence between the imaging and the magnetic field may include a gradient magnetic field in a specific direction corresponding to the imaging requirement, a voltage sequence of each gradient coil of the gradient magnetic field in the specific direction, a respective current sequence, and a direction sequence traversed by the total spatial gradient magnetic field, and may further include a magnitude change sequence of the gradient magnetic field in each specific direction, a magnitude change sequence of the total spatial gradient magnetic field, and a voltage sequence and a current sequence of a main magnetic field coil pair in each preset direction traversed by the total spatial gradient magnetic field.

Referring to the schematic diagram of the spherical coordinate system in fig. 4, the relationship between the gradient magnetic fields in three directions and the total gradient magnetic field in space includes:

wherein G is_xRepresents the magnitude of the gradient magnetic field in the X direction; g_yRepresents the magnitude of the gradient magnetic field in the Y direction; g_zRepresents the magnitude of the gradient magnetic field in the Z direction; g represents a spaceMagnitude of total gradient magnetic field, θ and

the direction of the total gradient magnetic field in space is determined by the two angles of the spherical coordinate system, and the direction of the total gradient magnetic field in space is changed when any angle is changed; arctan (·) represents an arctangent function; arccos (·) represents an inverse cosine function.

Alternatively, the above relationship may also be expressed as:

G_x、G_yand G_zThe projection components of G on the coordinate axes. Thus, it can be understood that by adjusting G_x、G_yAnd G_zCan be combined to obtain a desired total gradient magnetic field magnitude G in space, and a direction characterizing the total gradient magnetic field in space

As is well known, magnetic nanoparticles have superparamagnetism, and when an external magnetic field exists, the magnetic moments of the magnetic nanoparticles existing in liquid can be deviated to the direction of the external magnetic field, so that magnetic flux changes are generated in a receiving coil, and a voltage signal is generated. Because a certain time delay is needed for the magnetic moment to be biased towards the direction of the applied magnetic field, under the excitation of the pulse, the voltage signal can generate relaxation attenuation in a time domain, and therefore the voltage signal can also be called as a voltage relaxation attenuation signal. For easy understanding, please refer to fig. 5, fig. 5 is a schematic diagram of a pulsed alternating magnetic field strength distribution, a pulsed current signal and a voltage relaxation decay signal of magnetic nanoparticle response provided by an embodiment of the present invention. FIG. 5 is a view showing an example of a magnetic field strength of-2.9 mT, in which a trapezoidal broken line indicates a change in the field strength of a pulsed alternating magnetic field; the curved dashed line represents the voltage signal in the ideal case; the curved solid line represents the actual voltage relaxation decay signal. From the curve of the actual voltage relaxation decay signal it can be seen that the voltage signal is in a decaying state from the moment the pulsed alternating magnetic field reaches a steady state.

In the embodiment of the present invention, the time domain attenuation area of the voltage relaxation attenuation signal generated by the response of the magnetic nanoparticles under the excitation of the pulsed alternating magnetic field with different magnetic field strengths is studied, and the result is shown in fig. 6, where fig. 6 is a relationship curve between AUC and pulsed alternating magnetic field strength H provided in the embodiment of the present invention. The attenuation area of the voltage signal, i.e., the voltage relaxation attenuation signal, in the time domain, referred to as AUC for short, in the embodiment of the present invention, as can be seen from fig. 6, the AUC has a nonlinear relationship with the pulse alternating magnetic field strength H. Thus, it is feasible to use AUC as a characteristic parameter for magnetic nanoparticle concentration encoding and reconstruction imaging.

Therefore, according to the embodiment of the invention, the corresponding relation between the imaging and the magnetic field can be predetermined through experiments according to different imaging requirements, and the currents of the gradient coil pairs in the specific direction are sequentially changed under the regulation rule guided by the corresponding relation, so that the magnitude of the gradient magnetic field in the specific direction is changed, the direction of the total spatial gradient magnetic field can be changed along a certain track in the space, and the purpose of traversing a plurality of preset directions is realized; and in each preset direction, under the regulation rule guided by the corresponding relation, the size of the pulse magnetic field is changed for a preset value for several times by changing the current mode of the main magnetic field coil, so that the pulse magnetic field excites the magnetic nanoparticles in the target to be detected to generate voltage signals with different preset values for several times in the preset direction, and the one-dimensional space coding of the concentration of the magnetic nanoparticles can be realized by utilizing the time domain attenuation Area (AUC) of the voltage signals with different times. It can be understood that when the direction of the total gradient magnetic field in space changes in a plane or in a space formed by a plurality of planes, two-dimensional space coding and three-dimensional space coding of the magnetic nanoparticle concentration can be correspondingly realized based on one-dimensional space coding under different changes of the field intensity space distribution of the pulsed magnetic field, and a reconstructed image of the magnetic nanoparticle concentration distribution in the target to be detected can be obtained through corresponding decoding reconstruction processing. Details of this section will be described later.

The current provided by the embodiment of the invention to each coil can be realized by using a computer and other control devices, for example, a waveform generator can be used for increasing the mains supply voltage to alternating current with a certain numerical value, the boosted alternating current is converted into direct current through rectification, and a frequency converter is used for obtaining pulse alternating current with a certain frequency, for example, the frequency is 1.67-50 KHz; and pre-driving the scanning sequence by using a front-end controller, further performing power driving, and distributing current to each coil under the high-voltage control of variable-frequency output. In addition, the current applied to each coil can be fed back to the front of the pre-driving through a feedback loop, so that closed-loop control is formed.

For the gradient coil, the specific way of obtaining the current may be that the sequence generator sends the time sequence parameter and the amplitude parameter of each coil voltage to the front-end controller according to the sequence parameter set by the computer software. The front-end controller outputs the amplitude parameter to the gradient controller according to the time sequence, and the gradient controller respectively sends voltage signals to the gradient coils in three directions in the given time sequence according to the obtained parameter and amplifies the voltage signals. It will thus be appreciated that the individual voltages received by the gradient coils in each direction constitute a sequence of voltages, each of which will drive a corresponding current.

With regard to the main magnetic field coils, the specific way in which the currents are obtained resembles gradient coils and will not be described in detail here.

And S3, continuously acquiring a voltage signal generated by the excitation of the magnetic nanoparticles by the pulsed magnetic field.

The magnetic field induces a current, and the direction and magnitude of the magnetic field are related to the direction and magnitude of the induced current. The change in the induced magnetic field can be reflected by a change in the voltage in the coil. The embodiment of the invention can utilize the receiving coil pair to receive the change of the magnetic flux caused by the magnetization response of the magnetic nano particles under the excitation of the pulse magnetic field. That is, the step may include:

and continuously acquiring a voltage signal generated by the magnetic nano particles under the excitation of the pulsed magnetic field by using the receiving coil pair.

In the receiving coil pair of the embodiment of the invention, the axial direction of the two coils is Z direction and has a distance. The type of the receiving coil pair is not limited, and any one of the existing coils can be selected according to the requirement.

And S4, acquiring the time domain attenuation area of the voltage signal as the imaging parameter of the half pulse oscillation period for each half pulse oscillation period.

According to the embodiment of the invention, the obtained voltage signals of each time are firstly subjected to relevant signal processing. And may specifically include an analog signal processing stage and a digital signal processing stage.

In an alternative embodiment, the analog signal processing stage may include the following processes: carrying out low-noise amplification on the obtained voltage signal; the Analog signal is converted into a Digital signal, which can be realized by an Analog-to-Digital Converter (ADC) having a certain sampling frequency, and a number of points can be sampled in a half pulse oscillation period corresponding to the excitation frequency of the pulse current.

In an alternative embodiment, the digital signal processing stage may include the following processes: and (3) reducing the signal of the excitation magnetic field aiming at the voltage signal of a complete pulse oscillation period or a half pulse oscillation period, and only keeping the voltage signal generated by the magnetic nano particles. And then extracting the time domain attenuation area of the voltage signal of each half pulse oscillation period from the voltage signal after signal processing, namely AUC (AUC) is taken as the imaging parameter of the half pulse oscillation period.

As before, AUC can be taken as a characteristic parameter for magnetic nanoparticle concentration encoding and reconstruction imaging. Regarding the manner of obtaining the AUC, in an alternative embodiment, the obtaining the time domain attenuation area of the voltage signal for each half of the pulse oscillation period includes the following two steps:

and selecting partial voltage signal sampling points corresponding to the stabilized pulse magnetic field intensity in the half pulse oscillation period from all sampling points of the voltage signal in the half pulse oscillation period aiming at each half pulse oscillation period.

And performing time domain integration on the selected partial voltage signal sampling points to obtain the time domain attenuation area of the voltage signal in the half pulse oscillation period.

Referring to fig. 7, it is understood that fig. 7 is a schematic diagram of changes in the pulsed magnetic field strength and AUC provided by the embodiment of the present invention. Fig. 7 shows the difference of the excitation field amplitudes in the original image by the difference of colors, and fig. 7 is a graph showing the effect of performing the gray processing on the color original image. The first row in fig. 7 shows the variation of the amplitude of the pulsed magnetic field, in the embodiment of the present invention, the amplitude of the pulsed magnetic field, i.e., the excitation field amplitude, varies, for example, in the range of-10 mT to 10mT, and the oblique lines represent the spatial total gradient field. The four graphs in the second row show the graphical representations of the excitation field amplitude and voltage signals for specific excitation field amplitudes-10 mT, -5mT, 5mT and 10 mT. Each of these four figures represents one half pulse oscillation period. Taking one of the graphs as an example, the trapezoid represents a field intensity curve of the pulsed magnetic field in the half pulse oscillation period, and the straight line at the top end of the trapezoid represents the field intensity amplitude, i.e., the excitation field amplitude. The voltage signal obtained is shown as a dashed line. In the actual signal receiving process, a plurality of voltage signal acquisition points are obtained for the voltage signals shown by the dotted lines, but the actually required AUC is a part of voltage signal sampling points corresponding to the stationary pulse magnetic field strength in the half pulse oscillation period, that is, a corresponding voltage signal sampling point after the start of the straight line at the top end of the trapezoid, which is indicated by a black part in the figure, and the numerical value of the area of the black part is the AUC numerical value corresponding to the figure. In addition, with respect to the lower four graphs of FIG. 7, it can be seen that as the excitation field amplitude increases, the AUC value decreases, which can be understood in connection with the curves of FIG. 6 for H at [ -10mT, -5mT ] and [5mT,10mT ].

Regarding how to select a partial voltage signal sampling point meeting the above requirement, the embodiment of the present invention may perform an observation experiment and a signal acquisition experiment on voltage signals of the magnetic nanoparticles under various pulse current oscillations in advance, and determine in advance a duration of a corresponding half pulse oscillation period and a duration for stabilization after each pulse current is loaded to be stabilized, for example, 2 microseconds and the like, of each magnetic nanoparticle under various pulse current oscillations. And establishing a data corresponding relation by using the experimental data. Therefore, during actual imaging, the stable duration used by a half pulse oscillation period of a certain magnetic nanoparticle under a certain pulse current can be searched from the data corresponding relation, and therefore, partial voltage signal sampling points after the stable duration is finished can be selected from a plurality of voltage signal sampling points obtained in the half pulse oscillation period. Or, through experiments, the data correspondence relationship can be made to include a proportion value of some voltage signal sampling points in the voltage signal sampling points of some magnetic nanoparticles under some pulse current and after the pulse magnetic field intensity is stable, that is, a part of voltage signal sampling points after the stable use time period ends accounts for all voltage signal sampling points of the half pulse oscillation period, for example, 80% or 90%. Therefore, during actual imaging, the proportion value of half pulse oscillation period of a certain magnetic nanoparticle under a certain pulse current can be searched from the data corresponding relation, and therefore, partial voltage signal sampling points of the proportion value can be directly selected from a plurality of voltage signal sampling points obtained in half pulse oscillation period in a reverse order. Of course, the manner of selecting the partial voltage signal sampling points meeting the above requirements based on the preliminary experiments is not limited to the above description, and may be reasonably designed according to the needs.

The process of performing time domain integration on a part of voltage signal sampling points belongs to the prior art, and is not described herein. It will be appreciated that as the pulse current varies, an imaging parameter can be obtained for each half of the pulse oscillation period.

And S5, reconstructing and imaging the concentration distribution of the magnetic nanoparticles in the target to be detected based on the system matrix by using the obtained multiple imaging parameters.

The imaging principles of embodiments of the present invention are briefly introduced below:

according to the intensity of the pulse excitation magnetic field, the shape and the size of the relaxation effect are different, and the shape and the size of the voltage signal attenuation are different. The embodiment of the invention adopts an excitation magnetic field in a form of a nearly square wave pulse, and is expressed as follows:

wherein H (t) represents the magnetic field strength; a represents the magnitude of the magnetic field strength; f represents the excitation frequency; t tableShowing time; n represents a coding iteration number; n represents the number of codes of the sample in the imaging area;

a square wave is shown.

The embodiment of the invention does not use a selection field and a focusing field in the existing magnetic particle imaging technology MPI, but adopts the technical scheme of full-area non-uniform pulse excitation, so that each point in the whole space is a magnetic field free area and can be excited by a pulse alternating magnetic field, thereby contributing to a voltage signal and greatly enhancing the signal-to-noise ratio. The time domain attenuation Area (AUC) of the voltage signal extracted from the voltage signal every time is equal to the linear superposition of AUC of all points/pixels of the magnetic nano-particles in the whole space. The research of the embodiment of the invention finds that AUC and magnetic field intensity A are in a nonlinear relation and in a direct proportion relation with magnetic particle concentration c. Therefore, the spatial encoding and the cross-sectional imaging can be performed through the relationship, namely, the AUC is used as the imaging parameter of the embodiment of the invention and is used for image reconstruction of the magnetic particle concentration.

In order to meet the requirement of spatial encoding by the pulse excitation magnetic field based on AUC, the following settings are considered in the embodiment of the present invention: the magnetic field is excited by a pulse having a magnitude varying in a single direction of XYZ, and is uniform in the remaining two directions perpendicular to the single direction to facilitate layer selection. Therefore, the embodiment of the invention sets the pulse non-uniform alternating gradient magnetic field in the XYZ direction, and changes the strength of the gradient magnetic field in each direction by changing the loaded pulse current, so that the gradient magnetic fields are superposed to change the direction of the total gradient magnetic field in space. And in each direction of the spatial total gradient magnetic field, the intensity of the main magnetic field is linearly changed by changing the pulse current loaded by the main magnetic field, so that the intensity of the final pulse magnetic field is changed in each direction of the spatial total gradient magnetic field. As described above, the gradient magnetic field directions in all directions are the same in the embodiment of the present invention, and are all Z directions. According to the embodiment of the invention, linear changes of field intensity of the pulse excitation type gradient magnetic field and the main magnetic field are utilized, and the nonlinear relation between AUC and the strength of the pulse type magnetic field is matched, so that under the action of the final pulse magnetic field, the field intensity spatial distribution is not uniform, the magnetic field spatial distribution is different every time, and each encoding and signal acquisition are mutually independent, and thus, the unique solution of the magnetic particle concentration matrix can be obtained.

One-dimensional spatial encoding and decoding of time-domain attenuation Area (AUC) of voltage signal, discretized magnetic field strength A, magnetic particle concentration c, and S representing AUC_decayObey the following relationship:

wherein S is_decay(t) represents the time domain attenuation area of the voltage signal, i.e., AUC; s_decay(a (r, t)) represents AUC per unit concentration of magnetic particles at the excitation magnetic field strength a; s (r) represents the receive coil sensitivity.

In the embodiment of the invention, the time corresponds to the current change, and the formula is discretized:

wherein S is_decay(i) Represents AUC; n represents the encoding number of samples in the imaging area, and in the formula, Delta V represents the volume size of the voxel of the data sampling point; g (i)_n,r_n) The system matrix g is an element of the system matrix g, and is independent of the magnetic particle concentration. The system matrix is used for representing the spatial distribution of AUC of voltage signals generated by magnetic particles with unit concentration under the action of a pulse magnetic field, and correction is realized by using the sensitivity of the receiving coil obtained by actual measurement in the construction process. Based on the system matrix, the magnetic particle concentrations corresponding to different moments (corresponding to different currents) of the change of the pulse magnetic field in each direction can be reversely deduced, so that the imaging is realized by using an image reconstruction method.

The operation matrix form of AUC is simplified as:

gc＝S_decay

wherein c represents a magnetic particle concentration matrix after one-dimensional decoding; s_decayRepresents AUCA matrix; if the system matrix g under the unit concentration is known, the magnetic particle concentration of each coding point can be calculated to obtain c. Therefore, the embodiment of the present invention may obtain the system matrix g in unit concentration through experiments in advance, obtain a plurality of AUCs through signal reception, and use c ═ g^-1S_decayAnd calculating the concentration of the magnetic particles to realize one-dimensional reconstruction. In actual reconstruction, c can not be calculated by direct inversion generally, and a regularization least square singular value decomposition method and an iterative solution algorithm can be used for assisting matrix solution. The solving process using the system matrix will not be described in detail here.

Therefore, according to the AUC of the magnetic nanoparticle response excited by the pulse magnetic field, the AUC is proportional to the concentration of the magnetic particles and the nonlinear relation of the pulse excitation magnetic field strength, multi-directional excitation and space encoding are carried out by combining the changing gradient magnetic field in three directions of XYZ and the changing main magnetic field, a plurality of AUCs can be obtained, one-dimensional reconstruction is carried out on the concentration space distribution of the magnetic particles through a system matrix, and two-dimensional or three-dimensional reconstruction processing is carried out on the basis of one-dimensional reconstruction data in a plurality of directions by using a related image reconstruction method, so that a two-dimensional or three-dimensional concentration space distribution image of the magnetic nanoparticles in the target to be detected can be obtained.

Hereinafter, alternative modes of each coil in the embodiment of the present invention will be described.

An X-direction gradient coil pair comprising:

a pair of Golay-type lateral gradient coils symmetric along a YZ plane, wherein each Golay-type lateral gradient coil comprises two Golay coils extending along the Z-direction. Each Golay coil is distributed on the cylindrical surface in a 120-degree circular arc, the field angle of the near circular arc is 68.7 degrees, and the field angle of the far circular arc is 21.3 degrees. Referring to fig. 8, fig. 8 is a schematic structural and spatial orientation diagram of an X-direction gradient coil pair according to an embodiment of the present invention.

A Y-direction gradient coil pair comprising:

a pair of Golay-type transverse gradient coils symmetric along an XZ plane, wherein each Golay-type transverse gradient coil comprises two Golay coils extending along the Z direction. Each Golay coil is distributed on the cylindrical surface in a 120-degree circular arc, the field angle of the near circular arc is 68.7 degrees, and the field angle of the far circular arc is 21.3 degrees. Referring to fig. 9(a), fig. 9(a) is a schematic diagram illustrating a shape of a coil and a current flow direction in a Y-direction gradient coil pair according to an embodiment of the present invention; the shape of the coil and the current flow direction are similar in the X-direction gradient coil pair, and are not illustrated here. Wherein, theta₀Representing the opening angle of the near circular arc; theta_rRepresenting the opening angle of the distant arc; z is a radical of_rAnd z₀Indicating different positions on the Z-axis.

Please refer to fig. 9(b) for the structure and spatial orientation of the Y-direction gradient coil pair, and fig. 9(b) is a schematic diagram of the structure and spatial orientation of the Y-direction gradient coil pair according to an embodiment of the present invention.

A Z-direction gradient coil pair comprising:

and the pair of circular Maxwell coils are axially overlapped, axially face to the Z direction and have intervals. Referring to fig. 10(a), fig. 10(a) is a schematic diagram illustrating a coil shape and a current flow direction in a Z-direction gradient coil pair according to an embodiment of the present invention; where d represents the coil pitch and R represents the coil radius. For the structure and spatial orientation of the Z-direction gradient coil pair, please see fig. 10 (b).

The gradient coil pairs in each direction are distributed in a staggered manner to surround a cylindrical space, as shown in fig. 11(a), fig. 11(a) is a schematic structural and spatial orientation diagram of the gradient coil pairs in three directions provided by the embodiment of the present invention. To facilitate understanding of the position relationship between the target to be measured and each gradient coil, please refer to fig. 11(b), where fig. 11(b) is a schematic diagram of the position relationship between the pair of gradient coils in three directions and the target to be measured according to the embodiment of the present invention. Fig. 11(b) shows that the target to be measured is a human body, the plane on which the human body lies is an XZ plane, the face of the human body faces the positive Y direction, and the cylinder in which the human body is located is a cylindrical space surrounded by the gradient coil pairs in each direction in a staggered distribution manner.

A main magnetic field coil pair comprising:

and the pair of circular Maxwell coils are axially overlapped, axially face to the Z direction and have intervals. The coil shape is shown in fig. 10(a), but the two coils are loaded with an alternating current in the same direction.

A receive coil pair comprising:

and a pair of circular Homholtz coils which are axially overlapped, axially face in the Z direction and have a spacing. And the spacing of the receiving coil pairs is larger than that of the gradient magnetic field excitation coil pairs in the Z direction.

That is, the receive coil pair is parallel to the gradient coil pair in the Z direction. One coil in the receiving coil pair is positioned outside one gradient coil in the gradient coil pair in the Z direction, and the other coil in the receiving coil pair is positioned outside the other gradient coil in the gradient coil pair in the Z direction.

For simplicity, the main magnetic field coil pair and the receiving coil pair are not illustrated in the drawings, and it can be understood by those skilled in the art that the main magnetic field coil pair and the receiving coil pair are also in the cylindrical space of the embodiment of the present invention in actual imaging.

In an alternative embodiment, the object to be measured may be placed on the carrying device during scanning imaging.

For example, the carrying device can be in the form of a bed body, a support and the like, and plays a role in carrying and fixing the target to be measured. The plane of the carrier is parallel to the XZ plane and the long axis is parallel to the Z axis. When preparing to perform scanning imaging, the carrying device can be moved so that the target to be measured is entirely located in the central imaging area. So as to scan and image the magnetic particle concentration distribution of the whole area of the target to be measured.

In practice, it may only be necessary to scan and image the magnetic particle concentration distribution of a local region of the target to be measured, and at this time, the central imaging region does not necessarily cover the entire region of the target to be measured. Therefore, in this case, in an alternative embodiment, before scanning imaging, the scanning position of the target to be measured can be determined by using the laser in the horizontal and vertical directions, and the position of the carrying device can be adjusted to align the scanning position with the central imaging area. For example, the object to be measured is a human body, the scanning part is a head, the patient who has injected magnetic nanoparticles can lie on the back of the body on the carrying device, and the bed body is pushed to push the head of the patient to a central imaging area in the cylindrical space through laser positioning.

In this case, since the scanning portion is only a local region of the object to be measured, magnetic nanoparticles are present in the remaining region other than the scanning portion even after the object to be measured is injected with the magnetic nanoparticles. Therefore, in order to accurately image the magnetic particle concentration only at the scanning portion of the target object, it is necessary to eliminate as much as possible the interference caused by the response voltage signal generated by the magnetic nanoparticles in the remaining portion of the target object except the scanning portion.

Therefore, in an optional embodiment, the magnitude of the selected gradient magnetic field is changed to superpose the generated total gradient magnetic field in space, and before the magnetic nanoparticles traverse at least one preset direction in space, the magnetic particle imaging method further includes:

and (3) restraining the magnetic nano particles except the target scanning area in the target to be detected by using external magnetic field saturation.

In this embodiment, the target scanning area corresponds to the central imaging area and also to the local scanning portion of the target. By applying a certain external magnetic field to the magnetic nanoparticles except the target scanning area, the part of the magnetic nanoparticles can be restrained and saturated, so that the part of the magnetic nanoparticles can not generate voltage signals, and the obtained voltage signals are ensured to be only from the magnetic nanoparticles in the target scanning area.

In an optional embodiment, the method for restraining magnetic nanoparticles in a target to be measured except a target scanning area by using saturation of an applied magnetic field includes:

and loading current to the shielding coils arranged under the other regions except the target scanning region in the target to be detected.

Specifically, a shielding coil assembly is arranged in the bearing device, and the shielding coil assembly comprises a plurality of coils which are arranged in parallel along the length direction of the bearing device; the coil of the shield coil assembly opposite to the central imaging region is the central imaging region coil, and the rest are the peripheral region coils. The central imaging region coil covers the projection range of the central imaging region in the XZ plane. In the imaging process, the peripheral area coil is loaded with current, the central imaging area coil is not loaded with current, namely, only the peripheral area coil is in an open state, so that a static magnetic field is generated to saturate and constrain the magnetic nano particles in the peripheral area, and the magnetic nano particles in the target to be detected, which are only positioned in the central imaging area, are excited by a pulse magnetic field, so that interference signals are avoided.

The type of the coil included in the shielding coil assembly is not limited herein, and in an alternative implementation manner, the shielding coil assembly may be implemented by using a rectangular coil, as shown in fig. 12, where fig. 12 is a schematic structural diagram of the shielding coil assembly according to the embodiment of the present invention.

For example, the shielding coil assembly may include 15 rectangular coils arranged along the length of the bed, each coil having a width of 10 cm and a length of 30 cm, and each coil has 200 turns and is loaded with a dc current of 30 amperes. During imaging, 2-5 central imaging area coils are closed, so that the magnetic nano particles in the central imaging area can be excited and oscillated by a pulse magnetic field to generate a voltage signal. And the coils of the other peripheral regions are opened to generate a static magnetic field of 30mT for saturation constraint of the magnetic nano particles in the peripheral regions to avoid generating interference signals.

In an optional embodiment, after performing reconstruction imaging on the concentration distribution of magnetic nanoparticles in the target to be detected based on the system matrix, the pulsed magnetic particle imaging method further includes:

and displaying and outputting the imaging result.

The imaging result can be displayed by using an image display, and the image display displays a distribution image of the magnetic nanoparticle concentration in the target to be detected, so that doctors and other personnel can conveniently observe the distribution image.

Outputting the imaging result may be achieved using a laser hologram camera or the like. A laser holographic camera is a device for taking a hologram by using laser as coherent light, is used for image printing to form a film for diagnosis, and is connected with a computer through a DICM interface. The external memory is used for connecting the computer to realize data storage and copying.

Further, the imaging result may also be stored and transmitted, or the like.

The above functions can be realized by using a PACS-RIS system. The PACS refers to a Picture Archiving and Communication System (PACS), which is a comprehensive system that has been rapidly developed in recent years with the progress of digital imaging technology, computer technology, and network technology and aims to comprehensively solve the problems of acquisition, display, storage, transmission, and management of medical images. The RIS is a radiology information management system (RIS), which is a software system for optimizing the workflow management of the radiology department of a hospital, and a typical flow includes links such as registration appointment, diagnosis, image generation, film production, report, audit, film distribution and the like.

In the following, the imaging process of different dimensions of the embodiment of the present invention is explained.

One-dimensional imaging

When the target imaging dimension is one-dimensional, the obtained multiple imaging parameters are utilized to reconstruct and image the concentration distribution of the magnetic nanoparticles in the target to be detected based on the system matrix, and the method comprises the following steps:

and carrying out one-dimensional reconstruction by using a system matrix to obtain one-dimensional reconstruction data including the magnetic particle concentration information of the preset value layers in the target to be detected in the preset direction, and forming a one-dimensional distribution diagram of the magnetic particle concentration in the target imaging direction.

The system matrix of the embodiment of the invention can be obtained in advance through experiments and is expressed as follows:

the magnitude of the pulsed magnetic field in the same preset direction is changed for a preset number of times, and the obtained AUC is expressed as:

then use c ═ g^-1S_decayC can be calculated.

Wherein, N of the parameter subscript is a preset value; i.e. i₀,i₁,…,i_N-1Representing a coil current which enables the pulse magnetic field in the preset direction to change for N times, namely a current sequence of a main magnetic field coil pair in each preset direction; r is₀,r₁,…,r_N-1Representing N position points in the preset direction; s_decay(i₁) Indicates a coil current of i₁AUC collected; g (i)_N-1,r₀) Magnetic particles representing unit concentration at a current i_N-1Under the action of the pulse alternating magnetic field, in the preset direction₀AUC generated by each location point; the meaning of the remaining elements is analogized. c denotes one-dimensional reconstruction data, which contains elements each of which is the concentration of magnetic particles at each position point in the central imaging region.

The system matrix can be used for obtaining one-dimensional reconstruction data of magnetic particle concentration information of a preset value layer in a preset direction, and the one-dimensional reconstruction data is characterized to be in an image form, namely a one-dimensional distribution diagram of the magnetic particle concentration in the target imaging direction.

The predetermined number of times is determined according to imaging resolution requirements. The larger the predetermined value is, the higher the imaging resolution is. It is understood that the data dimension of the obtained AUC is a predetermined value.

The specific solving process using the system matrix will not be described in detail here.

(II) two-dimensional imaging

When the imaging dimension of the target is two-dimensional, the obtained multiple imaging parameters are utilized to carry out reconstruction imaging on the concentration distribution of the magnetic nanoparticles in the target to be detected based on a system matrix, and the reconstruction imaging comprises A1-A2:

a1, changing the pulse magnetic field in each preset direction for a preset value to obtain a plurality of imaging parameters, and performing one-dimensional reconstruction by using a system matrix to obtain one-dimensional reconstruction data including the magnetic particle concentration information of the preset value layers in the target to be measured in the preset direction.

The one-dimensional reconstruction data is obtained as described in the one-dimensional imaging section.

A2, performing two-dimensional filtering back projection on all one-dimensional reconstruction data obtained from multiple preset directions changing in a specific plane to obtain a two-dimensional projection graph representing the concentration distribution of magnetic nanoparticles in a target to be detected in a target imaging plane.

Wherein the specific plane is determined according to the target imaging plane. The target imaging plane may be an XY plane, an XZ plane, a YZ plane, and any other planes.

In particular, for two-dimensional imaging, θ and

one angle is fixed, the other angle is traversed, the direction of the total gradient magnetic field in the space is changed in a specific plane corresponding to the fixed angle, in the preset direction formed by each traversal angle, the preset value change of the pulse magnetic field in the preset direction is realized by changing the current of the main magnetic field, and an AUC value can be obtained by changing the pulse magnetic field in each time.

Sum of theta in two-dimensional imaging

The numerical range of the imaging device is determined according to the imaging plane, and the traversing stepping of the traversing angle and the preset value times of the change of the pulse magnetic field in a certain preset direction are determined according to the imaging resolution requirement. Step of traversing angleThe smaller the feed, the larger the predetermined value, the higher the imaging resolution. It can be understood that the data dimension of the obtained AUC is the number of changes of the traversal angle × the predetermined value.

The mathematical principle of the filtered back projection reconstruction method is radon transform, and the method is commonly used in CT imaging reconstruction. For the specific transformation, please refer to the related prior art, which is not described herein.

(III) three-dimensional imaging

When the imaging dimension of the target is three-dimensional, the obtained multiple imaging parameters are utilized to carry out reconstruction imaging on the concentration distribution of the magnetic nanoparticles in the target to be detected based on a system matrix, and the reconstruction imaging comprises the following steps of B1-B3:

and B1, performing linear change on the pulse magnetic field in each preset direction for a preset value time to obtain a plurality of imaging parameters, and performing one-dimensional reconstruction by using a system matrix to obtain one-dimensional reconstruction data including the magnetic particle concentration information of the preset value layers in the target to be measured in the preset direction.

And B2, performing two-dimensional filtering back projection on all the one-dimensional reconstruction data respectively obtained from a plurality of preset directions belonging to the same plane to obtain a two-dimensional projection diagram related to the plane.

Please refer to S1302 for understanding, which is not described herein.

And B3, performing three-dimensional reconstruction on the obtained two-dimensional projection graphs respectively related to each plane to obtain a three-dimensional reconstruction graph representing the concentration distribution of the magnetic nanoparticles in the target to be measured.

The three-dimensional reconstruction is to calculate and obtain a distribution image of the concentration of the magnetic particles in the target to be measured in a three-dimensional space according to data information in the two-dimensional magnetic particle concentration distribution image projected along different directions. The adopted method can be chromatographic synthesis, filtering back projection reconstruction, iterative reconstruction or artificial intelligence reconstruction and the like. The specific procedures of these methods are not described in detail herein.

Specifically, for three-dimensional imaging, θ and

one angle is fixed, namely, the angle is used as a fixed angle, the other angle is used as a traversal angle to traverse in a corresponding step within a certain range, and in a preset direction formed by each fixed angle and the traversal angle, the size of the pulse magnetic field in the preset direction is changed for a preset value time by using a mode of changing the current of the main magnetic field.

And after the traversal of the traversal angle is finished, changing the original fixed angle by one step, and traversing the original traversal angle again according to the mode until the traversal of the traversal angle is finished.

And repeatedly executing the process by changing the original fixed angle for many times until the original fixed angle reaches the traversal upper limit value of the original fixed angle.

In the above process, first, the sum of theta

The traversing process in each preset direction in the space can be realized by starting traversing at a fixed angle without limitation. Theta and

the numerical range of the imaging angle is determined according to the three-dimensional imaging requirement, and the traversing stepping and the preset value of the traversing angle are determined according to the imaging resolution requirement.

It is understood that what is obtained

The imaging process of the embodiments of the present invention is illustrated below with reference to specific parameter values. It should be noted that the parameter values mentioned below are not intended to limit the embodiments of the present invention, but are merely an example of an implementation manner to facilitate understanding of the scheme, and in practical use, suitable values may be specifically selected according to needs.

Relevant parameters of magnetic field

The diameter of two circular Maxwell coils of the main magnetic field is 40 cm, the thickness and the width are both 5 cm, the number of turns of the coils is 200 turns, and the distance between the two coils is 40 cm. The two coils are loaded with a co-current in the maximum current range of 20-60 amperes, and a pulsed uniform alternating magnetic field with a maximum value of 10-20mT is generated in the central imaging region, with an excitation frequency of 1.67-5.0 kHz. In each preset direction of the spatial total gradient magnetic field, the current loaded on the main magnetic field coil is changed 256 times, and the size of the pulse uniform alternating magnetic field is changed 256 times, the strength is changed from-5.0 mT to 5.0mT, and each time, the change is 0.039 mT.

A pair of Golay-type transverse gradient coils of the X-direction gradient coil pair are applied with reverse pulse alternating currents at an excitation frequency of 1.67-50kHz to generate a pulsed non-uniform alternating gradient magnetic field of 50mT/m in the central imaging region. The axial magnetic field components are distributed in a linear gradient mode along the x direction and are uniformly distributed on a yz plane, the variation range of the magnetic field intensity within the range of 20 cm of the central imaging area is less than 5%, and the constant magnetic field surface is ensured to be a plane instead of a curved surface. During imaging scanning, the magnitude of the gradient magnetic field is changed by simultaneously increasing the current of the two gradient coils, so that the magnetic field strength of the gradient in the direction is changed from-50 mT/m to 50mT/m, and each time, the magnetic field strength is changed to 0.39 mT/m.

A pair of Golay-type transverse gradient coils of a Y-direction gradient coil pair are applied with reverse alternating currents at an excitation frequency of 1.67-50 kHz. A pulsed non-uniform alternating gradient magnetic field of 50mT/m is generated in the central imaging region. The axial magnetic field components are distributed in a linear gradient mode along the y direction, are uniformly distributed on an xz plane, and the variation range of the magnetic field intensity within the range of 20 cm of the central imaging area is smaller than 5%, so that the isomagnetic field surface is a plane instead of a curved surface. During imaging scanning, the magnitude of the gradient magnetic field is changed by simultaneously increasing the current of the two gradient coils, so that the magnetic field strength of the gradient in the direction is changed from-50 mT/m to 50mT/m, and each time, the magnetic field strength is changed to 0.39 mT/m.

In two circular maxwell coils of the gradient coil pair in the Z direction, the diameter of each coil is 40 cm, the thickness and the width are both 5 cm, the number of turns of the coils is 200 turns, and the distance between the two coils is 40 cm. The two gradient coils are supplied with pulsed alternating currents of opposite directions, the maximum current having a value in the range of 20 to 60 amperes and an excitation frequency of 1.67 to 50 kilohertz. A pulsed non-uniform alternating gradient magnetic field of 50mT/m is generated in the central imaging region. The axial magnetic field components are distributed in a linear gradient mode along the z direction and are uniformly distributed on the xy plane, the variation range of the magnetic field intensity in the central imaging area within the range of 20 cm is smaller than 5%, and the isomagnetic field surface is ensured to be a plane instead of a curved surface. During imaging scanning, the magnitude of the gradient magnetic field is changed by simultaneously increasing the current of the two gradient coils, so that the magnetic field strength of the gradient in the direction is changed from-50 mT/m to 50mT/m, and each time, the magnetic field strength is changed to 0.39 mT/m.

And two circular Homholtz coils of the receiving coil pair are used for receiving the magnetization vector change in the z direction. Each coil had a diameter of 40 cm, a thickness and width of 5 cm, and a spacing of 50 cm between the two coils.

One-dimensional spatial encoding and reconstruction

Specifically, the excitation frequency is 3.3KHz, and the signal sampling frequency is 16.5 MHz. The one-dimensional spatial encoding and reconstruction process is illustrated with the x-direction as an example. After the scanning part of the target to be detected is located in the central imaging area, determining the pulse alternating voltage value of each of the two gradient coils of the gradient coil pair in the X direction in the predetermined corresponding relation between the imaging and the magnetic field so as to drive and generate the corresponding pulse alternating current value. It should be added that in the one-dimensional spatial encoding and reconstruction, the voltage (current) of the gradient coil is fixed, and the foregoing voltage sequence (current sequence) of the main magnetic field coil pair in each preset direction can be understood as a plurality of identical voltage (current) values. And determining respective pulse alternating voltage sequences of two coils of the main magnetic field coil pair in X-direction imaging in the predetermined corresponding relation of imaging and magnetic field so as to drive and generate corresponding pulse alternating current sequences.

Gradient magnetic field G in X direction_xI.e. the total gradient magnetic field in space G. Pulsed alternating current value of the gradient magnetic field in the X direction, G_xIn the case where (θ is 0 °,

) For example, the magnetic field is-50 mT/m. The gradient magnetic field in the other two directions is always 0.

Where (θ is 0 °,

) In the preset direction, the current of the main magnetic field coil pair changes according to the pulse alternating current sequence, so that the size of the magnetic field changes once after each half pulse oscillation period of the main magnetic field, and the half pulse oscillation period of 256 times is completed in total, so that the size of the main magnetic field changes from-5 mT/m to 5mT/m by taking 0.039mT as stepping.

Thus, in the preset direction (θ is 0 °,

) Thus, 256 AUCs can be obtained, and the 256 AUCs are one-dimensionally reconstructed using the corresponding system matrix, so that the magnetic particle concentration of each of the 256 layers along the preset direction can be obtained, that is, the preset direction (θ is 0 °,

) One-dimensional reconstruction of the data.

A simulation experiment is carried out on the one-dimensional space encoding and reconstructing process, the obtained result is shown in fig. 13, and fig. 13(a) is an original image of the simulation experiment; only the white areas in the original image correspond to the magnetic nanoparticles. Fig. 13(b) is a one-dimensional projection view reconstructed by the simulation experiment using the method of the embodiment of the present invention.

The one-dimensional spatial encoding and reconstruction process with respect to the Y-direction and Z-direction is similar to that in the X-direction and will not be described repeatedly herein.

Two-dimensional space encoding and reconstruction

With respect to two-dimensional imaging, the imaging plane may be an XY plane, an XZ plane, a YZ plane, and an arbitrary plane. The XY plane is used as an example. Specifically, when the target imaging dimension is two-dimensional and the target imaging plane is an XY plane:

after the scanning part of the target to be detected is positioned in the central imaging area, determining respective pulse alternating voltage sequences of two gradient coils of the gradient coil pair in the X direction and the Y direction in a predetermined corresponding relation between imaging and a magnetic field so as to drive and generate corresponding pulse alternating current sequences; and determining respective pulse alternating voltage sequences of the main magnetic field coil pair and the two coils in each preset direction of the spatial total gradient field during XY plane imaging so as to drive and generate corresponding pulse alternating current sequences.

The gradient magnetic fields in the X direction and the Y direction are superposed to form a total spatial gradient magnetic field G. The first current value of the alternating current sequence of the gradient magnetic field in the X-direction and the Y-direction, G is (0 °,

) For example, a magnetic field of magnitude of-50 mT/m. The magnitude of the gradient magnetic field in the Z direction is always 0.

In the case where (θ is 0 °,

) In the initial preset direction, the current of the main magnetic field coil pair is changed according to the corresponding pulse alternating current sequence, so that the size of the magnetic field is changed once after each half pulse oscillation period of the main magnetic field, and the half pulse oscillation period of 256 times is completed, so that the size of the main magnetic field is changed from-5 mT/m to 5mT/m by taking 0.039mT as stepping, and the size of G is kept at-50 mT/m. Thus, in the preset direction (θ is 0 °,

) Then, 256 AUCs can be obtained, and the 256 AUCs are one-dimensionally reconstructed using the corresponding system matrix, so that a predetermined direction (θ is 0 °,

) The one-dimensional reconstructed data of (1) contains 256 values of the concentration of the magnetic particles.

Passing the respective current values after the first current value through the gradient magnetic field in the X direction and the Y direction according to the pulse alternating current sequence of the gradient magnetic field in the X direction and the Y directionCurrent change of field so as to maintain

Constant, theta increases along 1 deg., and in each preset direction (theta,

) And the current of the main magnetic field is changed for 256 times according to the corresponding pulse alternating current sequence to obtain 256 AUCs, and the corresponding system matrix is used for one-dimensional reconstruction to obtain one-dimensional reconstruction data in the preset direction. The above process is repeated until a predetermined direction (θ is 180 °,

) The data is reconstructed one-dimensionally.

And (4) performing two-dimensional filtering back projection on the 180 one-dimensional reconstruction data to obtain a two-dimensional projection diagram aiming at the XY plane.

It will be appreciated that 256 × 180 signal encodings due to half a pulse oscillation excitation are performed in total during two-dimensional imaging, i.e. the data dimension of AUC is 256 × 180. The signal sampling frequency is 16.5MHz, the excitation frequency is 3.3KHz, the number of sampling points in a half excitation oscillation period is 5000, 256 multiplied by 180 half oscillation periods are needed, and the time is 6.98 seconds.

With respect to XY plane imaging, spatial total gradient magnetic field direction variation can also be achieved by simultaneously adjusting the magnitude of the gradient magnetic fields in the three directions with corresponding pulsed alternating current sequences, but wherein the Z-direction gradient coil pair is loaded with a pulsed current sequence such that the magnitude of the Z-direction gradient magnetic field is always 0.

Similarly, with respect to XZ plane imaging, spatial total gradient field direction changes can be achieved by adjusting the currents of the gradient fields in the X and Z directions; regarding YZ plane imaging, the change of the total gradient magnetic field direction in space can be realized by adjusting the currents of the gradient magnetic fields in the Y direction and the Z direction, and the detailed process is not described in detail. However, in each preset direction of the spatial total gradient magnetic field, the field intensity of the pulse magnetic field is changed by changing the pulse alternating current of the main magnetic field 256 times.

A simulation experiment is carried out on the two-dimensional space encoding and reconstructing process, and the obtained result is shown in fig. 14, wherein fig. 14(a) is an original image of the simulation experiment, specifically, a maximum intensity projection image of a human head blood vessel image obtained by nuclear magnetic resonance; only the white areas in the original image correspond to the magnetic nanoparticles. Fig. 14(b) is a two-dimensional image reconstructed by the simulation experiment using the method of the embodiment of the present invention. Wherein the two-dimensional filtered back projection is obtained by inverse Radon transform. As can be seen from the simulation result, the two-dimensional image reconstructed by the method of the embodiment of the invention can clearly display the original magnetic particle distribution condition in the target to be measured.

Three-dimensional space coding and reconstruction

Specifically, after a scanning part of a target to be detected is located in a central imaging area, determining respective pulse alternating voltage sequences of two gradient coils of a gradient coil pair in each direction in a predetermined corresponding relation between imaging and a magnetic field so as to drive and generate corresponding pulse alternating current sequences; and determining a sequence of pulsed alternating voltages of the main magnetic field coil pair in each preset direction of the total gradient field in space to drive the generation of a corresponding sequence of pulsed alternating currents.

The gradient magnetic fields in the three directions are superposed to form a spatial total gradient magnetic field G. The first current value of the sequence of pulsed alternating currents of the gradient magnetic field in three directions, G is (0 °,

) For example, a magnetic field of magnitude of-50 mT/m.

In the case where (θ is 0 °,

) In the initial preset direction, the current of the main magnetic field coil pair is changed according to the corresponding pulse alternating current sequence, so that the size of the magnetic field is changed once after each half pulse oscillation period of the main magnetic field, and the half pulse oscillation period of 256 times is completed, so that the size of the main magnetic field is changed from-5 mT/m to 5mT/m by taking 0.039mT as stepping, and the size of G is kept at-50 mT/m. Therefore, the temperature of the molten metal is controlled,in a predetermined direction (θ is 0 °,

According to each current value of the pulse alternating current sequence of the gradient magnetic fields in the three directions after the first current value, the current change of the gradient magnetic fields in the three directions is kept

Constant, theta increases along 1 deg., and in each preset direction (theta,

) And changing the current of the main magnetic field coil pair 256 times according to the alternating current sequence corresponding to the main magnetic field to obtain 256 AUCs, and performing one-dimensional reconstruction by using the corresponding system matrix to obtain one-dimensional reconstruction data in each preset direction. The above process is repeated until a predetermined direction (θ is 180 °,

) The one-dimensional reconstruction of the data of the orientation. Using two-dimensional filtering back projection to 180 one-dimensional reconstruction data to obtain a target

A corresponding two-dimensional projection map.

Then, the currents through the three-directional gradient magnetic fields are varied in accordance with the current values after the pulse alternating current sequence of the three-directional gradient magnetic fields, so that

Is a step increment of 12 DEG, holding

Changing theta from 0 DEG to 180 DEG in the previous mode, changing the current of the main magnetic field coil pair for 256 times according to the pulse alternating current sequence of the main magnetic field coil pair under each theta, and finally obtaining the target

A corresponding two-dimensional projection map.

Changing again according to the subsequent current value of the pulse alternating current sequence of the gradient magnetic field in three directions

Repeating the above process until the target is obtained

A corresponding two-dimensional projection map. And carrying out chromatography synthesis on all the obtained two-dimensional projection images to obtain a distribution image of the magnetic nanoparticle concentration in the target to be detected in a three-dimensional space.

It can be understood that the signal encoding caused by 256 × 180 × 15 times of half-pulse oscillation excitation is completed in the three-dimensional imaging process, and the data dimension of AUC is 256 × 180 × 15. The number of internal sampling points is 5000, and the total time is 256 × 180 × 15 half oscillation periods, and the total time is 1.75 minutes, which is calculated by that the signal sampling frequency is 16.5mhz and the excitation frequency is 3.3 khz.

In the scheme provided by the embodiment of the invention, the field intensity of the pulsed non-uniform alternating gradient magnetic field in at least one direction is selectively changed according to the imaging requirement, the total spatial gradient magnetic field traverses at least one preset direction in the space where the target to be detected injected with the magnetic nanoparticles is located, and in each preset direction, the size of the pulsed magnetic field obtained by superposing the total spatial gradient magnetic field and the main magnetic field is changed for a preset value time, so that the field intensity spatial distribution of the pulsed magnetic field is different under each change of the main magnetic field intensity. Therefore, it is possible to supply a pulsed excitation field of different magnetic field strength in each preset direction to excite the magnetic nanoparticles, thereby realizing one-dimensional spatial encoding of magnetic nanoparticle concentration distribution information of a predetermined number of layers in the preset direction. On the basis, through the continuous change of the direction of the total spatial gradient magnetic field in one or more planes, the two-dimensional or three-dimensional spatial coding of the magnetic nanoparticle concentration can be realized by utilizing the field intensity spatial distribution change of the pulsed magnetic field, and voltage signals generated by exciting the magnetic nanoparticles along a plurality of directions and a plurality of gradient sizes are obtained. By extracting the time domain attenuation area of the voltage signal from the obtained voltage signal, based on the system matrix, the distribution image of the magnetic nanoparticle concentration after different dimensionalities are reconstructed can be obtained.

Meanwhile, the existing nuclear magnetic resonance imaging technology carries tissue information such as muscles and bones, and has a certain interference item for observing blood vessels. The embodiment of the invention utilizes the characteristic that the magnetic nanoparticles only exist in blood, digital subtraction is not needed in imaging, less motion artifacts exist, and the magnetic nanoparticle magnetic resonance imaging method can be used for targeted imaging. Compared with the existing PET and SPECT imaging technologies, the embodiment of the invention has higher sensitivity and image resolution, no ionizing radiation and easier production and storage of the tracer. The two-dimensional reconstruction method provided by the embodiment of the invention can replace the existing DSA angiography technology, and can provide quick and effective reference information for diagnosis and treatment of vascular diseases.

In a second aspect, an embodiment of the present invention provides a pulsed magnetic particle imaging system. As shown in fig. 15, the pulsed magnetic particle imaging system may include:

the pulse excitation magnetic field module comprises a main magnetic field coil pair and gradient coil pairs in the X direction, the Y direction and the Z direction; the main magnetic field coil pair is used for providing a pulse alternating main magnetic field in the Z direction under a controlled state; each direction gradient coil pair is used for providing a gradient magnetic field which is pulsed in the direction and is non-uniformly alternated under a controlled state; the target to be measured, into which the magnetic nanoparticles are injected, is placed in the spatial central region of the pulse excitation magnetic field module, and the long axis of the target is parallel to the Z axis; the two coils of each coil pair are respectively arranged in parallel and oppositely at intervals.

The control module is used for controlling the gradient coil pairs in all directions to selectively generate a pulse non-uniform alternating gradient magnetic field in at least one of the X direction, the Y direction and the Z direction according to imaging requirements, changing the size of the selected gradient magnetic field, enabling the total spatial gradient magnetic field generated by superposition to traverse at least one preset direction in the space where the target to be detected is located, and changing the size of the main magnetic field by controlling the main magnetic field coil pairs in each preset direction, so that the size of the pulse magnetic field obtained by superposing the total spatial gradient magnetic field and the main magnetic field obtains preset value changes.

And the receiving coil pair is used for generating induced voltage under the excitation of the pulse magnetic field.

And the signal processing module is used for carrying out signal processing on the voltage signal obtained from the receiving coil and obtaining the time domain attenuation area of the voltage signal as an imaging parameter of each half pulse oscillation period aiming at each half pulse oscillation period.

And the image reconstruction module is used for reconstructing and imaging the concentration distribution of the magnetic nanoparticles in the target to be detected based on the system matrix by using the obtained multiple imaging parameters.

The above-mentioned parts are respectively explained for facilitating the understanding of the embodiments of the present invention.

1) Pulse excitation magnetic field module

The magnetic field direction of the main magnetic field of the embodiment of the invention is the Z direction.

In an alternative embodiment, the means for providing a pulsed alternating main magnetic field in the controlled state by the pair of main magnetic field coils comprises:

loading equidirectional pulse alternating current to the main magnetic field coil pair by the control module; when the loaded equidirectional pulse alternating current is constant, the magnetic field intensity of the main magnetic field is in pulse uniform alternation; when the loaded equidirectional pulse alternating current changes, the magnetic field intensity of the main magnetic field is in pulse non-uniform alternation.

Specifically, the magnetic field strength of the main magnetic field generated by the pulsed magnetic particle imaging system is pulse-uniform and alternating, that is, the strength in the magnetic field direction is constant and uniform. However, in each preset direction of the total gradient magnetic field in the space, the magnetic field intensity of the main magnetic field is in pulse non-uniform alternation and has a certain linear change trend.

The two main magnetic field coils of the embodiment of the invention are axially overlapped and face to the Z direction, and the two main magnetic field coils have a certain distance. The main magnetic field coil can be implemented by any coil in the prior art, such as a normal conducting coil or a superconducting coil, and the shape of the coil can be rectangular, circular, and the like. And are not intended to be limiting herein.

According to the embodiment of the invention, the type of the main magnetic field coil can be reasonably selected according to the use requirement, and the specific structure is set, and the specific details are not described in detail herein. The main magnetic field coil is provided with a power interface, and the control module provides current for the main magnetic field coil and controls the main magnetic field coil through a power signal.

In an alternative embodiment, the pair of main magnetic field coils comprises:

and the pair of circular Maxwell coils are axially overlapped, axially face to the Z direction and have intervals.

Refer to fig. 10(a), but the two coils are loaded with the same-direction alternating current.

In an alternative embodiment, the gradient coil pair in each direction provides a pulsed non-uniform alternating gradient magnetic field in the direction under controlled conditions, comprising:

the control module loads reverse pulse alternating current with the same current to the gradient coil pair in the direction to generate a gradient magnetic field with pulse non-uniform alternation in the direction.

The gradient coil pair in the Z direction can be a Maxwell coil pair; the X-direction and Y-direction gradient coil pairs may be saddle coil pairs, such as Golay coil pairs in Golay. The gradient coil pair in each direction is provided with a positive electrode joint and a negative electrode joint which are connected with the control module, and the control module controls the current to be output to the control module so as to control the size of the gradient field in the direction.

In an alternative embodiment, the X-direction gradient coil pair includes:

a pair of Golay-type lateral gradient coils symmetric along a YZ plane, wherein each Golay-type lateral gradient coil comprises two Golay coils extending along the Z-direction. Each Golay coil is distributed on the cylindrical surface in a 120-degree circular arc, the field angle of the near circular arc is 68.7 degrees, and the field angle of the far circular arc is 21.3 degrees. Please refer to fig. 8.

In an alternative embodiment, the Y-direction gradient coil pair includes:

a pair of Golay-type transverse gradient coils symmetric along an XZ plane, wherein each Golay-type transverse gradient coil comprises two Golay coils extending along the Z direction. Each Golay coil is distributed on the cylindrical surface in a 120-degree circular arc, the field angle of the near circular arc is 68.7 degrees, and the field angle of the far circular arc is 21.3 degrees. Please refer to fig. 9.

In an alternative embodiment, the Z-direction gradient coil pair includes:

and the pair of circular Maxwell coils are axially overlapped, axially face to the Z direction and have intervals. Please refer to fig. 10.

The gradient coil pairs in each direction are distributed in a staggered manner to surround a cylindrical space, as shown in fig. 11 (a). The position relationship between the target to be measured and each gradient coil is shown in fig. 11 (b).

For the description of the main magnetic field coil pairs and the gradient coil pairs in each direction and the associated magnetic fields, reference is made to the relevant contents of the first aspect, and a repeated description is not made here.

2) Control module

The control module is mainly used for controlling the pulse current and can be operated on a computer. The system specifically comprises a waveform generator and a front-end controller corresponding to the waveform generator, wherein the waveform generator can boost the mains supply voltage to alternating current with a certain numerical value, convert the boosted alternating current into direct current through rectification, and obtain pulse alternating current with a certain frequency through a frequency converter, such as the frequency of 1.67-50 KHz; the front-end controller pre-drives the scanning sequence, further performs power driving, and distributes pulse current to each coil under the high-voltage control of variable-frequency output. In addition, the current applied to each coil can be fed back to the front of the pre-driving through a feedback loop, so that closed-loop control is formed.

Therefore, the control module can distribute current to the gradient coil pairs in any direction in the pulse excitation magnetic field module, provide reverse pulse alternating current with multiple changes, control the direction of the spatial total gradient field after the gradient magnetic fields in all directions are superposed to change, and provide the main magnetic field coil pairs with homodromous pulse alternating current with preset value changes in each preset direction of the spatial total gradient field so as to control the magnitude of the pulse magnetic field to change for preset value times in each preset direction, so that the field intensity spatial distribution of each pulse magnetic field generated in the central imaging area changes differently to oscillate the magnetic nanoparticles in the target to be detected.

In an optional implementation manner, the control module controls, according to an imaging requirement, the gradient coil pairs in each direction to selectively generate a pulsed non-uniform alternating gradient magnetic field in at least one of an X direction, a Y direction and a Z direction, and changes a magnitude of the selected gradient magnetic field, so that a total gradient magnetic field in a space generated by superposition traverses at least one preset direction in a space where a target to be measured is located, and changes a magnitude of a main magnetic field by controlling the main magnetic field coil pair in each preset direction, so that the magnitude of the pulsed magnetic field obtained by superposing the total gradient magnetic field in the space and the main magnetic field obtains a predetermined value change, including:

the control module selects gradient coil pairs in at least one direction matched with imaging requirements in the corresponding relation between the imaging and the magnetic field, obtains a voltage sequence of the selected gradient coil pairs in each direction, and obtains a voltage sequence of the main magnetic field coil pairs in each preset direction.

The imaging requirements comprise a target imaging dimension, a target imaging direction when the target imaging dimension is one-dimensional, and a target imaging plane when the target imaging dimension is two-dimensional; the corresponding relation between the imaging and the magnetic field is determined in advance by carrying out experiments on the current change and the field intensity of the gradient magnetic field coil pair and the main magnetic field coil pair according to the relation between the gradient magnetic fields in three directions and the total spatial gradient magnetic field.

The gradient magnetic fields in the selected specific directions are different, the varied reverse pulse alternating currents provided to the gradient magnetic field exciting coil pairs in the selected specific directions also have difference, but the alternating current sequence of the main magnetic field coil pairs is the same in each preset direction of the total spatial gradient magnetic field.

The relationship between the gradient magnetic fields in three directions and the total gradient magnetic field in space is described in the above related contents.

Of course, the control module may perform mechanical control of the system, remaining signal control, and the like, and the specific contents thereof will not be described in detail herein. For the specific content of the control module, please refer to the related content of the first aspect, and the description is not repeated here.

3) Receiving coil pair

It is known that a magnetic field induces a current, and the direction and magnitude of the magnetic field are related to the direction and magnitude of the induced current. The change in the induced magnetic field can be reflected by a change in the voltage in the coil. The receiving coil of the embodiment of the invention is used for receiving the change of magnetic flux caused by the magnetization response of the magnetic nano particles and generating corresponding voltage signals.

In an alternative embodiment, the receiving coil pair includes:

and a pair of circular Homholtz coils which are axially overlapped, axially face in the Z direction and have a spacing. And the spacing of the receiving coil pairs is larger than that of the gradient magnetic field excitation coil pairs in the Z direction. Please refer to the related contents of the first aspect.

4) Signal processing module

The signal processing module may be run on a computer. For processing the voltage signals detected from the pair of receiving coils.

In an optional embodiment, the signal processing module performs signal processing on the voltage signal obtained from the receiving coil pair, and obtains, for each half pulse oscillation period, a time-domain attenuation area of the voltage signal as an imaging parameter of the half pulse oscillation period, including:

the signal processing module performs low-noise amplification processing and analog-to-digital conversion processing on the voltage signals obtained from the receiving coil pairs.

And selecting partial voltage signal sampling points corresponding to the stabilized pulse magnetic field intensity in the half pulse oscillation period from all sampling points of the voltage signal in the half pulse oscillation period for each half pulse oscillation period of the processed voltage signal.

And performing time domain integration on the selected partial voltage signal sampling points to obtain the time domain attenuation area of the voltage signal as an imaging parameter of the half pulse oscillation period.

For the content of this section, please refer to the introduction of the first aspect specifically, and the description will not be repeated here.

5) Image reconstruction module

The image reconstruction module may be run on a computer. For the imaging process of different dimensions, specifically:

one-dimensional imaging:

when the target imaging dimension is one-dimensional, the image reconstruction module utilizes the obtained multiple imaging parameters to reconstruct and image the concentration distribution of the magnetic nanoparticles in the target to be detected based on the system matrix, and the method comprises the following steps:

the image reconstruction module performs one-dimensional reconstruction on a plurality of imaging parameters obtained by changing the size of the pulse magnetic field for a preset value time in a preset direction corresponding to the target imaging direction by using a system matrix to obtain one-dimensional reconstruction data including the magnetic particle concentration information of a preset value layer in the target to be detected in the preset direction, and forms a one-dimensional distribution diagram of the magnetic particle concentration in the target imaging direction.

Two-dimensional imaging:

when the target imaging dimension is two-dimensional, the image reconstruction module utilizes a plurality of obtained imaging parameters to reconstruct and image the concentration distribution of magnetic nanoparticles in the target to be detected based on the system matrix, and the method comprises the following steps:

the image reconstruction module performs one-dimensional reconstruction on a plurality of imaging parameters obtained by changing the size of the pulse magnetic field for a preset value for multiple times in each preset direction by using a system matrix to obtain one-dimensional reconstruction data including the concentration information of magnetic particles of the preset value layers in the target to be detected in the preset direction.

Performing two-dimensional filtering back projection on all one-dimensional reconstruction data obtained from a plurality of preset directions which change in a specific plane to obtain a two-dimensional projection graph which represents the concentration distribution of magnetic nanoparticles in a target to be detected in a target imaging plane; wherein the specific plane is determined according to the target imaging plane.

Three-dimensional imaging:

when the target imaging dimension is three-dimensional, the image reconstruction module utilizes the obtained multiple imaging parameters to reconstruct and image the concentration distribution of the magnetic nanoparticles in the target to be detected based on the system matrix, and the method comprises the following steps:

the image reconstruction module carries out linear change on the size of the pulse magnetic field in each preset direction for a preset value time to obtain a plurality of imaging parameters, and carries out one-dimensional reconstruction by utilizing a system matrix to obtain one-dimensional reconstruction data including the concentration information of magnetic particles in preset value layers in the target to be measured in the preset direction.

And performing two-dimensional filtering back projection on all the one-dimensional reconstruction data respectively obtained from a plurality of preset directions belonging to the same plane to obtain a two-dimensional projection graph related to the plane.

And performing three-dimensional reconstruction on the obtained two-dimensional projection drawings respectively related to each plane to obtain a three-dimensional reconstruction drawing representing the concentration distribution of the magnetic nanoparticles in the target to be measured.

With regard to the imaging process with different imaging requirements, please refer to the description of the first aspect, and the description will not be repeated here.

In addition, in an optional embodiment, the pulsed magnetic particle imaging system further includes:

and the bearing device is used for placing the target to be detected.

Wherein, a shielding coil component is arranged in the bearing device; the shielding coil assembly comprises a plurality of coils which are arranged in parallel along the length direction of the bearing device; a coil in the shielding coil assembly, which is opposite to a central imaging area of the pulsed magnetic particle imaging system, is a central imaging area coil, and the rest are peripheral area coils; wherein, the central imaging area is matched with the target scanning area of the target to be measured.

In an alternative embodiment, the pulsed magnetic particle imaging system further comprises:

and the bearing device control module comprises a laser positioning unit and a shielding coil assembly control unit.

The laser positioning unit is used for determining a target scanning area of a target to be detected by utilizing lasers in the horizontal and vertical directions, adjusting the position of the bearing device and aligning the target scanning area to the central imaging area.

And the shielding coil component control unit is used for loading current to the coils in the peripheral area in the imaging process so as to generate an external magnetic field to saturate and restrain magnetic nano particles except the target scanning area in the target to be detected.

Through the above manner, the local scanning part of the target to be detected, that is, the target scanning area, can be effectively imaged individually, and for this part, please refer to the related description of the first aspect, and a repeated description is not made herein. The outline structure of the pulsed magnetic particle imaging system according to the embodiment of the present invention is understood with reference to fig. 11, and it should be noted that, for simplicity, the main magnetic field coil pair and the receiving coil pair are not shown in fig. 11 (a).

In an alternative embodiment, the functions of the control module of the carrying device can also be realized by the control module.

a magnetic nanoparticle injector. The magnetic nanoparticle injector can be arranged outside the space region of the pulse excitation magnetic field module, namely outside the cylindrical structure, and is used for injecting the magnetic nanoparticles into the target to be detected.

In the pulsed magnetic particle imaging method and system of the embodiment of the invention, the magnetic nanoparticles can be superparamagnetic iron oxide nanoparticles (Resovist), and the magnetic nanoparticles are colloidal suspension with a concentration of 0.5mmol of Fe/mL and the like. The injection dose is set according to the weight of the object to be measured. The magnetic nanoparticles are injected intravenously, manually by a doctor, automatically by an instrument, and the like.

a closed housing; the closed shell can be used for placing the pulse excitation magnetic field module therein to form a hollow cylindrical accommodating space.

image display, laser camera and external memory.

The image display is used for displaying a distribution image of the magnetic nanoparticle concentration in the target to be detected, and is convenient for doctors and other personnel to observe. A laser holographic camera is a device for taking a hologram by using laser as coherent light, is used for image printing to form a film for diagnosis, and is connected with a computer through a DICM interface. The external memory is used for connecting the computer to realize data storage and copying.

Also, in an optional embodiment, the DICM interface is further connected to a PACS-RIS system.

With regard to the specific contents and related effects of the pulsed magnetic particle imaging system, please refer to the relevant parts of the pulsed magnetic particle imaging method described in the first aspect, which will not be described in detail herein.

The above description is only for the preferred embodiment of the present invention, and is not intended to limit the scope of the present invention. Any modification, equivalent replacement, or improvement made within the spirit and principle of the present invention shall fall within the protection scope of the present invention.

Claims

1. A pulsed magnetic particle imaging method, comprising:

generating a pulsed, uniformly alternating main magnetic field;

2. A method according to claim 1, characterized in that the means for generating a pulsed, uniformly alternating main magnetic field comprise:

3. The pulsed magnetic particle imaging method of claim 1, wherein the means for generating a pulsed non-uniformly alternating gradient magnetic field in either direction comprises:

4. The pulsed magnetic particle imaging method according to claim 1, wherein the selecting, according to the imaging requirement, a gradient magnetic field that generates pulsed non-uniform alternation in at least one of an X direction, a Y direction and a Z direction, and changing the magnitude of the selected gradient magnetic field to superpose a generated total gradient magnetic field in a space where a target to be measured into which magnetic nanoparticles are injected is located, traversing at least one preset direction, and changing the magnitude of the main magnetic field in each preset direction to superpose the total gradient magnetic field and the main magnetic field to obtain a predetermined value change in magnitude of the pulsed magnetic field, comprises:

in the corresponding relation between imaging and a magnetic field, selecting a gradient coil pair in at least one direction matched with imaging requirements, obtaining a voltage sequence of the selected gradient coil pair in each direction, and obtaining a voltage sequence of a main magnetic field coil pair in each preset direction;

according to a mode of generating a current sequence by driving a voltage sequence, carrying out current loading on the gradient coil pair in at least one direction to generate a gradient magnetic field in a selected direction and change the size of the gradient magnetic field, so that a total spatial gradient magnetic field generated by superposition traverses at least one preset direction in a space where the target to be detected is located; in each preset direction, a voltage sequence of the main magnetic field coil pair is used for driving to generate a current sequence, so that the current of the main magnetic field coil pair is changed according to the sequence of the current sequence after each half pulse oscillation period, the size of the main magnetic field is changed, and the size of the pulse magnetic field is changed for a preset value;

the imaging requirement comprises a target imaging dimension, a target imaging direction when the target imaging dimension is one-dimensional, and a target imaging plane when the target imaging dimension is two-dimensional; the imaging and magnetic field corresponding relation is determined in advance by carrying out experiments on current change and field intensity of the gradient coil pair and the main magnetic field coil pair according to the relation between the gradient magnetic fields in three directions and the total spatial gradient magnetic field.

5. The pulsed magnetic particle imaging method of claim 1 or 4, wherein the obtaining a time-domain attenuation area of the voltage signal for each half pulse oscillation period comprises:

for each half pulse oscillation period, selecting partial voltage signal sampling points corresponding to the stable pulse magnetic field intensity in the half pulse oscillation period from all sampling points of the voltage signals in the half pulse oscillation period;

6. The pulsed magnetic particle imaging method according to claim 5, wherein when the target imaging dimension is one-dimensional, the reconstructing and imaging the concentration distribution of the magnetic nanoparticles in the target to be measured based on the system matrix by using the obtained plurality of imaging parameters includes:

7. The pulsed magnetic particle imaging method according to claim 6, wherein when the target imaging dimension is two-dimensional, the reconstructing and imaging the concentration distribution of the magnetic nanoparticles in the target to be measured based on the system matrix by using the obtained plurality of imaging parameters includes:

carrying out one-dimensional reconstruction on a plurality of imaging parameters obtained by changing the size of the pulse magnetic field for a preset value for a plurality of times in each preset direction by using a system matrix to obtain one-dimensional reconstruction data including the concentration information of magnetic particles in a preset value layer in the target to be detected in the preset direction;

performing two-dimensional filtering back projection on all one-dimensional reconstruction data obtained from a plurality of preset directions which change in a specific plane to obtain a two-dimensional projection diagram which represents the concentration distribution of magnetic nanoparticles in the target to be detected in the target imaging plane; wherein the specific plane is determined from the target imaging plane.

8. The pulsed magnetic particle imaging method according to claim 7, wherein when the target imaging dimension is three-dimensional, the reconstructing and imaging the concentration distribution of the magnetic nanoparticles in the target to be measured based on the system matrix by using the obtained plurality of imaging parameters includes:

carrying out linear change on the magnitude of the pulse magnetic field in each preset direction for a preset value time to obtain a plurality of imaging parameters, and carrying out one-dimensional reconstruction by using a system matrix to obtain one-dimensional reconstruction data including the concentration information of magnetic particles in the preset value layers in the target to be detected in the preset direction;

performing two-dimensional filtering back projection on all one-dimensional reconstruction data which are respectively obtained from a plurality of preset directions belonging to the same plane to obtain a two-dimensional projection drawing related to the plane;

9. A pulsed magnetic particle imaging system, comprising:

10. The pulsed magnetic particle imaging system of claim 9, wherein the pair of main magnetic field coils are arranged in a manner to provide a pulsed alternating main magnetic field under controlled conditions, comprising:

loading, by the control module, a co-directional pulsed alternating current to the pair of main magnetic field coils; when the loaded equidirectional pulse alternating current is constant in magnitude, the magnetic field intensity of the main magnetic field is in pulse uniform alternation; when the loaded equidirectional pulse alternating current changes, the magnetic field intensity of the main magnetic field is in pulse non-uniform alternation.