Magnetic resonance imaging radio frequency magnetic field distribution measurement and human tissue data inversion method

Abstract

The invention relates to a magnetic resonance imaging radio frequency magnetic field distribution measurement and human tissue data inversion method, which comprises the following steps of: 1, respectively scanning a homogeneous water model and an imaging body by using spin echo sequences (90 degrees, TE and TR), (45 degrees, TE and TR) and (45 degrees, TE / 2 and TR) in sequence to obtain amplitude and argument data; and 2, according to the amplitude data of the sequences (90 degrees, TE and TR) and (45 degrees, TE and TR), calculating to obtain sensitive field distribution of the emission magnetic field and the receiving magnetic field of the imaging body without proton density coupling, and according to the argument data of the sequences (45 degrees, TE and TR) and (45 degrees, TE / 2 and TR), calculating to obtain phase distribution of the emission magnetic field and the receiving magnetic field of the imaging body. Based on magnetic resonance electrical characteristic imaging, the complex permittivity of the imaging body tissue is inversed through the emission magnetic field measurement quantity, and then the magnetic conductivity distribution of the tissue is obtained through further inversion of the receiving magnetic field and the existing complex permittivity.

Description

Technical Field

[0001] The invention belongs to the field of medical electromagnetic research, and in particular relates to a method for measuring radio frequency magnetic field distribution in magnetic resonance imaging and inverting human tissue data. Background Art

[0002] According to different magnetic properties, substances in nature are mainly divided into three categories: diamagnetic, paramagnetic and ferromagnetic. The main components of biological human tissue are water, proteins composed of carbon, hydrogen and oxygen, fat, etc., as well as trace metal elements. Macroscopically, human tissue is weakly diamagnetic, which is recorded as a real negative value magnetic susceptibility, which represents the degree of magnetization in an external magnetic field. When it comes to magnetic resonance imaging scanning, human tissue is in the main static magnetic field B0 and the magnetic field gradient B G The magnetized effect χ DC , and in the alternating exciting magnetic field The magnetized effect χ RF The former DC The impact on the proton density image is mainly manifested in the non-uniform distribution of the resonant Larmor frequency and the spatial fluctuation of the precession phase. RF It mainly affects the uniformity of the excitation magnetic field amplitude and the distribution of the angular distribution, and then affects the consistency of the distribution of the transverse vector formed by the deflection of the steady-state magnetization intensity vector in the fault. The superposition of the longitudinal and transverse relaxation effects presents different contrast weighting mechanisms on the image. The static magnetization effect can be well eliminated by radio frequency spin echo scanning technology, and the alternating (radio frequency) magnetization effect is one of the objects of the electromagnetic characteristic imaging magnetic field measurement and numerical inversion described in the present invention. According to Maxwell's equations, the alternating excitation and receiving magnetic fields generate conduction and displacement currents related to the induced electric field in the tissue. The relevant measurement characteristic parameters are recorded as conductivity σ and permittivity ε, which are the other two measurement inversion objects of the present invention.

[0003] Magnetic susceptibility χ m Both the magnetic permeability μ and the magnetic properties of a substance are used to characterize the magnetic properties of the substance. The following derivation relationship exists between the two. For a linear medium, the magnetization intensity is defined as M = χ m H, H is the strength of the external magnetic field. The magnetic field strength is expressed by the magnetic flux density formed in the medium, H = (B / μ0)-M, μ0 is the vacuum magnetic permeability. Therefore, the measured magnetic flux density is:

[0004] B=μ0(H+M)=μ0(1+χ m )H=μ0μ r H=μH (1)

[0005] In the formula, μ r is the relative magnetic permeability of the medium. The target parameter of the inversion of the present invention is μ r =1+χ m , χ mIt is also called volume susceptibility, which is different from mass susceptibility and molar susceptibility. For electromagnetic imaging based on nuclear spin radio frequency magnetic field resonance, eliminating B0 and B G Related μ DC After the contribution of , the measured excitation emission and detection reception magnetic field inversion characteristics are respectively distributed as follows:

[0006]

[0007] In the formula, and represent the amplitudes of the transmitting and receiving magnetic flux density vectors, respectively. is the equivalent external magnetic field strength, and '~' indicates that the field quantity is related to the effect when a unit current flows through the receiving coil. Because there is no external magnetic field during the signal receiving period, It is only proportional to the magnetization intensity M formed inside the medium, and the units of M and H are the same.

[0008] A special case is that in the field of medical electromagnetics research, magnetic susceptibility is usually used rather than magnetic permeability to characterize the magnetic properties of human tissue. The reason is that most types of biological tissue have a weak magnetic susceptibility of about -5.0×10 -6 level, so it is usually more convenient to study electromagnetic effects in ppm. For inorganic ferromagnetic materials, μ r =1+χ m ≈χ m , then magnetic permeability is normally used as the physical property analysis parameter.

[0009] The interaction between a medium and an electromagnetic field or space-time harmonics can be described by three parameters: magnetic permeability, electrical conductivity, and permittivity (μ, σ, ε). When analyzing the electrical properties of biological tissues with weak electrical conductivity, the effects of temperature, electromagnetic frequency response, and the physical state of the material are usually ignored, and the tissue is assumed to be an isotropic, inhomogeneous medium. During magnetic resonance scanning, the time-harmonic magnetic flux density will cause weak interfering induced conduction currents and displacement currents:

[0010]

[0011] In the rectangular coordinate system, is a vector composed of orthogonal partial derivatives in three-dimensional space, called nabla; J C is the conduction current, J D is the displacement current, E is the electric field strength, i is the imaginary unit, and ω is the angular frequency.

[0012] The ratio of the two current amplitudes is J C / J D =σ / ωε. In the RF band of magnetic resonance, the J of the human body C and J DThe amplitude levels are of the same order of magnitude, but with the increase of field strength J D Increase, for example, J of fat, gray matter, muscle, and cerebrospinal fluid at 7.0T (300MHz) C / J D They are approximately 0.4, 0.6, 0.7 and 1.7 respectively.

[0013] The existing magnetic resonance electrical imaging and the magnetic resonance electromagnetic imaging of the present invention rely on measuring the distribution of the radio frequency magnetic field. In a coordinate system rotating at the resonance Larmor frequency, the process of magnetic resonance imaging scanning human tissue to collect signals with radio frequency and gradient pulse sequences involves a right-handed excitation magnetic field (propagating from the transmitting coil to the imaging body) and a left-handed detection magnetic field (propagating from the imaging body to the receiving coil). The component H of the magnetic field intensity or the equivalent magnetic field intensity of the magnetization intensity in the Cartesian coordinate system is x and H y It is expressed as:

[0014]

[0015] In the formula, (B + ,B - ) and (H x ,H y ) are all complex numbers defined on the complex plane, and '*' indicates conjugation. + With B - or H + With H - They are of equal status, that is, in the process of electromagnetic property inversion, the inversion of different magnetic field distribution data sets measured by different excitation parameters for the same test image body should obtain the same electromagnetic property parameter value distribution.

[0016] Nuclear magnetic resonance uses static magnetic field, gradient magnetic field and radio frequency magnetic field as physical means to realize morphological anatomical imaging of human tissue, including voxel-based spatial position encoding and image reconstruction. In conventional magnetic resonance imaging, image reconstruction is completed based on the Fourier inverse transform from wave number domain to image domain. When the human body is placed in the static magnetic field B0, the hydrogen proton spin system in the body forms a macroscopic magnetization intensity vector Where 0 represents the initial state before the resonance excitation begins, M0 is the amplitude and is greater than 0, and M is negative along the z-axis. The negative sign means that the spin system is in a low-energy steady state without being excited, and indicates the diamagnetic nature of the hydrogen-containing proton material. For a nuclear spin system in thermal equilibrium within the sample, the amplitude of the macroscopic magnetization intensity is related to the spin density, static magnetic field strength, and system temperature T s For homogeneous samples, there are:

[0017]

[0018] Where Ns is the spin density per unit volume, γ is the gyromagnetic ratio, for hydrogen protons γ = 42.576MHz / T, is the reduced Planck constant, I is the spin quantum number, k is the Boltzmann constant, T s is the image body temperature. 1 H is the imaging nuclide, Some are constants. The number or density of hydrogen protons contained in organs and tissues in different parts of the human body is not the same. The external static magnetic field B0(r) = μ(r)H0(r) generated by magnetic resonance scanning superconductors or permanent magnets, where μ(r) is the tissue magnetic permeability distribution and H0 is the magnetic field strength. Not only does it have slight changes related to the spatial position r, but the actual distribution changes over time. These can be improved or even ignored by the uniform field and locking field system. Magnetic resonance imaging is also affected by certain changes in human body temperature, including physiological or pathological effects and the specific absorption rate of radio frequency during scanning. The temperature T s It is also a quantity that varies with space and time. In summary, under strict conditions, the magnetization intensity vector is

[0019] Radio frequency excitation transmits magnetic field for imaging Perpendicular to the B0 direction, '_' indicates a complex value, that is, it is periodic, and the frequency is the same as the precession frequency of M. in represents the amplitude of the circularly polarized magnetic field, Represents the initial phase. In the same direction, the flip angle is formed after interaction τ is the RF pulse width, and the generated transverse component can be expressed by the following formula:

[0020]

[0021] In the formula, i is an imaginary unit, which means that there is a phase increment on the complex plane That is, if If applied along the positive direction of the x-axis, for the right-handed precession M, its rotation and tilt will end in the positive direction of the y-axis.

[0022] The electromagnetic characteristic imaging is performed using the commonly used large-size birdcage-type whole-body excitation transmitting coil (whole-body) and the separate head and neck receiving coil. Based on the specific imaging scan sequence parameters, that is, setting the specific (α, TE, TR), where α is the excitation deflection angle, TE is the echo time, and TR is the repetition time, two-dimensional tomography is performed to measure the sensitive field amplitude and phase of the radio frequency magnetic field distribution. And set a certain frequency or phase encoding layer selection gradient, frequency encoding gradient and phase encoding gradient to determine the imaging field of view and voxel size, and obtain the lumped signal in the wave number domain but expressed as a complex value in the time domain:

[0023]

[0024] Where, voxel position vector r; M ⊥ and M 0,⊥ The initial phases are ω0t and Receiving magnetic field It represents the magnetic field generated when a unit current passes through the receiving coil, which is a circularly polarized left-handed magnetic field during the reception of magnetic resonance signals; is B0 and magnetic field gradient B G Acting together on the initial phase of the nuclear spin, if only B0 V is the volume of the imaging area.

[0025] For the image domain data reconstructed from the raw data of the scanning k-space linear array, which contains two parts of data, amplitude and phase, the equivalent spin density m(r) is defined by the following formula:

[0026]

[0027] In the formula, That is, the amplitude image displayed by common medical imaging films. The existing technology "RF field imaging" B1map is to measure the amplitude of the RF excitation transmission magnetic field, which is obtained through a special scanning sequence and data post-processing. However, it is impossible to measure the distribution of the transmission phase and the start of the reception phase because the end of the transmission phase and the start of the reception phase cannot be defined in the transmission and reception phase. We still need to "find another way out". Due to the special biological structure and function of the human eye, which is sensitive to amplitude ratio but not phase difference, the radiology department generally does not provide phase image films as a basis for differential diagnosis. In the imaging method based on gradient echo, B0 and B G The interaction with inhomogeneous tissues causes distortions in the phase image distribution related to magnetic susceptibility, so phase data is used in the reconstruction algorithms of susceptibility-weighted imaging (SWI) and quantitative susceptibility mapping (QSM). SWI and QSM can only rely on limited phase change information to invert the distribution of static magnetic susceptibility. Due to the limitations of the imaging principle itself and the difficulty in calculating the mapping from phase to dipole field source, the clinical application technology level of this type of imaging technology needs to be improved.

[0028] Since there is a coupling relationship between the magnetization intensity and the unit receiving magnetic field in the magnetic resonance detection signal, the existing technology mainly focuses on the measurement method of the transmission magnetic field distribution. According to the principle of nuclear magnetic resonance imaging, the ideal transmission magnetic field should have a uniform spatial distribution of amplitude, so as to excite the magnetization intensity vector to form a uniformly distributed deflection angle in the horizontal direction. In addition, because it can be generally considered that the magnetic permeability of the human body is approximately constant and equal to the magnetic permeability of the vacuum, the transmission magnetic field in formula (6) is different from the receiving magnetic field. The symbol has no wavy lines and the phase angle The actual non-uniformity depends only on the geometric structure of the transmitting coil.

[0029] A dual-angle method is also proposed in the prior art to measure the amplitude distribution of the transmitted magnetic field. The same imaging body is scanned twice with a gradient echo sequence with excitation deflection angles of α and 2α respectively, and the obtained image amplitude data are divided, and the result is:

[0030]

[0031] Because the flip angle 2α is set not to exceed 2π, the image data is divided and the inverse cosine is taken to finally obtain the excitation emission magnetic field Although according to the reciprocity theorem, the receiving magnetic field amplitude in the equivalent spin density m(r) expression of formula (8) The meaning is the magnetic field generated when a unit current passes through the receiving coil. However, because (TE, TR) controls the actual evolution of the transverse and longitudinal vectors in the continuous scanning sequence cycle, the above direct division of the magnetic vector amplitudes in pixels with different orientations without considering the (TE, TR) condition does not conform to the physical principle. In magnetic resonance imaging scanning, the images obtained by setting different echo times and repetition times show specific contrast between tissues with different longitudinal and transverse relaxation times, which is essentially a comparison of the longitudinal magnetic vector M. || Restoration of the steady state and initial transverse magnetic vector M 0,⊥ Therefore, it is wrong to forcibly use the amplitude image, that is, the reconstructed complex equivalent spin density amplitude, as the transverse magnetic vector amplitude to calculate the emission magnetic field amplitude distribution.

[0032] In order to increase the sensitivity of measuring the amplitude distribution of the transmitted magnetic field, Smith et al. proposed a dual-angle method in which the refocusing flip angle β is equal to 2 times the excitation deflection angle α. For a spin echo SE scanning sequence, under the conditions that the longitudinal relaxation time T1 of the imaging body is much greater than the echo time T1>>TE and the repetition time is much greater than T1 TR>>T1, the magnetic resonance signal amplitude SI is approximately equal to:

[0033]

[0034] Where x is the voxel position and C is a contrast variable related to tissue characteristics and sequence parameters. When β is equal to 2 times α, the signal amplitude is proportional to the cube of sinα. Furthermore, the quotient λ of the amplitude image data obtained from the two scans can be SE , the distribution of the transmitted magnetic field amplitude is obtained according to the following formula:

[0035]

[0036] Although theoretically this method has the best measurement sensitivity, the function sin 3 The slope is maximum when θ=60° and 120°. However, since the refocusing flip angle is not equal to 180°, the lateral component of the magnetic vector is not fully deflected, the magnetic vector is unevenly distributed on both sides of the average velocity position, and the refocusing plane has an inclination, that is, the lateral component on one side is sparse and the other side is dense, the echo signal is distorted. Combined with the effect of the spatial encoding gradient pulse group in each sequence cycle, the original magnetic resonance image is seriously interfered with by artifacts.

[0037] Direct measurement or indirect evaluation of the distribution of the phase component of the transmitted magnetic field is more difficult. The prior art believes that on conventional clinical magnetic resonance imaging scanning equipment, the transmit phase can only be measured by using a transmit-receive shared RF coil, such as a birdcage coil. Compared with the one-transmit-one-receive RF excitation-detection method, the single-channel transmit-receive shared coil exchanges the signal-ground polarity at the port when it acts as a receiving coil to detect the magnetic resonance induced signal. Therefore, when processing the phase of the equivalent spin density, it creates a favorable condition that the phase of the transmitting magnetic field and the receiving magnetic field can be considered to be approximately equal in value and opposite in sign. The following formula gives the equivalent spin density under the "transmit-receive phase assumption" condition,

[0038]

[0039] In the formula, Circularly polarized unidirectional emission magnetic field The initial phase Although the imaging excitation amplitude The distribution is approximately uniform, but because it contains the magnetic permeability parameter, its distribution is closely related to the actual characteristics of the tissue. Similarly, the phase component It must also be related to the voxel position.

[0040] The classic Bloch equations completely describe the longitudinal and transverse relaxation process of the spin system in the magnetic field. The macroscopic magnetization intensity vector in each voxel is regarded as a precessing arrow. After the RF excitation magnetic field pulse, its horizontal component moves in a circle at the Larmor frequency while its amplitude gradually decreases. From the perspective of the voxels, the horizontal component of the vector gradually dephases; while the longitudinal component has no phase angle factor, and its amplitude gradually recovers from zero to the original steady state. Combining the principles of pulse excitation and echo signal detection of magnetic resonance imaging scanning sequence, in order to calculate the transmitted and received magnetic field data from the 2D complex matrix of amplitude and phase angle reconstructed by the image, how to obtain the corresponding sensitive field (amplitude) and phase (unwrapped phase angle) distribution based on and using the basic processing methods in the laws of electromagnetic wave propagation and interaction, as well as combining the vector and its evolution characteristics, is the basis for electromagnetic imaging derived from the field of magnetic resonance imaging and even other related biomedical technologies and applications such as EEG, EEG, ECG and CMG.

[0041] The present invention focuses on implementing the measurement of the transmitting magnetic field and receiving magnetic field with separate transmission and reception based on the conventional coil facilities and conventional sequence technical conditions in clinical magnetic resonance, so as to be used for the inversion distribution reconstruction of the electrical and magnetic properties of human tissue imaging, thereby further assisting the differential diagnosis of diseases with the presentation of relevant information imaging, and discovering and evaluating problems that cannot be discovered by existing morphological anatomical structure imaging. In terms of electrical properties, human tissue exhibits both certain conductive properties and certain insulating dielectric properties; and in terms of magnetic properties, the human body as a whole can basically be regarded as having weak diamagnetism, that is, the magnetic permeability is slightly less than 1, and has a negative magnetic susceptibility. In order to describe the behavior of the radio frequency magnetic field in magnetic resonance imaging, under time-harmonic conditions, according to Faraday's law and Maxwell-Ampere's law in Maxwell's equations:

[0042]

[0043] Wherein, E is the magnetic field strength; ω is the propagation wave angular frequency, which is equal to the aforementioned Larmor resonance frequency ω0 in the present invention; J f The free current induced makes the human body conductivity σ, then J f =σE; D is the electric displacement vector. If the human body permittivity ε has D = εE. Since magnetic resonance imaging is only related to the excitation emission and detection reception magnetic field, and since the goal of electromagnetic property calculation and distribution reconstruction is to invert conductivity, permittivity and magnetic permeability, equation (13) can be written as:

[0044]

[0045] In the formula, is the complex permittivity The curl on both sides of the second equation in equation (14) is calculated and the first equation is substituted into the result, or the second equation is substituted into the first equation to eliminate the electric field, and when it is assumed that the human body magnetic permeability is equal to the vacuum magnetic permeability, a complex-valued equation containing one known quantity and one unknown quantity is obtained, with the magnetic field as the measured quantity and the complex permittivity as the target inversion quantity.

[0046] Magnetic Resonance Electrical Properties Tomography (MR-EPT) is an imaging method that uses magnetic resonance imaging to measure radio frequency magnetic fields to invert the distribution of electrical characteristic parameters of human tissue. In terms of imaging operation and data processing flow, it is mainly divided into implementing pulse sequence scanning to collect initial amplitude and phase angle data of magnetic resonance imaging, substituting the data into the calculation formula to obtain the magnetic field sensitivity field and phase data, and finally substituting the magnetic field data into a specific inversion algorithm to obtain the complex permittivity distribution result and image reconstruction.

[0047] Considering that the electromagnetic characteristics of real human tissue are very complex, with effects such as anisotropy and frequency dispersion, it is generally necessary to carry out characteristic inversion method research under the assumption of tissue homogeneity. Depending on whether the characteristic gradient assumption is ignored and the magnetic field sensitivity field and phase measurement method used, magnetic resonance electrical imaging has developed several different technical implementation routes. According to formula group (14), a mainstream Helmholtz equation method is proposed in the prior art, which ignores the gradient of complex permittivity, that is, there will be systematic errors in the inversion results at the interface of different tissues. Taking the transmitted magnetic field as an example, the core equation obtained is:

[0048]

[0049] The corresponding relationship between the real and imaginary parts of the complex solution of the equation is used to calculate the conductivity and permittivity of each voxel. However, the main problem with this method is that when the transmitted magnetic field is divided into amplitude and phase parts to calculate the permittivity and conductivity respectively, the expression for calculating the relative permittivity by amplitude is always negative, which does not conform to the actual physical situation. Therefore, the main research is to accurately invert the conductivity from the initial phase of the emission, but there is currently no reasonable measurement technology for the real phase distribution.

[0050] The prior art provides an electrical properties inversion method that takes into account the complex permittivity gradient factor, namely, gradient electrical properties tomography (gEPT). Since the complex permittivity, like the field itself, varies with space, the derivation of the core equations of this method, which are composed of scalars and vectors, increases the difficulty and complexity. Compared with the Helmholtz equation method, the applicable conditions of the gradient electrical properties tomography method are very different from those of clinical magnetic resonance imaging. It requires a special multi-channel independently controllable pulse transmission and reception common coil to complete the magnetic field measurement, and is divided into two steps: first, inverting the electrical properties gradient distribution from the magnetic field data, and then inverting the electrical properties from the gradient using the least squares method. Define the gradient term Still taking the calculation of electrical characteristics from the emitted magnetic field as an example, the relevant core equation is:

[0051]

[0052] In the formula, It indicates that the magnetic field sensitivity field transmitted by the i-th channel and received by the other channels is taken as the reference by the phase of the r-th channel; is the transmit reference phase of the rth channel; the right-hand component of the gradient vector g + Defined as g + =g x +ig y Although this method improves the distortion problem at the interface of the solution caused by the assumption of homogeneous electrical properties, it needs to overcome the ill-posedness of the solution in the calculation process, and the rationality of this method remains to be verified, especially the rationality of many unconventional excitation and detection methods used in magnetic resonance imaging.

[0053] Say goodbye to the problems of indirect solution by gradient and the lack of widespread use of multi-channel transceiver shared coils in clinical 3T and below. The existing technology proposes a convection-reaction equation method (convection-reaction Electrical Properties Tomography, crEPT). This method is similar in form to solving typical partial differential equations in fluid mechanics. More importantly, it derives the inversion core equation that the complex permittivity and magnetic field satisfy a linear relationship. Still starting from Maxwell's equations (14), the core equation of the convection-reaction method containing the electrical property gradient factor, which is homologous to gradient electrical property imaging, is derived:

[0054]

[0055] In the formula, u is the inverse of the complex permittivity. We further introduce the auxiliary vector The above core equation is derived and simplified into an equation about u:

[0056]

[0057] In the formula, α is an arbitrarily set diffusion coefficient, which is used to control the stability of the solution of the equation. According to the above formula, after presetting the initial values ​​of some electrical characteristics at the boundary of the solution domain, that is, the Dirichlet boundary condition, the complex permittivity distribution of the electrical characteristic imaging can be iteratively solved. In addition to solving the electrical characteristics by measuring the transmitted magnetic field, the corresponding electrical characteristics can also be solved by receiving the magnetic field. From the physical phenomenon point of view, It propagates inward from the boundary of the solution domain. It propagates from the solution domain to the boundary. Theoretically, although the amplitude and phase distribution of the transmitting magnetic field and the receiving magnetic field may be different, they both satisfy Maxwell's equations, and the electrical properties obtained should be similar. Although the above convection-reaction equation method does not ignore the electrical property gradient near the interface of different tissues, and only one known quantity can be measured to invert the only unknown quantity, like other magnetic resonance electrical property imaging, there is a homogeneous assumption that the magnetic permeability is equal to the vacuum magnetic permeability.

[0058] Although the magnetic properties of human tissues are not significant, some local magnetic properties are relatively significant, and can even be reflected and applied in auxiliary differential diagnosis of clinical imaging. The magnetic evolution process of hemoglobin molecules at different stages in the hematoma associated with cerebral hemorrhage has obvious contrast differences in magnetic resonance imaging images. The magnetic contrast agent injected for brain tumor patients undergoing enhanced MRI scans can regulate the contrast of the image to a certain extent to see the lesions clearly, and assist doctors in distinguishing between radiation necrosis, pseudoprogression, and early diagnosis of tumors. Therefore, various potential magnetic property imaging methods, including existing susceptibility-weighted imaging, quantitative susceptibility imaging, perfusion imaging, and blood oxygen level-dependent functional MRI, have important application value in clinical medicine.

[0059] Magnetic resonance electrical imaging and related magnetic imaging can measure the distribution of electromagnetic characteristic parameters of human tissues. Some physiological and pathological changes are sometimes not reflected by structural imaging too early. As a potential functional imaging, electromagnetic imaging is expected to play a role in the diagnosis of malignant diseases such as difficult-to-treat and difficult-to-detect tumors, cerebrovascular diseases, breast cancer, prostate cancer, etc. However, it is still difficult to effectively, accurately and reliably measure the distribution of RF transmitting magnetic fields and receiving magnetic fields in existing magnetic resonance electrical imaging. Most measurement methods are either based on single-channel birdcage-type transmitting-receiving common coils that are no longer used in clinical practice, or rely on ultra-high field transverse electromagnetic multi-channel independent transceiver controllable coils that are only available on scientific research magnetic resonance systems; and when inverting electrical characteristics, there are generally prerequisites such as homogeneous assumptions and magnetic permeability vacuum approximation, and it is often difficult to fully measure the amplitude and phase information of the magnetic field to seek to derive more complex calculation formulas and physical quantity processing methods. Summary of the invention

[0060] In order to solve the problem that the radio frequency detection coil receiving magnetic field distribution cannot be measured and the transmitting magnetic field sensitive field and phase distribution cannot be measured simultaneously in the existing magnetic resonance electrical characteristic imaging, the present invention provides a method for measuring the radio frequency magnetic field distribution of magnetic resonance imaging and inverting human tissue data, which uses a homogeneous water model to calibrate the sensitive field and phase distribution data of the measured magnetic field, and involves a method for independently measuring two magnetic field distributions, and the measurement results include two sets of data of the sensitive field amplitude component and phase component of the magnetic field. The present invention also relates to a method for inverting human tissue data, which is a method for inverting the distribution of electrical and magnetic characteristic parameters in human tissue based on magnetic resonance radio frequency magnetic field inversion imaging, and in particular, it relates to a method for first inverting the complex permittivity with a set of measured magnetic fields, and then further inverting the magnetic permeability with another set of magnetic fields combined with the complex permittivity.

[0061] In order to solve the problem that there is a lack of methods to decouple the receiving magnetic field amplitude from the proton density in the magnetic resonance signal in the existing magnetic resonance electrical imaging, and that the birdcage body coil excitation and multi-channel coil receiving configuration of clinical magnetic resonance cannot meet the needs of electrical imaging for multi-group discrete magnetic field measurements in a single scan, the present invention can measure the sensitive field and phase distribution of the receiving magnetic field and the transmitting magnetic field. Based on the Maxwell equations, the core inversion equation is derived by eliminating the electric field. From the two complex-valued known quantities of the receiving magnetic field and the transmitting magnetic field, the two unknown quantities of the electrical characteristic and the magnetic characteristic parameters are solved, where the electrical characteristic is the complex permittivity and the magnetic characteristic is the real-valued permeability.

[0062] It should be pointed out that, like the existing magnetic resonance electrical characteristic imaging method, the present invention does not consider the anisotropy of human tissue in the calculation process of inverting electromagnetic characteristic parameters, that is, for a given imaging target fault, it is assumed that the same electromagnetic characteristics will be obtained by inverting the measurement in directions other than the RF circular polarization direction of the excitation and detection magnetic fields. It is assumed that in the frequency band of the magnetic resonance RF magnetic field, the electromagnetic characteristics of human tissue have no frequency dispersion effect. It is assumed that human tissue is a magnetic linear medium, that is, when excited by a 90° deflection angle pulse sequence and scanned by a 45° pulse sequence, a magnetic flux density B proportional to the magnetic field intensity corresponding to the pulse is generated in the human tissue, and the relevant proportional coefficient is the ratio of the current twice passing through the transmitting coil to generate the magnetic field intensity H, which is equal to the magnetic permeability μ. In addition, the human body is weakly diamagnetic, and the typical hysteresis loop, magnetic memory and other effects of ferromagnetic materials are not considered. Unlike the permittivity, the permeability is considered to be a real number, that is, there is no phase difference between the alternating B and H, no time lag effect, and no magnetic loss. In addition, the changes in permeability and complex permittivity with temperature are also negligible.

[0063] The magnetic resonance imaging scanning process involves the main static magnetic field, the gradient magnetic field and the radio frequency magnetic field. The non-uniformity of the main static magnetic field can be compensated in the signal by the spin echo sequence. The eddy current caused by the rising and falling edges of the gradient magnetic field can be eliminated by changing the polarity of the phase image data by two scans and averaging. In the magnetic resonance imaging gradient echo technology, because the static magnetic susceptibility difference is retained in the phase relationship of the signal readout process, the aforementioned magnetic susceptibility weighting and quantitative magnetic sensitivity imaging are used to reproduce the low-frequency magnetic characteristics of the tissue, which is the target of the dipole field source model inversion in the latter. In the magnetic resonance electrical characteristic imaging and the magnetic resonance electromagnetic characteristic imaging of the present invention, a spin echo or fast spin echo sequence is used to collect the original magnetic resonance image data, and the static magnetic susceptibility difference, including the contribution of B0 non-uniformity and gradient magnetic field spatial encoding non-uniformity, is compensated by a high-amplitude radio frequency refocusing pulse. The transmission magnetic field and the receiving magnetic field distribution including the influence of radio frequency coil non-uniformity and the difference in electromagnetic characteristics of human tissue are measured and calculated, and the electromagnetic characteristic parameter distribution of the radio frequency band is inverted.

[0064] In order to achieve the above object, the present invention adopts the following technical scheme:

[0065] A method for measuring radio frequency magnetic field distribution in magnetic resonance imaging and inverting human tissue data, the method comprising the following steps:

[0066] Step (1) Scanning and imaging body radio frequency magnetic field measurement data:

[0067] Step (1.1) performs spin echo scanning on a homogeneous water phantom, and the sequence excitation deflection angle, echo time and repetition time parameters are (90°, TE, TR) respectively, to obtain amplitude image and angular image data;

[0068] Spin echo scanning was performed on the homogeneous water phantom, and the parameters of the sequence excitation deflection angle, echo time and repetition time were (45°, TE, TR) respectively, to obtain the amplitude image and the angle image data;

[0069] Spin echo scanning was performed on the homogeneous water phantom, and the parameters of the sequence excitation deflection angle, echo time and repetition time were (45°, TE / 2, TR) respectively, to obtain the amplitude image and the angle image data;

[0070] Spin echo scanning is performed on the imaging human body part, and the parameters of the sequence excitation deflection angle, echo time and repetition time are (90°, TE, TR) respectively, to obtain amplitude image and angular image data;

[0071] Spin echo scanning is performed on the imaging human body part, and the parameters of the sequence excitation deflection angle, echo time and repetition time are (45°, TE, TR) respectively, to obtain amplitude image and angular image data;

[0072] Spin echo scanning is performed on the imaging human body part, and the parameters of the sequence excitation deflection angle, echo time and repetition time are (45°, TE / 2, TR) respectively, to obtain amplitude image and angular image data.

[0073] In each of the above spin echo scans, the geometric parameters such as field of view, slice thickness, slice spacing, frequency encoding number, phase encoding step and number of slices were the same and kept consistent;

[0074] Step (1.2) processes the homogeneous water model and the scan data of the imaged human body part, and obtains the radio frequency receiving magnetic field sensitivity field distribution data and the transmitting magnetic field sensitivity field distribution data of the human body tissue;

[0075] Step (1.3) processes to obtain radio frequency receiving magnetic field phase distribution data and transmitting magnetic field phase distribution data of human tissue;

[0076] Step (2) Invert the conductivity and permittivity from the magnetic field measurement data:

[0077] Assuming that the magnetic permeability of human tissue is constant and equal to the magnetic permeability of vacuum, the distribution of conductivity and relative permittivity in the tissue is obtained by inverting the transmitted magnetic field or received magnetic field data induced by the human body using the magnetic resonance electrical characteristic imaging method;

[0078] Step (3) Invert the magnetic permeability or magnetic susceptibility from the magnetic field measurement data and the existing electrical properties:

[0079] The recovery method assumes that the magnetic permeability of human tissue varies with space. The human body receives or transmits the magnetic field and combines the electrical characteristics data that have been obtained to invert the distribution of magnetic permeability in the tissue based on the partial differential equation of the magnetic field.

[0080] Under the action of the static magnetic field, the nuclear spin system precesses at the Larmor frequency. The transverse radio frequency magnetic field creates conditions for the generation of magnetic resonance signals. The magnetic field gradient is applied to perform frequency and phase encoding on the precessing distributed magnetization intensity vector in the wavenumber domain. Finally, the amplitude and phase distribution of the magnetic vector in the spatial domain are obtained through the inverse Fourier transform of magnetic resonance imaging.

[0081] The present invention is based on the reconstructed image of the equivalent magnetic vector distribution signal acquired by the spin echo and fast spin echo scanning sequence, that is, the equivalent spin density distribution information, combined with the sequence parameter flip angle, echo time and repetition time control to realize the manipulation of the required magnetic vector amplitude ratio and phase difference, and obtains the radio frequency receiving magnetic field and transmitting magnetic field distribution data after the parameterized image data is specially processed as follows.

[0082] Specifically, taking fast spin echo imaging as an example, in the rotating coordinate system, the magnetic vector system precesses right-handedly. Assuming that the initial free induction decay signal after the RF excitation pulse with a deflection angle of 90 degrees and a phase of 0 degrees is t = 0, at which time the magnetic vectors all point to the positive direction of the y' axis, that is, they are in the same phase. After half the echo interval TS / 2, the magnetic vector system has a certain degree of dephasing. Assuming that the magnetic vector precesses at an average speed along the -x' axis at t = TS / 2, a 180-degree deflection angle refocusing pulse is applied at this time, and the magnetic vector system changes its direction and the average speed magnetic vector is along the x' axis. The directions of other magnetic vectors are arranged in reverse order according to the speed of precession, that is, they obey the principle of spin echo refocusing pulse. After TS / 2, the magnetic vector system is phase-coherent again (the first time is when the 90-degree pulse ends), and the echo signal reaches a peak value. At this time, the magnetic vector system as a whole points to the positive direction of the y' axis. If the echo chain length is n, it means that n magnetic vector system refocusing will occur, that is, n echo signals with different amplitudes will be generated.

[0083] The present invention follows the common practice of setting the echo time TE of the fast spin echo scanning sequence, takes an even number n value, sets TE=TS*n / 2, and fills the sampling data of the k=n / 2th echo into the central position of the k-space. In the above-described example, the multi-echo generation process in which the magnetic vector system is initially along the positive direction of the y' axis, dephases at the -x' axis at TS / 2, and then re-phases along the positive direction of the y' axis after TS / 2, the average velocity magnetic vector does not necessarily need to rotate 90 degrees, and the multiple dephase and re-phase conditions after other specific rotation angle values ​​are also applicable to the magnetic field measurement of the present invention. In the process of each dephase to the maximum degree and refocusing to the complete same phase of the magnetic vector system, the relative phase between the magnetic vectors is the object of receiving the magnetic field measurement, rather than measuring the total phase accumulated by the transverse magnetic vector rotating in the complex plane in one sequence cycle. The phase difference between transverse magnetic vectors corresponding to the maximum degree of phase dispersion during each half echo time corresponds to the opposite number of the phase difference at the beginning of the phase gathering period signal acquisition, which is the initial phase of the physical quantity of equivalent proton density of magnetic resonance imaging. The present invention further derives the phase of the measured receiving magnetic field and the transmitting magnetic field from the initial phase. Depending on the length of the set half echo time, the initial phase distribution of the magnetic vector system may be greater than 2π. At this time, it is necessary to perform phase unwinding after obtaining the spatial distribution of the initial phase of the equivalent proton density, and then implement the calculation processing of the receiving magnetic field and the transmitting magnetic field phase.

[0084] The parameter group (α, TE, TR) of the scanning pulse sequence implemented in an imaging examination controls the process of the precession of the magnetic vector system. The imaging scan based on the spin echo sequence also needs to consider the influence of the radio frequency refocusing pulse on the precession of the system. Referring to the change law of the magnetic vector amplitude and phase after the radio frequency excitation pulse acts on the system given by formula (6), after the refocusing pulse with a flip angle of 180° acts on the magnetic vector system with a certain degree of dephasing, the transverse magnetization intensity vector is:

[0085] Mrf,⊥ =M i,⊥ e iπ (19)

[0086] In the formula, the subscript 'rf' indicates refocusing, and 'i' indicates initial before flipping. That is, the amplitude of the transverse component of the magnetic vector does not change, but the phase has an inverse order with equal phase difference between the fast and slow precessing magnetic vectors; for the longitudinal component, the 180° refocusing pulse does not change its amplitude. Therefore, unlike the effect of the excitation pulse deflecting the steady-state magnetic vector to produce a coherent initial transverse component, the use of a refocusing pulse in a non-oblique flip plane can meet the conditions of "maintaining the angular amplitude ratio" and "consistent amplitude and constant angle difference" of the magnetic vector precession process required by the radio frequency magnetic field measurement scanning sequence of the present invention. In addition, when the phase encoding and frequency encoding gradient polarities are constant, the formation mechanism of the echo signal determines that each 180° pulse refocusing period repeats a strictly symmetrical process in which the same specific fast-moving magnetic vector first lags behind the specific slow-moving magnetic vector, becomes coherent, and then ends ahead of the specific slow-moving magnetic vector. The advantage of this is that on the reconstructed magnetic vector image obtained by inverse Fourier transform, the equivalent spin density amplitude and phase distribution based on voxel measurement obtained between different sequences and the same sequence scan are highly consistent, which simplifies the parameter-dependent Bloch equation describing the precession of the magnetic vector.

[0087] For ordinary or fast spin echo sequences, due to the limited repetition time TR, it is generally impossible to reach a target human tissue longitudinal relaxation time T1 greater than 4 to 5 times in imaging, so after a number of equally spaced radio frequency excitation pulses, the steady-state longitudinal magnetization intensity vector Mss will reach saturation, that is, less than the maximum ambient thermal steady-state longitudinal magnetization intensity M0 before the start of scanning. In order to adapt to the conditions of measuring the magnetic field amplitude and phase distribution of the present invention, the sequence repetition time needs to be greater than or equal to 4 to 5 times the tissue transverse relaxation time T2, so that the residual transverse steady-state free precession M'ss during the excitation cycle is minimized. It is known that in the cranial brain, the T1 and T2 reference values ​​of cerebrospinal fluid, gray matter and white matter are 3050ms and 323ms, 1300ms and 99ms, 830ms and 82ms, respectively. Therefore, the sequence TR can be set to 2000ms. Although the longitudinal magnetic vector cannot be completely restored, it can achieve the effect of the longitudinal magnetic vector reaching a new steady state with uniform amplitude after several excitations, thereby ensuring the consistency of the longitudinal magnetic vector from beginning to end during the TR cycle, so as to effectively measure the magnetic field distribution. Regarding the serial echo time TE, there is no need to make the transverse magnetic vector coherence of the tissue fully dissipate before the refocusing pulse is implemented, that is, TE≥T2*, T2* is the magnetic sensitive transverse relaxation time, which is shorter than the intrinsic spin-spin transverse relaxation time T2; because the head and tail signals between adjacent echoes before and after TE time overlap in one cycle, it does not affect the refocusing pulse after flipping the transverse vector and keeping the phase difference between vectors unchanged to obtain the amplitude consistency between continuous echoes. Fig.32The figure shows a schematic diagram of the formation of new steady states of longitudinal and transverse vectors under equally spaced radio frequency excitation. In Fig.32 Figure (a), four radio frequency excitation pulses are continuously applied at equal time intervals TR. Since TR is shorter than 5 times the longitudinal relaxation time T1 of the imaging object, after the first excitation, the longitudinal vector recovers but does not fully recover to the amplitude of the thermal equilibrium steady state M0. After the second excitation and so on, the longitudinal component deflected after the same TR time has a lower amplitude than that of the previous steady state. After several excitations, a new steady state Mss of the longitudinal vector lower than M0 is formed. In Fig.32 Figure (b), the formation of residual transverse steady state free precession during the pulse sequence scanning is described. The same four equally spaced pulses act on the target object. Since the subsequent excitations are not fully coherent with the transverse magnetic vector, that is, TR < T2*, there is always circularly polarized steady state precession M’ss in the transverse direction. In spin echo-like sequences, the existence of M’ss does not affect the acquisition of magnetic resonance imaging signals, the filling of k-space, and even the Fourier inverse transform, etc.

[0088] The physical quantities required to be measured in the magnetic resonance electromagnetic property tomography (MREMPT) of the present invention, namely the radio frequency excitation magnetic field and the receiving magnetic field, whose amplitudes (sensitive fields) and phases can be calculated from the amplitude and phase data in the image domain obtained by the conventional magnetic resonance Fourier inverse transform, that is, the amplitudes and phases of the baseband signals after removing the Larmor frequency carrier. The signal acquisition and detection devices in the magnetic resonance imaging scanning equipment play roles such as gain, filtering, intermediate frequency conversion, digital conversion, and quadrature detection, and specifically include several cascaded links from the receiving coil, preamplifier, image rejection filter, downconverter to the analog-to-digital converter, quadrature phase-sensitive detection, and digital filter. Although there are many links, its core purpose is spectrum shifting, removing the carrier frequency, and demodulating the amplitude A and phase information. A typical magnetic resonance raw signal S0 can be expressed as:

[0089]

[0090] where A is the signal amplitude, ω0 is the resonance frequency related to the main magnetic field B0 provided by the magnet, δω represents the precession resonance frequency deviation of the nuclear spin system, that is, the frequency offset of voxel spatial encoding in imaging, is the initial phase of the lumped or voxel contribution signal, that is, in Equation (8). After detection, the in-phase and quadrature components S I = Acos(δωt + φ) and S Q=-Asin(δωt+φ). Since the physical meaning of the amplitude and phase data information of the magnetic resonance image obtained by inverse Fourier transform is the amplitude and phase of the transverse component of the rotating precession magnetization intensity vector, the amplitude and phase data in the image file can be used to process and obtain the distribution of the RF transmission and reception magnetic field measurement. Fig.31 The figure is a functional diagram of a typical cascade unit of a magnetic resonance imaging system receiver. The original magnetic resonance signal from the detection coil first passes through a pre-radio frequency amplifier to improve the signal-to-noise ratio within the resonance frequency band, then passes through a variable gain amplifier to adjust the dynamic range of the signal to match the range of the analog-to-digital converter, and then passes through a bandpass filter to further suppress the noise outside the band, preparing for mixing and anti-aliasing in the digital domain. After bandpass filtering, the magnetic resonance signal is converted into a 14-bit digital signal by the analog-to-digital converter. The signal power is divided into two paths for orthogonal phase-sensitive detection, and the local oscillation signals used are cos(2πfIFt) and sin(2πfIFt) respectively. The two detected signals are passed through a digital low-pass filter to finally obtain the real and imaginary parts of the baseband k-space data. The data is subjected to inverse Fourier transform to obtain the amplitude and angular distribution of the transverse component of the precession magnetic vector in the image domain, that is, the original magnetic resonance measurement equivalent proton density image data. Based on the proton density data, the amplitude and phase distribution of the RF transmitting magnetic field and the receiving magnetic field can be calculated.

[0091] Based on the fast spin echo signal acquisition scanning pulse sequence, magnetic resonance imaging of human tissue is performed to measure the spatiotemporal evolution of the magnetization intensity vector in the receiving magnetic field induced by the detection coil, and then the measurement of the transmission magnetic field distribution is calculated based on the receiving magnetic field, such as cranial brain, breast or prostate imaging. In order to achieve the measurement of magnetic field distribution and correct the amplitude and phase image distortion caused by the coil magnetic field inhomogeneity in the tissue signal, it is necessary to scan a set of data on the homogeneous water phantom with a sequence group of the same parameters to obtain the corresponding signal amplitude and phase image data. Specifically, the human tissue is scanned three times, and the pulse sequence parameters are set to (90°, TE, TR), (45°, TE, TR) and (45°, TE / 2, TR), respectively, to obtain amplitude image and angular image data, respectively. The water phantom is scanned in the same way to obtain 3 groups of reference amplitude and angular image data, which are listed together in Table 1 of the original image data acquired by magnetic resonance imaging.

[0092] Table 1

[0093] sequence Amplitude data Argument data organize I(90°,TE,TR) m_9_h a_9_h II(45°,TE,TR) m_4_h a_4_h III(45°,TE / 2,TR) m_t_h a_t_h Water Model IV(90°,TE,TR) m_9_p a_9_p V(45°,TE,TR) m_4_p a_4_p VI(45°,TE / 2,TR) m_t_p a_t_p

[0094] In Table 1, m represents magnitude, a represents angle, h represents human tissue, and p represents phantom.

[0095] In order to facilitate the processing of magnetic field data, formula (8) is rewritten as follows:

[0096]

[0097] In the formula, according to its definition, the present invention assumes that M0 is only determined by Ns(r) in measuring magnetic field and inversion characteristics, that is, it is assumed that the main magnetic field is uniformly distributed before measuring the magnetic field, B0(r)=B0 and the imaging body temperature is uniformly distributed, T s (r) = T s 0. Considering the unavoidable influence of tissue contrast, sequence time parameters, etc. on imaging in actual scanning, the magnetic field measurement data, i.e., magnetic resonance signal intensity SI, for the spin echo sequence SE can be written as follows according to the Bloch equation:

[0098]

[0099] In the formula, the parameters related to voxel characteristics include magnetization intensity vector, excitation deflection angle and refocusing flip angle, longitudinal and transverse relaxation time, and receiving sensitive field S(r), i.e., unit receiving magnetic field. It can be seen that compared with formula (21), formula (22) adds a contrast term, i.e.:

[0100]

[0101] Under the condition that the two sequence parameters, especially TE, TR and β, are equal or consistent with each other, the numerator and The factor term can be eliminated by image data division like the term with the same weight M0(r), for example, β≡180° in the present invention. The denominator term has a nonlinear coupling relationship In order to make the magnetic field measurement method of the present invention effective and feasible, TR>>T1 is required so that the denominator is always approximately equal to 1. Under 3.0T, taking the cerebrospinal fluid tissue T1=3050ms with the most serious potential error as an example, when TR=10000, If α is 60° or 120°, the corresponding term value is further halved, and C(r) is approximated to:

[0102]

[0103] Therefore, when calculating the amplitude distribution of the transmitting magnetic field and the receiving magnetic field from the original magnetic resonance measurement image data, the sequence parameter TR>>T1 condition is implicit, and only the sinα(r) term is retained to participate in the entire calculation process.

[0104] It should be pointed out that the significance of the setting of the special values ​​of the deflection angle, echo time and repetition time in Table 1 is that, combined with the evolution of the transverse and longitudinal components of the magnetic vector in magnetic resonance imaging, the transverse components of the magnetic vector in a synchronous cycle of two different deflection angles α, the same TE and TR sequences maintain equal angular distribution and divisible differential amplitudes, while the transverse components in the two same α, doubled TE and equal TR sequence cycles maintain a constant angular distribution, and the target receiving magnetic field phase is obtained by simply multiplying and restoring the single phase difference. In the sequence, for the T2-weighted imaging magnetic field measurement sequence, setting a completely consistent repetition time TR ensures that at the beginning of each cycle, regardless of the sequence flip angle, the quasi-approximate longitudinal steady state of the magnetic vector system returns to the same, that is, the longitudinal magnetization intensity is saturated, and the transverse coherent magnetization intensity in the quasi-approximate steady state can decay naturally. In addition to not meeting the condition that the repetition time is greater than 5 times the transverse relaxation time of the tissue, the T1-weighted imaging sequence is not suitable for measuring the magnetic field because the pixel signal noise ratio is low. After the division operation of the received magnetic field, the original high-noise voxel result in the tissue area becomes an uncertainty value. Using the "amplitude ratio and phase difference" processing method, for a set of deflection angle, echo time and repetition time parameter settings, combined with the homogeneous water model measurement, a set of received magnetic field and transmitted magnetic field distribution measurement results can be obtained. By changing different sequence parameters, multiple sets of magnetic field distribution measurement data for the same imaging body can be obtained. Although they are multiple sets of different measured magnetic fields mathematically, the solutions correspond to the same type of electromagnetic characteristic distribution.

[0105] When performing magnetic resonance imaging scanning to collect data for measuring magnetic fields, the special feature of fast spin echo compared to ordinary spin echo sequence is that the number of echo times and the position of the selected echo time in the echo sequence must be considered, because the sampling data of the line group echo signal fills the k space in order, which affects the accuracy and consistency of the magnetic field phase measurement. In one embodiment of the present invention, when scanning images with a fast spin echo sequence, the number of echo times set for sequences II and V is 1, the echo chain length is 32, and the echo time is 120ms, and the number of echo times set for sequences III and VI is also 1, the echo chain length is still 32, but the echo time is set to 60ms. In this way, the echo time of sequences III and VI is half of that of II and V. Although the instantaneous amplitude of the echo of sequences III and VI is higher than that of II and V, because the deflection angle is the same, the two groups of sequences follow the same specific e-exponential T2 attenuation quantitative law, so that the measurement and calculation of the receiving and transmitting magnetic fields have the prerequisite of "amplitude consistency and constant phase difference".

[0106] Regarding the setting of the receiving bandwidth in the magnetic resonance imaging scanning sequence parameters, the receiving bandwidth is related to the frequency encoding step number. In one embodiment of the present invention, the sampling frequency encoding step number is 128. If the total acquisition time of the complex sample points in the readout gradient time is 5.12ms, the single sample time is 40μs, the total receiving bandwidth is 25kHz, and the bandwidth allocated to each voxel is about 195Hz; in the k-space domain, the time domain frequency encoding step number of 128 means that there are 128 data points in the readout gradient direction. In the embodiment, the frequency encoding direction field of view FOV is set to 256mm, then the pixel width is 2mm, and the corresponding k-domain field of view k PE_FOV Equal to 0.5mm -1 , and Δk PE =1 / 256mm -1 It should be pointed out that the setting of the image domain field size does not affect the measurement of the magnetic field. However, because the receiving bandwidth affects the signal-to-noise ratio of pixel signal acquisition, the Bandwidth needs to be adjusted to ensure that the Rel.SNR of different sequences is consistent, which is conducive to calculating the magnetic field amplitude under the same conditions.

[0107] The use of spatial encoding magnetic field gradients, like the above-mentioned image geometry setting parameters, only affects the size of the field of view and resolution, and does not affect the measurement of the magnetic field based on the pixel signal. In the two-dimensional frequency encoding layer selection gradient mode, the layer thickness is proportional to the derivative of the product of the gradient and the pulse width, and the spatial resolution of the frequency encoding readout gradient and the cyclic execution phase encoding gradient corresponding to a single scanning tomography is also proportional to the inverse of the product of the gradient and the acquisition time (for phase encoding, it is the maximum value of the gradient). Therefore, it can be seen that the radio frequency magnetic field measurement of the present invention has nothing to do with the image space feature results such as the specific size and resolution generated by the setting of the main static magnetic field and the magnetic field gradient.

[0108] According to Faraday's law of electromagnetic induction, the time-varying magnetic field generates a vortex electric field with a spatial gradient distribution, which is an electromotive force different from the Coulomb electric field. Lenz's law shows that the resistance to changes in magnetic flux density is the source of coil collision noise during the execution of gradient pulses. Due to the permittivity of human tissue, the rising and falling edges of the magnetic field gradient pulse will also induce eddy currents, which are manifested in the phase image as spatial domain second harmonic interference. The correction method can be to perform another set of scans that change the layer selection, phase and frequency encoding gradient polarity on the basis of sequences I to VI. Before calculating the phase of the receiving magnetic field, the phase time of the normal scan and the reverse polarity scan of each sequence is averaged to obtain more accurate initial phase data that eliminates the influence of eddy currents.

[0109] According to the different ways of generating the layer selection gradient, the existing 2D and 3D imaging modes of magnetic resonance imaging provide a good method for the radio frequency magnetic field measurement of the magnetic resonance electromagnetic property imaging of the present invention. Based on the frequency-encoded layer selection gradient and the selective excitation radio frequency pulse, the original equivalent spin density measurement data of the tomographic voxel array obtained has no necessary correlation in the layer direction, so the corresponding calculated receiving magnetic field and transmitting magnetic field can only be applied to various existing electrical property imaging inversion algorithms of 2D approximation and the electromagnetic property imaging algorithm of the present invention. However, the spin density measurement data of the 3D pixel matrix obtained by the magnetic resonance phase-encoded layer selection gradient and the non-selective radio frequency pulse based on the volume imaging method has a natural correlation between the layers after reconstruction, and the receiving magnetic field and transmitting magnetic field distribution data obtained by calculation also have inter-layer correlation, so it can be applied to the electromagnetic property imaging inversion method that is more in line with the general physical meaning.

[0110] At present, the 2D imaging methods on medical magnetic resonance equipment mainly include technologies such as gradient echo, ordinary spin echo, echo plane and fast spin echo, while 3D imaging includes technologies such as fast spin echo and gradient echo. In 2D imaging technology, the scanning time of ordinary spin echo sequence is too long. Although the echo plane imaging sequence provided by the manufacturer is so-called based on spin echo, and there is also a phase encoding gradient polarity regulation and modification function, spin echo-echo plane imaging is essentially based on gradient echo, and if the number of excitations is not 1, the strip data combination artifacts related to the number of excitations on the reconstructed angular image are serious. Under the condition that there is no gradient polarity modification function provided by the manufacturer, although eddy current artifacts are inevitably present on the angular image of the fast spin echo sequence scanning, by setting a lower sequence parameter receiving bandwidth, a smaller readout gradient intensity can be used, which can make up for the inability to compensate for eddy current interference to a certain extent. Therefore, the present invention preferably uses fast spin echo sequence technology as an imaging technology for scanning equivalent spin density distribution. In the 3D fast spin echo technology, there are executable sequences such as CUBE and basic fast spin echo; CUBE is a high-speed scanning sequence specially designed by the manufacturer for isovoxel imaging, and is not suitable for the imaging of the present invention because of the potential change of the RF deflection angle during its pulse cycle. For 3D scanning that maintains the correlation of the measured quantity data between layers, the present invention preferably uses the basic fast spin echo sequence to measure the equivalent spin density and magnetic field distribution.

[0111] In the following data processing process for calculating the receiving magnetic field, it is relatively reasonable to assume that the magnetic permeability of the tissue is equal to the magnetic permeability of the standard homogeneous water model solution, so as to satisfy the condition that the value of the sine function of the transmitting magnetic field can be negated. Therefore, in order to maintain consistency, in the process of electromagnetic property inversion, the electrical characteristic parameters assuming constant magnetic permeability can be inverted with the measured quantity of the receiving magnetic field distribution, and the magnetic properties can be inverted with the transmitting magnetic field. In addition, in the measurement of the magnetic field data generated by the transmitting coil and induced by the receiving coil, there are two different types of differentiated source contributions, one is the natural distribution of the electromagnetic properties of the human body itself, and the other is the non-uniform distribution of the magnetic field generated by the non-ideal coil, which are respectively represented by B and B. o (r) and B c (r), where o stands for object and c stands for coil.

[0112] In the following processing to obtain the operation of measuring the magnetic field, the processing goal is to obtain the magnetic field amplitude and phase data information of the human body about the main parameter sequence, and the excitation and acquisition sequence parameters of the main sequence are given as (45°, TE, TR). For the sequences listed in Table 1, when calculating the magnetic field sensitivity field amplitude distribution, sequences II and V correspond to the main sequences of the tissue and water model, respectively, and sequences I and IV are amplitude measurement auxiliary sequences. When calculating the phase distribution, sequences III and VI are phase measurement auxiliary sequences of the tissue and water model, respectively.

[0113] 1. Calculation method of receiving magnetic field measurement data

[0114] (1) Amplitude component

[0115] First, the process of calculating the sequence I and II data of the processed tissue to obtain its receiving magnetic field sensitivity field distribution is as follows, and the equivalent spin density that eliminates M0(r) is equal to:

[0116]

[0117] Where r is the voxel position coordinate, γ is the gyromagnetic ratio, the superscript '90' represents the excitation deflection angle, the subscript 'h' represents human tissue, μ h is the magnetic permeability of human tissue, τ is the RF excitation pulse width, and the upper left subscript 'TE' indicates It is the receiving magnetic field under special TE setting. The top mark '~' indicates the alternating variable. The upper right mark '+' and '-' indicate transmission and reception respectively. The lower right mark '1' of the symbol B indicates the transmitting magnetic field, which is different from the lower mark '0' in B0 indicating the main magnetic field. In theory, the RF excitation coil transmits the magnetic field The amplitude of is uniformly distributed in space, that is, In the following text, there are Compared to the receiving magnetic field The underscript '_' indicates pluralization, which is also In addition, the excitation radio frequency pulses in the sequence of the present invention all adopt the same pulse width, and the deflection angle is modulated by changing the pulse amplitude to ensure that the spin system of different α, the same TE and TR sequences precess synchronously during the scanning process, that is, to ensure that the precession phase is always equal.

[0118] Magnetic resonance imaging requires Uniform distribution, according to B(r) = μ(r)H(r), its non-uniformity comes from the possible μ(r) and H(r) parts. In the following, the magnetic permeability will be inverted from the transmitted magnetic field based on this relationship. In order to further process the sensitive field distribution of the received magnetic field In the following derivation, keep Expand it into two parts, μ and H.

[0119] Then, the homogeneous water model is scanned in the same way with the measurement sequence IV parameters (90°, TE, TR) and sequence V parameters (45°, TE, TR), and the unit receiving magnetic field distribution expression containing the specific trigonometric function weights of the transmitting magnetic field under the same conditions is obtained:

[0120]

[0121] In the formula, μ p,0 is the magnetic permeability of the phantom water model. The subscript 'p' indicates the water model, and the subscript '0' indicates that the magnetic permeability is constant because the water model is homogeneous. It is only related to the non-uniformity of the receiving coil magnetic field c, where c represents the coil. The homogeneous characteristics of the water phantom are utilized to obtain the image distribution of the specific trigonometric function weights related to the transmitting magnetic field also contained in the tissue scanning data.

[0122] In order to further eliminate the inhomogeneity of the transmitting magnetic field of the excitation coil and the inhomogeneity of the receiving magnetic field related to the detection coil, the ratio of the equivalent proton density distribution of the imaged human tissue and the phantom is obtained as follows:

[0123]

[0124] When the magnetic permeability μ of human tissue is considered h (r) is approximately equal to the magnetic permeability μ of the homogeneous contents of the model p,0 , the above formula can eliminate the contribution of the non-uniformity of the transmitting coil magnetic field. Under the scanning conditions of specific sequence parameters (TE, TR), the receiving magnetic field sensitivity field distribution is:

[0125]

[0126] Here, the obtained parameters related to the specific parameters used (TE, TR) In the present invention, the inherent non-homogeneous imaging volume proton density coupling is eliminated in the magnetic resonance signal. In general, compared with various contrast imaging techniques such as T1W, T2W and PDW in clinical practice, the magnetic field measurement in the present invention is not limited to a certain weighted imaging sequence determined by (TE, TR), and even the tissue sensitivity and specificity reflected by the receiving magnetic field sensitive field distribution can be combined with the existing magnetic resonance system image results to assist in pathological differential diagnosis.

[0127] (2) Phase component

[0128] First, we analyze the precession process of a single spin in a tissue or water phantom, especially how the receiving magnetic field is formed during spin echo scanning sequence imaging. The magnetization intensity vector in the receiving magnetic field precesses at the Larmor frequency, and the transverse magnetization intensity component at the initial stage of signal generation M 0,⊥ The initial phase is With the excitation emission magnetic field The initial phase The phase difference is π / 2, and then a free induction decay signal is generated. During the scanning process, the echo signals formed again and again contain spatial encoding, which is the object of the detection coil signal acquisition, that is, the receiving magnetic field of the present invention. It should be noted that the instantaneous phase during the echo formation, including the front and back phase convergence and phase dispersion, is exactly the relative phase distribution displayed by the image domain phase image obtained by the inverse Fourier transform. Although sometimes the relative phase may exceed the 2π range due to the long acquisition time and needs to be unwound, the imaging measurement phase has nothing to do with how much phase a single spin has accumulated during the entire TR period, and has nothing to do with the phase at the moment nω0TE.

[0129] In order to measure the phase distribution of the receiving magnetic field during the imaging scan of the human body, it is necessary to eliminate the contribution of the transmitting magnetic field phase in the original magnetic resonance signal equivalent spin density phase distribution data. Similar to the measurement method of the RF magnetic field amplitude, the same imaging body is scanned twice, that is, involving main sequences II and V and auxiliary sequences III and VI. The same excitation deflection angle of 45° is implemented to obtain the same M 0,⊥ The initial phase The echo times of the two groups of sequences with the same echo signal type are in a fixed ratio during the corresponding cycle execution period. Since the transverse relaxation time constant of the voxel determines the relaxation rate, the precession angular frequency is the same under the same gradient geometry conditions, and the two scans have a constant phase difference related to TE for the same position. The same repetition time TR of the two groups of sequences can ensure that the magnetization intensity vector starts with almost the same amplitude and azimuth state, the transverse component decreases and tends to end in the current sequence cycle, and the longitudinal component recovers and tends to steady state in the next sequence cycle. In the present invention, according to the sequence implementation plan of Table 1, the goal of phase-related data processing is to obtain the main parameter sequence (45, TE, TR) scanning and measuring the radio frequency magnetic field phase distribution of the human body, and the specific data processing process is as follows. The phase difference is obtained by subtracting the sequence III angle data from the sequence II angle data:

[0130]

[0131] In the formula, and are the receiving and transmitting magnetic field phases, respectively. The superscript '45' indicates the excitation deflection angle, and the superscripts 'TE' and 't' correspond to the long echo time and short echo time series, respectively. The difference removes the transmitting magnetic field phase distribution constant and contains information about the non-homogeneity of the test subject and the non-homogeneity of the receiving coil magnetic field.

[0132] The water phantom is scanned with the main sequence V and auxiliary sequence VI of the magnetic field phase measurement under the same parameter conditions. Since the working fluid in the phantom is uniform, the non-uniformity of the magnetic field phase of the receiving coil can be obtained. Specifically, the contribution of the magnetic field phase of the excitation pulse emission is eliminated by subtracting the two sets of phase data. Under the conditions of the flip angle, echo time and repetition time being (45°, TE, TR) and (45°, TE / 2, TR) respectively, the phase difference of the two sequences of the water phantom is:

[0133]

[0134] In the formula, the subscript p represents the water model phantom. p It only reflects the non-uniformity of the phase distribution of the receiving coil magnetic field, and this distribution also exists in the phase difference of the human body, so Δp h Further amendments include:

[0135]

[0136] So far, for the same voxel and the same spatial encoding gradient conditions, because the echo time parameter value of the auxiliary sequence is 1 / 2 of the main sequence, the relative distribution of the received magnetic field phase measured by scanning human tissue with the main sequence (45°, TE, TR) is obtained. 1 / 2, then:

[0137]

[0138] It is worth mentioning that compared with the derivation calculation of the receiving magnetic field sensitivity field, which assumes that the human body magnetic permeability is homogeneous and water-modeled due to the presence of a nonlinear sin function acting on the transmitting magnetic field, the method for solving the single-fold receiving magnetic field phase quantity Δp of the human body tissue h In the calculation of After being independent of the magnetic properties of the tissue, relevant subtraction operations are performed.

[0139] 2. Calculation method of transmission magnetic field measurement data

[0140] (1) Amplitude component

[0141] The magnetic resonance signals of the main sequence II and the auxiliary sequence I are still measured by amplitude, that is, the equivalent proton density, and the calculated receiving magnetic field sensitivity field data corresponding to the unit current generated in the detection coil. Under the conditions of sequence time parameters (TE, TR), the emission magnetic field sensitivity field distribution data of the imaged human tissue can be calculated by the following steps. According to the formula (6) of the macroscopic magnetization intensity vector, two sets of ratio-type intermediate quantities are defined:

[0142]

[0143]

[0144] Where M0 is the macroscopic magnetization intensity vector amplitude of the imaged human tissue, the superscript '90' represents the sequence excitation deflection angle, and 'TE' indicates that the field quantity B is measured at a specific echo time TE.

[0145] In order to further eliminate the magnetization intensity in the result, that is, to finally eliminate the proton density coupling, the line t1(r) / t2(r):

[0146]

[0147] Comparing with formula (9), we can get:

[0148]

[0149] Wherein, t1 and t2 are the intermediate quantities of equations (33) and (34). It can be seen that, similar to the existing dual-angle method, the present invention also uses the arc cosine function to solve the distribution of the transmitting magnetic field sensitivity field. And in order to satisfy the relationship of the trigonometric identity sin 2θ=2sinθcosθ, the deflection angles of the main sequence and auxiliary sequence of the sensitive field measurement are taken as typical 90° and 45° respectively.

[0150] Regarding the excitation RF pulse width τ, MRI manufacturers usually control the deflection angle by changing the amplitude of the transmitting coil driving current rather than the pulse width. This makes it easier to calculate and adjust the echo time and repetition time when executing a multi-echo scanning sequence to maintain consistency between different sequences. In actual situations, considering the saturation effect of the magnetic vector, the current amplitude ratio of 2 times the deflection angle is about 1.9. If you want to obtain the transmit magnetic field sensitivity field distribution data of the amplitude sequence Can be calculated from the main sequence It can be approximated by multiplying by 2. It can reduce the number of imaging scans and the need for multiple sets of magnetic field measurements for certain electromagnetic characteristics imaging. Constrained by the unit relationship between radians and gyromagnetic ratio and pulse width magnitude formula, the magnitude of the transmitted magnetic field sensitivity field is in μT.

[0151] (2) Phase component

[0152] Distribution of sensitive field of the same emission magnetic field The phase distribution of the receiving magnetic field for human tissue scanning is calculated using the main sequence II and V and the auxiliary sequence III and VI. Then, according to the equivalent proton density formula (21) related to the inverse Fourier transform, the measured emission magnetic field phase distribution based on the sequence parameters (45°, TE, TR) is calculated as follows:

[0153]

[0154] In the formula, the phase difference π / 2 comes from the transverse component of the magnetic vector M 0,⊥ The phase is equal to the phase of the magnetic field emitted by the RF excitation pulse plus a constant π / 2, which is equivalent to M0 multiplied by the imaginary unit i.

[0155] So far, the imaging scanning pulse sequence parameters (45°, TE, TR) are used as the means of measuring the radio frequency magnetic field, and the image data is collected based on the original equivalent spin density distribution of magnetic resonance imaging, realizing a complete method for deriving and calculating the sensitive field and phase information of the excitation transmission magnetic field and the detection of the receiving magnetic field. In addition to assuming that the human tissue is homogeneous in magnetic permeability in the calculation of the amplitude of the receiving magnetic field sensitive field, the formulas for the calculation of the transmitting magnetic field sensitive field and the transmitting and receiving phases are strictly derived according to the amplitude and phase laws of electromagnetic wave propagation. In addition, in the process of using the measurement data of the homogeneous water model to assist the calculation of the human tissue magnetic field measurement, the effect of eliminating the non-uniformity of the magnetic field of the receiving coil and the transmitting coil is additionally obtained.

[0156] For the measurement of the transmit magnetic field sensitivity field amplitude of the present invention, when sequences I and II and IV and V adopt deflection angles of 90° and 45°,

[0157] 1=sin90°=2sin45°cos45°=2×0.7071×0.7071 (38)

[0158] Due to the limitation of the arccosine function's independent variable value range (0,1), beg The dynamic range is small; and there is a trend that the smaller the single flip angle of the corresponding target, the smaller the dynamic range. Therefore, when the actual imaging scanning measures the magnetic field, according to the value rule of the trigonometric function, the sequence uses 120° and 60°, that is:

[0159] 0.866=sin 120°=2 sin 60°cos 60°=2×0.866×0.5 (39)

[0160] Under the condition of a single refocusing pulse of a conventional spin echo SE sequence, according to the Bloch equation describing the spin precession process, the signal intensity formula (22) generated at a specific position r in the magnetic resonance receiving coil is rewritten as:

[0161]

[0162] Wherein, M0 is the initial magnetization intensity of the environment thermal equilibrium state, α is the excitation deflection angle, β is the refocusing flip angle, TR is the sequence repetition time, TE is the echo time, T1 is the longitudinal relaxation time of the target human tissue, T2 is the transverse relaxation time, and S is the receiving sensitive field, that is, the unit receiving magnetic field amplitude of the present invention. It can be calculated from the formula that when α is equal to 120° and 60° respectively and β is equal to 180°, the intensity SI is equal to:

[0163]

[0164]

[0165] That is, the intensity at a deflection angle of 120° is greater than that at 60°. Under the condition that the echo time and repetition time are consistent, although the deflection angle of double excitation exceeds the transverse plane, there is a certain amount of power waste from the perspective of signal-to-noise ratio, and the sequence group deflection angle 120° and 60° schemes produce the same amount of initial transverse magnetization components, but from the comparison of signal intensity, the latter is greater than the former, that is, it still continues the relationship that the double angle intensity is greater than the single angle in the 90° and 45° schemes.

[0166] 3. Magnetic resonance electromagnetic property imaging inversion method

[0167] Compared with Gauss's law and Gauss's magnetic law, which describe the interaction between matter and a single source electric field or magnetic field, Faraday's law of electromagnetic induction and Maxwell-Ampere's law establish the interaction relationship between matter and electromagnetic field, including electromagnetic waves. In order to invert the electrical and magnetic property parameters of a material using the radio frequency magnetic field distribution measurement under magnetic resonance conditions, the present invention must start from Faraday's and Maxwell-Ampere's laws that contain both electric and magnetic fields, and establish a core calculation equation to solve the electromagnetic properties from known magnetic field quantities, while the electric field quantity is expressed by the magnetic field quantity. In addition, considering that there is a complex-valued magnetic flux density vector known quantity in an equation, there are two scalar unknown quantities, namely the complex permittivity of the electrical property and the magnetic permeability of the magnetic property, where the complex permittivity is further divided into conductivity and permittivity.

[0168] Starting from the principle of magnetic resonance imaging, the present invention proposes a method for measuring the distribution of radio frequency transmitting magnetic field and receiving magnetic field based on the same imaging body in two consecutive links of magnetization intensity vector excitation and echo signal detection. The two magnetic fields propagate in opposite directions and have different physical meanings. For imaging scanning, changing the pulse sequence parameters can obtain different measured transmitting magnetic field and receiving magnetic field distributions. Different core calculation equations can be derived according to the different prerequisites of the electromagnetic characteristic (σ, ε, μ) inversion algorithm to ensure that there are enough known quantities of magnetic field measurement in the case of multiple unknown electromagnetic characteristic parameters, and the field quantities are both homologous to the imaging body and linearly independent of each other. In addition, at the computational execution level of the inversion algorithm, based on the linear equations constructed by regularization and least squares method, multiple sets of independent magnetic field measurement data are conducive to reducing the occurrence of coefficient matrix rank deficiency. First, the derived basic equation of related wave propagation expressed in magnetic field without assumptions such as homogeneity of matter is given:

[0169]

[0170] In the formula, for inhomogeneous substances Both and μ are related to the spatial position, so Note that the field quantity H is not a known quantity. Previously it needed to be replaced by (B / μ).

[0171] The complex-valued magnetic field vector B can be orthogonally decomposed into (B x ,B y ,B z ), in the cylindrical coordinate system it can also be decomposed into (B + ,B - ,B z ), where B + is the right-hand component, B - is the left-handed component. In magnetic resonance imaging, the transmitted magnetic field measurement quantity is expressed as The received magnetic field is expressed as The following derivation process uses the magnetic field rotation component in the cylindrical coordinate system to invert the target electromagnetic characteristic parameters. When writing the component form of equation (43), Are of equal status.

[0172] remember To receive magnetic field The conjugate of , rewrite formula (4) and directly use the rectangular coordinate component B of the magnetic flux density x and B y Represents the rotating emission magnetic field and receiving magnetic field

[0173]

[0174] In the formula, B x and B y are all complex-valued amplitudes of vectors, i is an imaginary unit, the superscript '*' indicates complex conjugation, and the subscripts 'x' and 'y' represent the x- and y-direction components of the field, respectively. Regardless of the magnetic resonance electrical properties imaging inversion method, such as the standard Helmholtz equation method for magnetic field vectors that ignores the electrical properties gradient, It is necessary to derive the corresponding scalar core equation related to the rotation component, so as to solve the distribution of the unknown electrical characteristic parameters with the known sensitive field amplitude or phase of the component. Continuing to derive the Helmholtz equation as an example, the derivation of the rotation coordinate system component form equation from the rectangular coordinate system orthogonal field component is as follows:

[0175]

[0176] The linear combination of the equations is:

[0177]

[0178] Furthermore, according to the definition expressions of the magnetic resonance transmitting and receiving magnetic fields:

[0179]

[0180] In the formula, the complex permittivity is the target parameter, given by or The inversion can be obtained based on the voxel data matrix using the least squares method.

[0181] In order to reduce the difficulty of solving equations and obtain relatively accurate electromagnetic properties, combined with a necessary number of assumptions, there are two different technical routes to implement the inversion of complex permittivity and permeability. In existing magnetic resonance electrical imaging, an assumption similar to vacuum permeability is used to invert conductivity and permittivity from the transmitted magnetic field distribution. As mentioned above, among the three mainstream inversion methods, the Helmholtz equation method has reconstruction system errors at the tissue interface due to ignoring the electrical characteristic gradient; the gradientEPT method introduces the characteristic gradient intermediate calculation process, which improves the accuracy of the results, but has problems with the field measurement method; in the algorithm that does not ignore the electrical characteristic gradient, the convection-reaction EPT method derives linear equations combined with the necessary Dirichlet boundary conditions when ignoring the longitudinal component of the radio frequency magnetic field and its gradient, which is the preferred algorithm for solving complex permittivity.

[0182] The first type of radio frequency magnetic property inversion method of the present invention is based on the existing crEPT. First, the complex permittivity parameter distribution of the imaged human tissue is calculated using the transmitted magnetic field distribution data. Then, under the condition that the complex permittivity is known, the inversion of all electromagnetic properties is completed according to another crEPT equation about magnetic permeability based on the received magnetic field data measured in the same scan.

[0183] The second method of the present invention takes into account that adding an unknown quantity to the complex domain partial differential equation about a vector will inevitably increase the difficulty of the solution process. Therefore, a simple Helmholtz-type core equation is derived by combining the necessary assumptions of homogeneity of several electrical and magnetic properties, and the relationship between the magnetic field and the electromagnetic properties is explicitly given. The least squares method is used to fit and solve the complex permittivity and permeability, and then the voxel-based conductivity, relative permittivity and permeability distribution are finally obtained according to the corresponding relationship between the real part and the imaginary part. The technical route of the magnetic resonance electromagnetic imaging magnetic property inversion method is shown in Table 2.

[0184] Table 2

[0185]

[0186] (1) Derivation of the core equations of the first type of characteristic inversion method

[0187] Because relatively speaking, the derivation and calculation of the measured quantity of the received magnetic field distribution contains the assumption that the magnetic properties of the magnetic permeability of the imaged human tissue are similar to the magnetic properties of the homogeneous water model, when the magnetic field data of the present invention is used to deconstruct the permittivity distribution based on the existing magnetic resonance electrical properties imaging inversion method, the transmitted magnetic field distribution data is used preferentially.

[0188] Under the condition of known conductivity and permittivity distribution and one or more sets of received magnetic field data, according to Maxwell's equations, the wave equation containing the unknown permeability and only the magnetic field is:

[0189]

[0190] In the formula, to distinguish it from magnetic resonance electrical imaging, the known quantity of complex permittivity is recorded as If we insist on the non-homogeneity of tissue magnetic permeability, that is, we do not make any assumption that the magnetic permeability gradient is 0, then the process of deriving the core equation of magnetic permeability inversion similar to the convection-reaction partial differential equation of electrical properties is very complicated. The following is the final execution equation result about the receiving magnetic field:

[0191]

[0192] Where v is the inverse of the scalar permeability, and the complex permittivity gradient is defined as where g' x and g' y are the x and y direction components of g' respectively.

[0193] However, making the magnetic permeability uniformity assumption at appropriate links can greatly simplify the derivation process, and is conducive to program writing, reducing the error rate and maintaining the stability of the solution. First, if a magnetic uniformity assumption is made on the magnetic field intensity divergence term, then:

[0194]

[0195] The advantage of the above formula is that there is an unknown quantity μ. If we make the second magnetic homogeneity assumption, we can take μ out and multiply it by the right side, and we have:

[0196]

[0197] It should be noted that after multiplying the right-hand side term, μ returns to the non-homogeneous property. In the same electrical characteristic imaging Helmholtz equation method, when deriving the equation component form of the formula, when using the receiving magnetic field to calculate the characteristics, pay attention to the difference between the receiving magnetic field conjugate and the transmitting magnetic field itself.

[0198] (2) Derivation of the core equations of the second type of characteristic inversion method

[0199] Compared with the first type of inversion method, the second type of characteristic inversion method uses a certain number of homogeneous assumptions to derive a combination of two Helmholtz-type equations about the transmitted magnetic field and the received magnetic field measurement, and further solves the two unknowns of scalar complex permittivity and magnetic permeability at the same time. The core calculation equation is derived from formula (43) without any approximate treatment. The specific process is as follows. Make a homogeneous assumption of the magnetic properties of human tissue and a homogeneous assumption of the electrical properties. Then, according to Gauss's magnetic law, we have and have And let v = μ -1 ,get:

[0200]

[0201] In the magnetic resonance electromagnetic characteristic imaging of the present invention, the right-handed radio frequency transmitting magnetic field and the left-handed receiving magnetic field is the measured quantity, i.e. the known quantity of the above equation. Therefore, using formula (44), the orthogonal component form of the above equation is:

[0202]

[0203] So far, the goal is to solve the unknown quantity in the system of equations Although the two equations in the system of equations have the same form, the two magnetic field measurement known quantities substituted into the calculation are different, that is, the response functions of the same inhomogeneous characteristic source in space under different field excitations are different, so that two unknowns, namely the distribution of electrical and magnetic characteristic parameters, can be solved independently from the two complex valued known quantities. If there are multiple sets of magnetic field measurement known quantities obtained by changing the sequence parameter scan, the well-posedness of the solution can be increased.

[0204] Further investigation Item, can be written as:

[0205]

[0206] At this time, the equation It is suitable to apply the method of convection-diffusion equation to solve it.

[0207] Based on the 3.0T magnetic resonance imaging system conditions, the present invention uses a fast spin echo sequence to scan the equivalent proton density image data of human tissue, further calculates the distribution of the radio frequency transmission magnetic field and the receiving magnetic field, thereby inverting the distribution of the electrical and magnetic characteristics of the tissue and reconstructing it. For cranial imaging under 3.0T conditions, the longitudinal relaxation time and transverse relaxation time reference values ​​of cerebrospinal fluid, gray matter and white matter tissue are 3050ms, 1300ms, 830ms and 323ms, 99ms, 82ms, respectively. In order to basically meet the measurement and scanning requirements of TR>3*T1 and TR>>T2, the preferred parameters of the fast spin echo sequence used in the present invention are (120°, 30ms, 9000ms), (60°, 30ms, 9000ms) and (60°, 15ms, 9000ms). Assuming that the number of points in the frequency encoding and phase encoding directions of the imaging geometric parameters is 128, when the receiving bandwidth Bw is 31.25kHz, the single data sampling time td is 16μs, and the single echo acquisition time tacq is 2.048ms. For the echo time TE = 30ms, and referring to the echo chain length etl 16 in conventional imaging, the echo time interval is calculated to be 3.75ms. Considering the inhomogeneous effect of the magnetic field of the magnetic susceptibility, this interval can fully accommodate tacq; and the total acquisition time of one excitation is 60ms, that is, 2 times TE. Referring to conventional magnetic resonance imaging examinations, the image contrast obtained by the sequence scanning used tends to be proton density weighted. Generally, for cranial brain imaging, the typical repetition time and echo time in transverse relaxation time weighted T2WI and fluid suppression inversion recovery sequence FLAIR are 4000ms and 90ms and 9000ms and 114ms respectively. The total acquisition time of the present invention (120°, 30ms, 9000ms) and (60°, 30ms, 9000ms) is within 3*T2_M, where M represents gray matter and white matter, i.e., within the effective duration of intrinsic relaxation signal attenuation. According to the gradient calculation formula:

[0208]

[0209] In the formula, R RO is the spatial resolution in the voxel readout direction, γ is the gyromagnetic ratio, G RO To read the gradient in the direction. Bandwidth 62.5kHz and field of view 256*256mm 2 Under these conditions, the readout gradient intensity G_RO is only 5.7 mT / m.

[0210] For TE = 15ms (60°, 15ms, 9000ms), under the same etl conditions, the echo time interval is reduced to 1.875ms. When Bw does not change, it depends on the transverse relaxation time T2* value of the tissue's magnetic susceptibility. There may be a certain degree of head-to-tail connection between adjacent echoes, but it does not affect the refocusing pulse flipping the transverse magnetic vector precession phase to form an echo and complete signal acquisition, and further calculate the RF magnetic field distribution from the measured quantity. The total acquisition time is reduced to 60ms. Because the sampling time does not change, the overall signal-to-noise ratio increases due to the forward shift of the echo chain time. In order to maintain the consistency of the signal-to-noise ratio of the acquired signal, according to:

[0211]

[0212] Where t represents the time domain, ω0 is the angular frequency, β is the coefficient related to the deflection angle, and B 1u is the magnetic field generated by the unit current through the receiving coil, that is, dV s is the voxel volume, k is the Boltzmann constant, T is the coil temperature, R c is the resistance, B is the bandwidth, that is, Bw, N RO N is the number of points in the read direction matrix. PE is the number of points in the phase direction matrix, N a is the number of excitations; that is, when other imaging sequence parameters remain unchanged, the bandwidth needs to be appropriately increased to suppress the increase in signal-to-noise ratio caused by the decrease in TE. Increasing the bandwidth of the (60°, 60ms, 9000ms) sequence, that is, reducing td, can reduce the time domain signal-to-noise ratio. Note that when td and tacq are reduced, the system will automatically increase the frequency encoding gradient intensity because the voxel spatial resolution remains unchanged. In this way, the double superposition effect will lead to an increase in the gradient slew rate-eddy current effect, especially the original magnetic resonance angle data of the sequences (60°, 120ms, 9000ms) and (60°, 60ms, 9000ms) used in the calculation of the precession phase, that is, the transmission and reception magnetic field phases are sensitive to eddy currents. Assuming that Bw is increased to 62.5kHz, td=8μs, tacq=1.024ms, and G_RO=11.5mT / m under the condition that the voxel spatial resolution remains unchanged; at this time, the echo interval of 1.875ms can still accommodate tacq. In summary, in order to facilitate the accuracy of magnetic field measurement, a preferred embodiment of the present invention is to set the bandwidth to a narrower value of 31.25kHz for the sequences (120°, 30ms, 9000ms) and (60°, 30ms, 9000ms), leaving an adjustment margin for the short TE sequence to reduce the eddy current interference effect, and the bandwidth of (60°, 15ms, 9000ms) can be adjusted according to the actual value displayed by the system after setting the short TE to obtain a relative signal-to-noise ratio Rel.SNR consistent with the long TE sequence.

[0213] In the longitudinal direction, the set TR is approximately three times greater than the T1 of the cerebrospinal fluid in the tissue with the longest longitudinal relaxation time, which can ensure that the magnetization intensity vector in most tissues is fully restored and basically avoid the influence of the longitudinal partial saturation effect.

[0214] Beneficial effects:

[0215] The present invention can measure the amplitude and phase components of the time-harmonic radio frequency transmitting magnetic field excitation in nuclear magnetic resonance imaging, and the amplitude and phase components of the receiving magnetic field induction detection signal. The present invention overcomes the problem in the prior art that the imaging body spin density in the detection signal cannot be removed from the coupling of the receiving magnetic field sensitive field amplitude, and also overcomes the problem in the prior art that the detection signal cannot be distinguished between the transmitting magnetic field phase and the receiving magnetic field phase. Conversely, after obtaining the receiving magnetic field amplitude, the spin density distribution of human tissue can be calculated, thereby having a far-reaching technical impact on the existing transverse relaxation time weighted, longitudinal relaxation time weighted and proton density weighted imaging principle systems.

[0216] Although the magnetic permeability of human tissue is relatively homogeneous and close to the magnetic permeability of vacuum, the present invention breaks the unrealistic premise assumption of tissue vacuum permeability in existing magnetic resonance electrical property imaging and proposes magnetic resonance electrical property and magnetic property imaging. In clinical medical practice, magnetic property imaging can play a role in assisting imaging differential diagnosis for pathological conditions related to changes in magnetic properties, such as brain tissue microbleeding masses and breast calcification foci.

[0217] In order to evaluate the local microcirculation of human body, the existing magnetic resonance perfusion imaging needs to inject exogenous gadolinium-based contrast agents to change the magnetic properties of target tissues and change the relaxation time to reflect the image specificity. However, its potential incomplete metabolism and deep deposition in brain tissue may bring some side effects. Like arterial spin labeling imaging and blood oxygen saturation level dependent functional magnetic resonance imaging, the magnetic resonance magnetic property imaging of the present invention provides an imaging method based on endogenous magnetic property differences.

[0218] Different from the existing magnetic susceptibility weighted imaging and quantitative magnetic susceptibility imaging, which estimate the static magnetic susceptibility distribution based on the precession phase difference, the magnetic resonance radio frequency magnetic and electrical characteristics imaging of the present invention uses a spin echo scanning sequence instead of a gradient echo sequence to collect magnetic field measurement signals, which can effectively compensate for the influence of the main static magnetic field and gradient magnetic field inhomogeneity on the signal. Combined with the pre-scanning uniform field measures of the magnetic resonance imaging equipment, and the matching homogeneous water model sensitive field and phase distribution measurement image data collected in the electromagnetic characteristics imaging of the present invention, the influence of the magnetic field inhomogeneity of the transmitting and receiving coils on the measurement results is effectively reduced or even eliminated.

[0219] In combination with the magnetic resonance transmitting magnetic field and receiving magnetic field distribution data measurement method provided by the present invention, two technical route methods that consider different physical conditions can be used to invert the magnetic characteristic parameters of human tissue. The first method is to first use the existing magnetic resonance electrical characteristic imaging gradient inversion algorithm to solve the electrical characteristics, and then use the receiving magnetic field as the measurement quantity to invert the magnetic characteristics by the magnetic characteristic inversion core equation. The magnetic inversion has different core equations according to the homogeneous conditions. In the second method, the difficulty of solving the algorithm core equation can be reduced under the condition of ignoring the characteristic gradient. The advantage is that the electrical characteristics and magnetic characteristics can be inverted at the same time. The disadvantage is that the characteristic results have errors at different tissue interfaces. Compared with the first method, the second method assumes magnetic homogeneity because there is no electrical characteristic, and the calculated electromagnetic characteristic parameter distribution is closer to the actual situation.

[0220] The feasibility and accuracy of measuring the RF magnetic field distribution in magnetic resonance imaging makes it possible to truly evaluate the specific absorption rate (Specific Absorption Rate) related to human safety, making it possible to quantify various existing weighted imaging methods such as T1, T2, T1 / T2 and PD, and can be used to more accurately correct the magnetic field non-uniformity in images of various RF coils such as separate transmitters and receivers, integrated transmitters and receivers, phased arrays, and parallel imaging applications. BRIEF DESCRIPTION OF THE DRAWINGS

[0221] The present invention is further described below in conjunction with the accompanying drawings and embodiments.

[0222] Figure 1 It is a schematic diagram of the magnetic resonance radio frequency magnetic field measurement and electromagnetic property imaging device of the present invention.

[0223] Figure 2 The present invention is a flow chart of measuring the radio frequency transmitting magnetic field and receiving magnetic field of human body magnetic resonance imaging and inverting the parameters of tissue electrical and magnetic characteristics.

[0224] The relative precession of the magnetization intensity vector associated with the imaging scanning pulse sequence and the timing of the resonance echo signal generation affect the magnetic field measurement. Figure 3 (a) is a schematic diagram of the precession process when the lateral components of the magnetic vector rotate with equal phases but unequal amplitudes. Figure 3 (b) is a schematic diagram of the precession process when the amplitude decay rate of the transverse component of the magnetic vector is equal and the phase difference is constant. Figure 3 (c) is a timing diagram of echo signal acquisition in two scans with equal sequence excitation deflection angles and half the echo time difference.

[0225] When using a spin echo sequence to acquire magnetic resonance signals, the relative change in the phase of the magnetization intensity vector precession affects the magnetic field measurement. Figure 4 (a) is when a specific echo interval causes the transverse component of the magnetic vector to refocus on a specific coordinate axis each time. Figure 4(b) is when a specific echo time interval causes the lateral component to be refocused onto a different coordinate semi-axis each time.

[0226] Figure 5 The flowchart is a process of obtaining receiving magnetic field distribution data by magnetic resonance acquisition and reconstruction image processing.

[0227] Figure 6 The flow chart is a process of acquiring and reconstructing images through magnetic resonance imaging and processing existing receiving magnetic fields to obtain transmitting magnetic field distribution data.

[0228] Figure 7 (a) is the data processing flow chart of the first type of inversion algorithm that obtains the output electromagnetic characteristic parameter distribution from the input magnetic field data. Figure 7 (b) is the data processing flow chart of the second type of inversion algorithm.

[0229] Figure 8 It is a target conductivity distribution diagram of a water model working fluid simulation containing an abnormal body array for verifying the electromagnetic property imaging algorithm of the magnetic resonance measurement radio frequency magnetic field inversion of the present invention.

[0230] Fig. 9 It is the target relative permittivity distribution diagram of the water model working fluid simulation containing anomaly array.

[0231] Fig.10 It is the distribution diagram of the excitation emission magnetic field sensitivity field obtained by forward simulation of the water model working fluid containing the abnormal body array.

[0232] Fig.11 It is the continuous phase distribution diagram of the excitation transmission-receiving magnetic field obtained by forward simulation of the water model working fluid containing the abnormal body array.

[0233] Fig.12 It is the conductivity reconstruction distribution map obtained by inverting the water model working fluid simulation containing anomaly array based on the Helmholtz equation method.

[0234] Fig.13 It is the reconstructed distribution map of relative permittivity obtained by inverting the water model working fluid simulation containing anomaly array based on the Helmholtz equation method.

[0235] Fig.14 It is the target proton density distribution map of the human head model numerical simulation used for the verification of the electromagnetic property imaging algorithm of the radio frequency magnetic field inversion in magnetic resonance measurement.

[0236] Fig.15 It is the target conductivity distribution map of the human head model numerical simulation, used for verifying the electromagnetic property imaging algorithm of the RF magnetic field inversion by magnetic resonance measurement.

[0237] Fig.16 It is the target relative permittivity distribution map of the human head model numerical simulation, used for verifying the electromagnetic property imaging algorithm of the RF magnetic field inversion in magnetic resonance measurement.

[0238] Without adding magnetic resonance spatial encoding gradient magnetic field, Fig.17a It is similar to the distribution diagram of the excitation emission magnetic field sensitivity field amplitude obtained by numerical simulation forward modeling of the human head model of microwave tomography. Fig.17b is the corresponding phase distribution diagram.

[0239] Without adding magnetic resonance spatial encoding gradient magnetic field, Fig.18a is the distribution diagram of the detection receiving magnetic field sensitivity field amplitude obtained by the simulation forward modeling, Fig.18b is the corresponding phase distribution diagram.

[0240] Fig.19 It is the conductivity reconstruction distribution map obtained by inverting the numerical simulation of the human head model based on the Dual method containing characteristic gradient terms.

[0241] Fig. 20 The numerical simulation of the human head model is based on the Dual method with characteristic gradient terms, and the relative permittivity reconstructed distribution map is obtained by inversion.

[0242] Fig.21a is the original proton density weighted equivalent proton density amplitude distribution image, Figure 21b It is the corresponding equivalent proton density angular distribution image (in an actual imaging volume measurement and imaging example of magnetic resonance magnetic field measurement and electromagnetic property imaging of the present invention, a homogeneous water phantom with an abnormal shape is scanned under 1.5T magnetic resonance conditions).

[0243] Fig.22a is the original proton density weighted equivalent proton density amplitude distribution image, Figure 22b It is the corresponding equivalent proton density angular distribution image, in order to investigate the influence of conductivity change on electromagnetic property imaging (in an actual imaging volume measurement and imaging example of magnetic resonance magnetic field measurement and electromagnetic property imaging of the present invention, a homogeneous water model containing the same abnormal shape and 2.5g NaCl dissolved therein is scanned under 1.5T magnetic resonance conditions).

[0244] Fig.23a is the original proton density weighted equivalent proton density amplitude distribution image, Figure 23b It is the corresponding equivalent proton density angular distribution image, in order to investigate the influence of the change of conductivity and magnetic permeability on the electromagnetic property imaging (in an actual imaging volume measurement and imaging example of the magnetic resonance magnetic field measurement and electromagnetic property imaging of the present invention, a homogeneous water model containing the same abnormal shape, 2.5g NaCl dissolved and 0.25ml of 377mg / ml gadoterate meglumine was scanned under 1.5T magnetic resonance conditions).

[0245] Fig.24a is the original proton density weighted equivalent proton density amplitude distribution image, Figure 24bIt is the corresponding equivalent proton density angular distribution image (in an actual imaging volume measurement and imaging example of the magnetic resonance magnetic field measurement and electromagnetic property imaging of the present invention, human brain tissue is scanned under 1.5T magnetic resonance conditions).

[0246] Fig.25a The figure shows the measured distribution of the transmitted magnetic field sensitivity field. Fig.25b Shown is the transmitted magnetic field phase distribution, and Fig.25c The corresponding receiving magnetic field sensitivity field distribution is shown. Fig.25d Shown is the received magnetic field phase distribution (in the magnetic field measurement embodiment of scanning a homogeneous water phantom containing an abnormal shape, proton density weighted imaging is used with 3 fast spin echo sequences (120°, 40ms, 10000ms), (60°, 40ms, 10000ms) and (60°, 20ms, 10000ms)).

[0247] Fig.26a The figure shows the measured distribution of the transmitted magnetic field sensitivity field. Figure 26b Shown is the transmitted magnetic field phase distribution, and Fig.26c The corresponding receiving magnetic field sensitivity field distribution is shown. Fig.26d Shown is the received magnetic field phase distribution (in the magnetic field measurement embodiment of the scan containing the same abnormal shape and 2.5g NaCl dissolved in a homogeneous water model, proton density weighted imaging was used to measure 3 fast spin echo sequences (120°, 40ms, 10000ms), (60°, 40ms, 10000ms) and (60°, 20ms, 10000ms)).

[0248] Fig.27a The figure shows the measured distribution of the transmitted magnetic field sensitivity field. Figure 27b Shown is the transmitted magnetic field phase distribution, and Fig.27c The corresponding receiving magnetic field sensitivity field distribution is shown. Fig.27d Shown is the received magnetic field phase distribution (in the magnetic field measurement embodiment of scanning a homogeneous water phantom containing the same abnormal shape, 2.5g NaCl dissolved and 0.25ml of 377mg / ml gadoterate meglumine, using proton density weighted imaging 3 fast spin echo sequences (120°, 40ms, 10000ms), (60°, 40ms, 10000ms) and (60°, 20ms, 10000ms)).

[0249] Fig.28a The figure shows the measured distribution of the transmitted magnetic field sensitivity field. Fig.28b Shown is the transmitted magnetic field phase distribution, and Fig.28c The corresponding receiving magnetic field sensitivity field distribution is shown. Fig.28dShown is the received magnetic field phase distribution (in the magnetic field measurement embodiment of scanning human brain tissue, proton density weighted imaging is used with 3 fast spin echo sequences (120°, 40ms, 10000ms), (60°, 40ms, 10000ms) and (60°, 20ms, 10000ms)).

[0250] Fig.29a The conductivity distribution diagram in the homogeneous water model with abnormal shape obtained by inverting the transmitted magnetic field calculation quantity. Fig.29b The permittivity distribution diagram in the homogeneous water model with abnormal shape obtained by inverting the transmitted magnetic field calculation quantity. Fig.29c The conductivity distribution diagram in the homogeneous water model with abnormal shape is obtained by inverting the received magnetic field calculation. Fig.29d It is the permittivity distribution map in a homogeneous water model with an abnormal shape obtained by inverting the received magnetic field calculation.

[0251] Fig.30a It is the tissue conductivity distribution map obtained by inverting the transmitted magnetic field calculation. Fig.30b It is the tissue permittivity distribution map obtained by inverting the transmitted magnetic field calculation. Fig.30c It is the tissue conductivity distribution map obtained by inverting the received magnetic field calculation. Fig.30d It is the tissue permittivity distribution map obtained by inverting the received magnetic field calculation.

[0252] Fig.31 It is a functional schematic diagram of a typical cascade unit of a magnetic resonance imaging system receiver.

[0253] Fig.32 It is a schematic diagram of the new steady-state formation of the longitudinal magnetization intensity vector and the transverse magnetization intensity vector, wherein Figure (a) is a schematic diagram of the new steady-state formation of the longitudinal magnetization intensity, showing that the longitudinal magnetization intensity vector Mss is smaller than the initial magnetization intensity M0 of the environmental thermal equilibrium; Figure (b) is a schematic diagram of the new steady-state formation of the transverse magnetization intensity vector, showing that the inter-cyclic transverse magnetization intensity vector M'ss is greater than 0 (a new steady-state magnetization intensity different from that before scanning will be formed in the imaging body under the excitation of multiple equally spaced radio frequency pulses).

[0254] In the figure: 1. Main magnet for nuclear magnetic resonance imaging, 2. Imaging body, 3. Excitation radio frequency transmitting coil, 4. Detection induced echo signal receiving coil. DETAILED DESCRIPTION

[0255] The following will disclose multiple embodiments of the present invention with drawings. For the purpose of clear description, many practical details will be described together in the following description. However, it should be understood that these practical details should not be used to limit the present invention. That is, in some embodiments of the present invention, these practical details are not necessary. In addition, for the purpose of simplifying the drawings, some conventional structures and components will be depicted in a simple schematic manner in the drawings.

[0256] In addition, in the present invention, the descriptions of "first", "second", etc. are only used for descriptive purposes, and do not specifically refer to the order or sequence, nor are they used to limit the present invention. They are only used to distinguish components or operations described with the same technical terms, and cannot be understood as indicating or implying their relative importance or implicitly indicating the number of technical features indicated. Therefore, the features defined as "first" and "second" may explicitly or implicitly include at least one of the features. In addition, the technical solutions between the various embodiments can be combined with each other, but they must be based on the ability of ordinary technicians in the field to implement them. When the combination of technical solutions is contradictory or cannot be implemented, it should be considered that such a combination of technical solutions does not exist and is not within the scope of protection required by the present invention.

[0257] The present invention measures the excitation transmission magnetic field and detects the receiving magnetic field distribution to invert the distribution of magnetic permeability, conductivity and relative permittivity characteristic parameters, and is applicable to various magnetic resonance imaging facilities based on superconducting magnets, permanent magnets and electromagnetic magnets, etc., and is not limited by the main magnetic field strength. On the contrary, for facilities with a field strength of 3.0T or more, the transmission magnetic field and receiving magnetic field measurements of the present invention can truly and effectively evaluate the specific absorption rate (Specific Absorption Rate), coil radio frequency magnetic field uniformity, etc. The magnetic resonance transmission magnetic field and receiving magnetic field measurements are applicable to the whole body birdcage coil transmission-local coil reception radio frequency configuration, such as head and neck joint receiving coil, breast coil, etc., and are also applicable to transverse electromagnetic multi-channel integrated transceiver coil, or surface coil array configuration in parallel imaging. Within the scanning capability of human tissue magnetic resonance imaging, the magnetic field measurement and electromagnetic characteristic imaging of the present invention can examine a wide range of tissues and organs, such as the brain, pelvis and limbs. The scanning pulse sequence is based on the formation mechanism of spin echo or fast spin echo signals, and implements two-dimensional frequency-encoded layer-selected tomography or three-dimensional phase-encoded layer-selected volume imaging. In the inversion of electromagnetic properties, the former involves a two-dimensional approximation method with a zero assumption of characteristic gradient, while the latter is not constrained by thin-layer characteristic boundary conditions. Since the electromagnetic characteristic parameter values ​​vary with frequency, the inverted permeability and complex permittivity are electromagnetic characteristic values ​​only under the condition of Larmor resonance frequency corresponding to a specific field strength, such as 0.5T, 1.5T and 3.0T systems corresponding to frequency responses of approximately 21.288MHz, 63.86MHz and 127.73MHz, respectively.

[0258] Figure 1 The device shown is for measuring the magnetic resonance radio frequency transmitting magnetic field and receiving magnetic field and imaging the electrical and magnetic properties of tissues. Among them, the main magnet 1 for nuclear magnetic resonance imaging generates a uniformly distributed magnetic polarization static main magnetic field acting on the imaging body 2, and excites the radio frequency transmitting coil 3 to generate a uniformly distributed alternating magnetic field acting on the spin system in the imaging body 2. The equivalent macroscopic magnetization intensity vector is appropriately deflected to create the conditions for the generation of nuclear magnetic resonance signals. The amplitude of the transverse component formed by the deflection of the magnetization intensity vector is linearly related to the intensity of the transmitting magnetic field, and the phase of the component is 90° different from the phase of the magnetic field wave. Conventional magnetic resonance imaging is a process of combining radio frequency and gradient scanning according to the pulse sequence timing, which is a process of spatially modulating the frequency and phase of the spin precession using the magnetic field gradient, which corresponds to the frequency and phase of the k-space harmonics in the Fourier transform relationship. The receiving coil 4 collects wave number domain discrete sample point data in the form of a time domain signal containing spatial encoding information, and obtains a distribution image of the equivalent spin density precession amplitude and phase after inverse transformation. Starting from the spin precession amplitude and phase information measured by the magnetic resonance imaging, the proton density is decoupled to obtain the receiving magnetic field sensitivity field distribution related to the magnetization intensity vector of the imaging body 2, and then the phase distribution of the receiving magnetic field is obtained by using the sequence scanning processing data with the same deflection angle, different echo time and the same repetition time parameter. Then, the existing receiving magnetic field sensitivity field is used, combined with the original two magnetic resonance image amplitude data, to obtain the transmitting magnetic field sensitivity field distribution under the target condition. Finally, the transmitting magnetic field phase distribution is obtained by subtracting the existing receiving magnetic field phase from the original magnetic resonance image phase. Based on the measured transmitting magnetic field and receiving magnetic field data, the standard Helmholtz wave equation or the magnetic field-characteristic core equation containing the gradient factor is solved to invert the distribution of the magnetic permeability, conductivity and relative permittivity parameters in the imaging body 2.

[0259] The workflow of the present invention is as follows Figure 2 As shown, the following steps are included:

[0260] Step 1: The technician places the patient imaging body into the magnetic resonance facility, adjusts the examination site to align with the geometric center of the gradient coil according to the examination and scanning requirements; records the configuration of the RF excitation and detection coil parameters and types used;

[0261] Step 2: Perform three fast spin echo sequence imaging scans, set different specific flip angles, echo times, repetition times and other sequence parameters, such as (90°, TE, TR), (45°, TE, TR) and (45°, TE / 2, TR); change the polarity of the layer selection, phase encoding and frequency encoding gradients, and repeat the three scans; keep the geometric parameters such as field of view, voxel, layer thickness, number of layers and the like consistent in the six scans;

[0262] Step 3: Save the scanned image data;

[0263] Step 4: Process the image data, compare and verify that the ratio of the maximum amplitude of the sequences with different flip angles is consistent with theoretical expectations, and implement phase unwrapping measures;

[0264] Step 5: Replace the homogeneous water phantom and repeat steps 1-4 to perform imaging at the same facility gradient center position, and keep the field of view, voxel, layer, layer thickness and other sequence parameters consistent with the parameters of the patient imaging volume imaging above;

[0265] Step 6: using the patient amplitude acquisition data scanned at different flip angles, echo times and repeated time series, combined with the corresponding data collected by the homogeneous water model, to process and obtain the amplitude distribution of the receiving magnetic field sensitivity field of the patient imaging body; using the patient phase acquisition data scanned at the flip angle, proportional echo time and repeated time series, combined with the corresponding data collected by the homogeneous water model, to process and obtain the receiving magnetic field phase distribution of the patient imaging body;

[0266] Step 7: Still based on the patient amplitude acquisition data scanned by the different flip angles, echo times and repeated time series, combined with the obtained receiving magnetic field sensitivity field amplitude, processing is performed to obtain the transmitting magnetic field sensitivity field amplitude of the patient imaging body; still using the patient phase acquisition data scanned by the flip angle, echo time and repeated time series, combined with the obtained receiving magnetic field phase, processing is performed to obtain the transmitting magnetic field phase distribution of the patient imaging body;

[0267] Step 8: Using the first type of inversion method for electromagnetic property imaging of the present invention, first use the existing magnetic resonance electrical property imaging to invert the transmitted magnetic field data to obtain the electrical properties, and then invert the magnetic properties based on the convection-reaction equation or the standard Helmholtz equation; or, using the second type of inversion method of the present invention, from the transmitted magnetic field and the received magnetic field data, based on the Helmholtz equation that ignores the characteristic gradient, synchronously invert the distribution of electrical and magnetic property parameters.

[0268] According to the law of the precession process of the macroscopic magnetization intensity vector during the imaging scanning process, the present invention combines the general characteristics of the wave vector amplitude and phase change to design a method for measuring the distribution of the RF transmitting magnetic field and the receiving magnetic field. For the scanning target imaging volume 2, the lumped longitudinal relaxation time and transverse relaxation time of each voxel are given, which means that the rate of recovery of the longitudinal magnetic vector component and attenuation of the transverse magnetic vector component in space is given. Figure 3 In Figure (a), the equivalent spin angular momentum A1 and A2 in two scans at a certain voxel position are examined. Under different conditions of the excitation emission magnetic field amplitude, different initial magnetic vector flip angles are generated, and then the magnitude of the transverse magnetic vector component is different, which corresponds to A1 being longer than A2. If the signal acquisition echo time and magnetic vector restoration repetition time of the two scanning sequences are consistent, according to the phase angle Then, in each echo formation, A1 and A2 always keep the phase equal, that is, the magnetic vectors are in the same direction, and the amplitude ratio of A1 and A2 depends on the size of the initial transverse vector, because the longitudinal relaxation time and transverse relaxation time of the same voxel are constant. The acquisition signal of magnetic resonance imaging is essentially a magnetic field radio frequency wave in the form of amplitude-phase, and the spatial decoding of the precession frequency and phase by the static magnetic field gradient defines the voxel array. Therefore, the present invention scans the imaging body with a sequence of two different deflection angles, consistent echo time and repetition time, which provides the possibility of decoupling the proton density from the detection signal to measure the sensitive field amplitude of the receiving magnetic field.

[0269] For the voxel-based spin precession amplitude and phase information obtained by Fourier inverse transform of conventional magnetic resonance imaging, the effect of the radio frequency pulse excitation transmitting magnetic field on the initial steady-state magnetic intensity vector is not only to deflect the absolute longitudinal magnetic vector in direction to produce the transverse component of the magnetic vector, but also the initial phase of the transverse component is related to the initial phase of the transmitting magnetic field. Specifically, due to the nutation effect, the initial phase of the transverse component is equal to the initial phase of the transmitting magnetic field plus π / 2. The phase distribution reconstructed by magnetic resonance imaging means the sum of the continuous receiving magnetic field phase and the transmitting magnetic field. Since the transmission and reception propagation directions are opposite, the transmitting magnetic field phase is negative when the sum is taken. Under certain conditions, such as the classic orthogonal dual-channel birdcage coil as the common coil for excitation and detection, the initial phase of the transmitting magnetic field is approximately equal to half of the phase of the detection signal because the polarity of the port is changed to achieve the transmission and reception conversion. In the process of forming the collected echo signal, the relative phase dispersion of the precession between the voxel arrays just represents the difference in the amplitude phase of the transverse magnetic vector, and is reconstructed and displayed by Fourier inverse transform. Similar to the aforementioned magnetic field sensitivity field amplitude measurement, the present invention needs to execute two scanning sequences with special parameters to realize the measurement of magnetic field phase. The two sequences keep the deflection angle consistent, the echo time ratio constant and the repetition time consistent, for example (45°, TE, TR) and (45°, TE / 2, TR). Figure 3 Figure (b) shows the corresponding running amplitude and phase during two scanning echoes at the same voxel position under the same deflection angle. Because the deflection angle is equal and the echo time is half the difference, although the amplitudes of A3 and A4 are not strictly equal, they decrease according to the same exponential decay curve, and there is a constant phase difference between the relative phase dispersion of the two precessions. Figure 3 Figure (c) shows the spin echo signals generated twice by the same voxel. When the longitudinal and transverse relaxation rates are consistent, the conditions for measuring the phase of the magnetic resonance radio frequency receiving magnetic field are created, and the amplitude of the precession of the transverse component of the magnetic vector is correlated and the phase is coherent. Because the phase constant of the transmitting magnetic field is transferred to the initial transverse component of the magnetic vector during excitation, the subtraction of the two original magnetic resonance phase distribution data can obtain a single receiving magnetic field phase related to the echo time ratio. The phase distribution data of the original magnetic resonance reconstruction image, that is, the continuous phase measurement of the transmitting-receiving magnetic field, is still used to further calculate the phase distribution of the transmitting magnetic field from the receiving magnetic field phase.

[0270] When scanning with a fast spin echo sequence with echo time parameter TE, such as (45°, TE, TR), where the interval between adjacent echoes TS is determined by TE and the echo chain length ETL, N data are collected in each echo and the sampling data interval is td, then the echo reception time is N*td, assuming that the received signal noise ratio at this time is rSNR. When scanning with a sequence with echo time parameter TE / 2, such as (45°, TE / 2, TR), when ETL remains unchanged, the interval between adjacent echoes is TS / 2. If the number of echo acquisition data N remains unchanged, the sampling data interval of each echo is correspondingly set to td / 2, and the echo reception time is N*td / 2. Because the echo time is short, the received signal noise ratio at this time will be reduced to about 0.707*rSNR. In order to keep the signal-to-noise ratio of the two scans consistent, that is, the two transverse components of the magnetic vector decay from the same amplitude and along the same curve under the same excitation deflection angle pulse, the present invention adjusts td in disguise by adjusting the bandwidth BW of the sequence (45°, TE / 2, TR). Since tacq is related to the image resolution, the magnetic resonance facility will automatically adjust the magnitude of the magnetic field gradient G according to the set echo sampling point number N and tacq without changing the imaging field of view and resolution.

[0271] The physical process of imaging can be regarded as the reciprocating cyclic motion of the magnetic vector in the rotating coordinate system. Figure 3 Figure (c) analyzes from the time dimension the process of using a scanning sequence with the same flip angle and different echo times to form a signal on the imaging body to measure the phase of the radio frequency magnetic field, that is, the precession phase of the transverse component of the magnetic vector. Figure 4 From the spatial dimension, we continue to analyze the process of the precession of the transverse component of the magnetic vector to illustrate the principle of measuring the magnetic field phase of the present invention. Figure 4 Figure (a) shows that when the refocusing pulse acts, the average forward magnetic vector is located on the coordinate axis, and the refocusing peaks of each magnetic vector system are located on the same coordinate semi-axis. Figure 4Figure (b) shows the situation when the average speed magnetic vector is not located on the coordinate axis during the refocusing pulse, but the refocusing peak of each magnetic vector system is located on the coordinate axis. In the rotating coordinate system x'y'z, a fast spin echo sequence with an echo time of TE is used, such as (45°, TE, TR). The scanning imaging body (2) is stimulated by a pulse at time t=0 to generate a partial transverse component of the magnetic vector. After TS / 2 time, the maximum relative phase difference π / 2 is formed between the transverse components. It is assumed that the medium speed magnetic vector is along the +y' axis at this time, and the first refocusing pulse is implemented. At TS, the first echo signal peak is formed. The selection of equal TS allows the echo signal peak to be formed at the same time and in the same direction, and the maximum degree of phase dispersion in each round is also consistent, all with a phase difference of π / 2. When the fast spin echo sequence with the same deflection angle excitation pulse and echo time of TE / 2, such as (45°, TE / 2, TR), the scanning imaging body 2 still generates the same magnitude of the partial transverse component of the magnetic vector at t=0, and the maximum phase difference π / 4 is formed between the time components after TS / 4. Assuming that the magnetic vector is at an angle of 45° with the +x' axis, the first refocusing pulse is implemented. When TS / 2 is reached, the first signal peak is formed, and the transverse component of the magnetic vector points to the negative direction of the -y' axis. After each subsequent TS / 2, an echo peak is formed on the positive and negative half axes of x'y'. Compared with the echo of the (45°, TE, TR) sequence, the (45°, TE / 2, TR) sequence has a larger echo amplitude and a higher signal-to-noise ratio because it has not undergone a long period of transverse relaxation. The present invention measures the radio frequency magnetic field phase, i.e., the precession phase of the transverse component of the magnetic vector, by controlling the constant difference between the maximum phase difference formed by the precession phase dispersion according to the length of the echo interval of different sequences, so as to achieve the measurement of the receiving phase under the same excitation phase, and conversely calculate the relatively uniform transmission phase by combining the existing original magnetic resonance phase and the receiving phase. In a fast spin echo sequence, if one echo time is set, the parameter TE specifically refers to the time corresponding to the echo located in the middle position in the echo signal chain.

[0272] On the basis of conventional magnetic resonance imaging examination, the present invention needs to save and record the angular image data in addition to the amplitude image, and record the excitation magnetic field amplitude B1+RMS of the imaging device during each scan to compare with the magnetic field measurement value. Specifically, one embodiment is to use an 8-channel phased array parallel imaging coil to perform RF magnetic field measurement and electromagnetic characteristic inversion for cranial brain and breast scanning, and set the measurement geometric parameters in combination with the magnetic field gradient to have a field of view FOV 256*256mm 2 , frequency encoding and phase encoding steps are 256*256, layer thickness is 2mm, layer spacing is 2mm, and number of layers is 12.

[0273] Preferably, in order to make the histogram distribution of the image data of the cosine value of the excitation deflection angle fall into the approximate normal distribution range centered at 0.5 as much as possible, rather than cos45°=0.7071, when calculating the transmit magnetic field sensitivity field data, the deflection angles of sequences I and II are set to 120° and 60° respectively. Different from the existing dual-angle method of B1-mapping technology, the corresponding refocusing pulse deflection angles of the fast spin echo sequences I and II in the present invention are both 180°.

[0274] In order to eliminate the eddy current effect caused by the magnetic field gradient pulse, each magnetic field phase measurement magnetic resonance scanning sequence needs to be scanned with the corresponding reverse polarity sequence. The normal polarity argument data and the reverse polarity argument data obtained are averaged by pixel as the original magnetic resonance scanning argument data, which is used to calculate the phase distribution of the receiving magnetic field and the transmitting magnetic field.

[0275] In order to measure the distribution of electrical and magnetic properties of the reconstructed brain, a fluid suppression scanning sequence (Weighted Fluid-Attenuated Inversion Recovery Imaging, FLAIR) of human tissue and water phantom is successively implemented according to the following embodiment. The sequence excitation and detection parameters of the magnetic resonance imaging raw image data acquisition are shown in Table 3.

[0276] Table 3

[0277] Sequence T2-FLAIR (α, TE, TR) Amplitude data Argument data Reverse polarity angle data Brain I(120°,110ms,9000ms) m_2α_h - m_2α_h_po II (60°, 110ms, 9000ms) m_α_h a_α_h a_α_h_po III(60°,55ms,9000ms) - a_t_h a_t_h_po Water Model IV(120°,110ms,9000ms) m_2α_p - m_2α_p_po V(60°,110ms,9000ms) m_α_p a_α_p a_α_p_po VI (60°, 55ms, 9000ms) - a_t_p a_t_p_po

[0278] exist Figure 5 In the embodiment of the magnetic resonance receiving magnetic field amplitude sensitive field and phase distribution measurement shown in the figure, a radio frequency based on the principle of amplitude ratio and phase difference is given. magnetic field Measurement method. First, scan the cranial brain tissue with the main sequence II whose parameters of deflection angle, echo time and repetition time are (60°, 110ms, 9000ms) respectively, and obtain the amplitude and phase angle data m_α_h and a_α_h, where m represents the amplitude, a represents the angle, α represents the deflection angle, and h represents the scanning part of the human head. Then scan with the auxiliary sequence I parameters whose parameters of deflection angle, echo time and repetition time are (120°, 110ms, 9000ms) respectively, and obtain the amplitude data m_2α_h. Based on the pixel level, divide the obtained amplitude data m_2α_h by the data m_α_h to obtain the amplitude distribution data m of the cranial brain receiving magnetic field sensitive field weighted by the sine ratio of the proton density decoupled transmitting magnetic field. h (r), where m h (r) contains the contribution of the receiving coil magnetic field inhomogeneity and is related to the selected TE and TR. In order to avoid nonlinearity, it is also necessary to measure the homogeneous water phantom magnetic field data with the same parameter sequence and assume that the magnetic permeability of brain tissue is the magnetic permeability of water. After further processing, the influence of the transmitted magnetic field is eliminated.

[0279] Continue to scan the homogeneous water model with the sequence V and IV with the same parameter group, and obtain the amplitude and argument data m_α_p and a_α_p, and the amplitude m_2α_p, where p represents the water model. Divide the obtained amplitude m_2α_p by m_α_p to obtain the water model receiving magnetic field sensitive field amplitude distribution data m weighted by the sine ratio of the proton density decoupled transmitting magnetic field p (r). Since the water model is homogeneous, the m p (r) only contains the fluctuation of the magnetic field distribution of the receiving coil, that is, the contribution of non-uniformity.

[0280] The received magnetic field amplitude m with transmission weighting obtained by the brain scan measurement is h (r) divided by the receiving magnetic field m of the water model p (r), under unit current conditions, the amplitude distribution of the brain receiving magnetic field sensitivity field is obtained by eliminating the amplitude ratio of the sine value of the transmitting magnetic field The superscript TE indicates a specific echo time condition.

[0281] Combination Figure 3 The spatial precession of nuclear spin angular momentum, i.e. the rotation law of the macroscopic transverse component of the magnetic vector, is shown in the figure. In order to measure the phase distribution of the receiving magnetic field of the cranial tissue based on magnetic resonance imaging scanning, a phase scanning auxiliary sequence III (60°, 55ms, 9000ms) is designed. Compared with the main sequence II, the auxiliary sequence III has the same deflection angle, the echo time is shortened by 1 / 2, and the repetition time is the same, which means that the amplitude of the transverse component of the magnetic vector corresponding to the sequential echoes of the same voxel between two scans is consistent, and the phase difference is constant. The angular data obtained by the sequence III scan are a_t_h and a_t_h_po.

[0282] Subtract the argument data a_α_h' and a_t_h' obtained from the two cranial brain scans after gradient polarity correction of eddy current interference to obtain the receiving magnetic field phase Δp without the transmitting magnetic field phase h (r) is related to the selected echo time TE. Due to the inhomogeneity of the actual receiving coil magnetic field, the brain phase distribution Δp h In (r), in addition to the contribution of the electromagnetic properties of the tissue, it also contains the contribution of the receiving coil field fluctuations.

[0283] Similarly, the water phantom was scanned with sequence V with consistent imaging parameters (60°, 110ms, 9000ms) to obtain the argument data a_α_p and a_α_p_po, and the auxiliary sequence VI with parameters (60°, 55ms, 9000ms) was scanned to obtain the data a_t_p and a_t_p_po. After correction for the eddy current effect, the phase difference Δp between the sequences was obtained by subtracting a_t_p' from a_α_p'. p (r). Data Δpp (r) distribution contains the data Δp h The same receive coil intrinsic phase field inhomogeneity in (r).

[0284] Further, the phase difference data Δp of the cranial brain intermediate process h (r) minus the water mode phase difference Δp p (r) is used to obtain the absolute phase of the received magnetic field during brain imaging signal acquisition. Since the echo time and repetition time of the two sequences are (110ms, 9000ms) and (55ms, 9000ms) respectively, in order to obtain the unit received magnetic field phase The obtained single phase quantity needs to be multiplied by 2. Compared with the contribution of the transmitted magnetic field amplitude to the equivalent spin density signal in the form of a functional, the transmitted magnetic field and the received magnetic field phase have a linear relationship in the signal, making it relatively simple to calculate the magnetic field phase based on the designed sequence acquisition data.

[0285] The sensitive field amplitude and phase distribution data of the received magnetic field under the unit current of the brain have been calculated. The general processing method of the periodic magnetic field wave in the space domain is still based on the amplitude ratio and phase difference. Combined with the measurement scanning data of the above sequence, the amplitude and phase of the transmitted magnetic field under specific parameters are obtained through the following data calculation. The data processing process is as follows: Figure 6 shown.

[0286] In order to obtain the transmit magnetic field amplitude of the brain scan main sequence II (60°, 110ms, 9000ms), first divide the original magnetic resonance image amplitude data m_α_h by the unit receive magnetic field amplitude The remaining proton density amplitude weighted transmit magnetic field sensitive field amplitude sinusoidal quantity data t2. Similarly, under the same voxel position conditions of the same imaging volume, the image amplitude m_2α_h of the auxiliary sequence I with 2 times the deflection angle, the same echo and repetition time is divided by The remaining emitted magnetic field sinusoidal quantity data t1 with the same proton density weight.

[0287] Considering that the transmitted magnetic field amplitude is nested inside the sine trigonometric function As well as the contribution of consistent proton density distribution, using the sine double angle identity, further divide the intermediate process quantity t1 by t2 to calculate the amplitude of the emitted magnetic field related to the brain's own electromagnetic characteristics and the non-uniform magnetic field of the coil That is, the magnetic field under the excitation of the RF pulse with a deflection angle of 60°, where τ is the pulse width. Note that this is different from The absence of a top mark '~' reflects the principle that magnetic resonance imaging requires the radio frequency excitation magnetic field amplitude to be uniform.

[0288] In the above, the phase distribution of the magnetic field received by the cranial brain scan has been measured and calculated. Subtract the original MRI image angle data a_α_h' corrected by eddy current from the main sequence II. Since the excitation magnetic field phase e index in the magnetization intensity expression is multiplied by 1 imaginary unit, that is, the transverse component of the magnetic vector is delayed, the difference is subtracted by π / 2 to obtain the transmission magnetic field phase distribution data Where α represents the condition of a deflection angle of 60°. Similarly to the amplitude of the transmitting magnetic field, according to the principle of magnetic resonance imaging, the phase distribution of the excitation magnetic field intensity under no-load conditions also tends to be uniformly distributed, approaching the ideal uniform distribution condition of the deflection angle.

[0289] In order to measure the distribution of electrical and magnetic properties of reconstructed imaging breast tissue, a transverse relaxation time weighted scanning sequence (T2-Weighted Imaging, T2WI) of human tissue and phantom was successively implemented using a multi-channel surface coil array according to the following embodiment. Considering the characteristics of breast tissue related metabolites and fat-based, the repetition time TR=4000ms was set to save scanning time. The sequence excitation and detection parameters for magnetic resonance imaging raw image data acquisition are shown in Table 4.

[0290] Table 4

[0291] Sequence T2WI (α, TE, TR) Amplitude data Argument data Reverse polarity angle data Breast I(120°,90ms,4000ms) m_2α_b - - II (60°, 90ms, 4000ms) m_α_b a_α_b a_α_b_po III(60°,45ms,4000ms) - a_t_b a_t_b_po Phantom IV (120°, 90ms, 4000ms) m_2α_p - - V(60°,90ms,4000ms) m_α_p a_α_p a_α_p_po VI (60°, 45ms, 4000ms) - a_t_p a_t_p_po

[0292] exist Figure 5 In the embodiment of the measurement of the amplitude sensitive field and phase distribution of the magnetic resonance receiving magnetic field shown, a radio frequency magnetic field measurement method based on the principle of amplitude ratio and phase difference is given. First, the breast tissue is scanned with the main sequence II whose deflection angle, echo time and repetition time parameters are (60°, 90ms, 4000ms) respectively to obtain the amplitude and phase angle data m_α_b and a_α_b, where b represents the scanning part of the human breast. Then, the auxiliary sequence I parameter scan is performed with the deflection angle, echo time and repetition time parameters being (120°, 90ms, 4000ms) respectively to obtain the amplitude data m_2α_b. Based on the pixel level, the obtained amplitude data m_2α_b is divided by the data m_α_b to obtain the breast receiving magnetic field sensitive field amplitude distribution data m_2α_b weighted by the sine ratio of the proton density decoupled transmitting magnetic field. b (r), where m b (r) contains the contribution of the receiving coil magnetic field inhomogeneity and is related to the selected TE and TR. In order to avoid nonlinearity, it is also necessary to measure the magnetic field data of the homogeneous phantom with the same parameter sequence, and assume that the magnetic permeability of breast tissue is the magnetic permeability of the phantom. After further processing, the influence of the transmitted magnetic field is eliminated.

[0293] Continue to scan the homogeneous phantom with the same parameter set sequence V and IV to obtain the amplitude and angle data m_α_p and a_α_p, as well as the amplitude m_2α_p. Divide the obtained amplitude m_2α_p by m_α_p to obtain the amplitude distribution data m of the phantom receiving magnetic field sensitive field weighted by the sine ratio of the proton density decoupled transmitting magnetic field p (r). Since the phantom is homogeneous, the m p (r) only contains the fluctuation of the magnetic field distribution of the receiving coil, that is, the contribution of non-uniformity.

[0294] The receiving magnetic field amplitude m with transmission weighting obtained by the breast scanning measurement is b (r) divided by the received magnetic field m of the phantom p (r), under unit current conditions, the amplitude distribution of the mammary gland receiving magnetic field sensitivity field is obtained by eliminating the amplitude ratio of the sine value of the transmitting magnetic field

[0295] Combination Figure 3 The spatial precession of nuclear spin angular momentum, i.e. the rotation law of the transverse component of the macroscopic magnetic vector, is shown in the figure. In order to measure the phase distribution of the receiving magnetic field of breast tissue based on magnetic resonance imaging scanning, a phase scanning auxiliary sequence III (60°, 45ms, 4000ms) is designed. Compared with the main sequence I, the auxiliary sequence III has the same deflection angle, the echo time is shortened by 1 / 2, and the repetition time is the same, which means that the amplitude of the transverse component of the magnetic vector corresponding to the sequential echoes of the same voxel between two scans is consistent, and the phase difference is constant. The gradient polarity normal and opposite argument data obtained by sequence III scanning are a_t_b and a_t_b_po respectively.

[0296] Subtract the argument data a_α_b' and a_t_b' obtained from the two breast scans after gradient polarity correction of eddy current interference to obtain the receiving magnetic field phase Δp without the transmitting magnetic field phase b (r) is related to the selected echo time TE. Due to the inhomogeneity of the actual receiving coil magnetic field, the breast phase distribution Δp b In (r), in addition to the contribution of the electromagnetic properties of the tissue, it also contains the contribution of the receiving coil field fluctuations.

[0297] Similarly, the phantom was scanned with the same imaging parameters (60°, 90ms, 4000ms) sequence V to obtain the angle data a_α_p and a_α_p_po, and the auxiliary sequence VI with parameters (60°, 45ms, 4000ms) was scanned to obtain the data a_t_p and a_t_p_po. After correction for the eddy current effect, the phase difference Δp between the sequences was obtained by subtracting a_t_p' from a_α_p'. p (r). Data Δp p (r) distribution contains the data Δp bThe same receive coil intrinsic phase field inhomogeneity in (r).

[0298] Further, the intermediate phase difference data Δp b (r) minus the phantom phase difference Δp p (r) is used to obtain the absolute phase of the received magnetic field during breast imaging signal acquisition. Since the echo time and repetition time of the two sequences are (90ms, 4000ms) and (45ms, 4000ms) respectively, in order to obtain the unit received magnetic field phase The obtained single phase quantity needs to be multiplied by 2.

[0299] The sensitive field amplitude and phase distribution data of the received magnetic field under the unit current of the breast have been calculated. The general processing method of the periodic magnetic field wave in the spatial domain is still based on the amplitude ratio and phase difference. Combined with the measurement scanning data of the above sequence, the amplitude and phase of the transmitted magnetic field under specific parameters are obtained through the following data calculation. The data processing process is as follows: Figure 6 shown.

[0300] In order to obtain the transmit magnetic field amplitude of the breast scan main sequence II (60°, 90ms, 4000ms), first divide the original magnetic resonance image amplitude data m_α_b by the unit receive magnetic field amplitude The remaining proton density amplitude weighted transmit magnetic field sensitive field amplitude sinusoidal quantity data t2. Similarly, under the same voxel position conditions of the same imaging volume, the image amplitude m_2α_b of auxiliary sequence I with 2 times the deflection angle, the same echo and repetition time is divided by The remaining emitted magnetic field sinusoidal quantity data t1 with the same proton density weight.

[0301] Considering that the transmitted magnetic field amplitude is nested inside the sine trigonometric function As well as the contribution of consistent proton density distribution, using the sine double angle identity, further divide the intermediate process quantity t1 by t2 to calculate the emission magnetic field amplitude related to the electromagnetic characteristics of the breast itself and the non-uniform magnetic field of the coil That is, the magnetic field under the excitation of the RF pulse with a deflection angle of 60°, where τ is the pulse width.

[0302] In the above, the phase distribution of the receiving magnetic field of breast scanning has been measured and calculated. The original MRI image angle data a_α_b' corrected by eddy current of main sequence II is subtracted from it. Since the excitation magnetic field phase e index in the magnetization intensity expression is multiplied by 1 imaginary unit, that is, the transverse component of the magnetic vector is delayed, the difference is subtracted by π / 2 to obtain the transmission magnetic field phase distribution data.

[0303] It is assumed that the electromagnetic properties of the MRI volume are isotropic, that is, when the circularly polarized right-handed transmitting magnetic field and the left-handed receiving magnetic field interact with the material, the amplitudes of the scalar permeability and complex permittivity inverted from the measured quantities under different magnetic field gradient encoding and decoding directions are consistent and do not appear as tensors. On the other hand, it is assumed that the electromagnetic properties of the imaging volume have linear properties, that is, for the same voxel, when the transmitting magnetic field and the receiving magnetic field of different amplitudes interact with the object, the internal magnetic flux density generated is in the same positive proportion to the amplitude of the external applied field, and there is no time lag effect in the induced magnetic field with the fluctuation of the excitation magnetic field and the internal magnetization intensity, that is, the magnetic permeability is a real number.

[0304] According to an embodiment of the present invention, the electromagnetic properties of human tissue are inverted from the measured magnetic field, that is, the magnetic permeability parameter distribution is calculated by the first type of property inversion method. Figure 7 Figure (a) shows the transmitted magnetic field measured during the process of scanning the imaging body using magnetic resonance. Under the assumption that the magnetic permeability in the body is uniformly distributed and equal to the magnetic permeability of vacuum, the electrical characteristic distribution of human tissue is first inverted by the convection-reaction equation method containing the permittivity gradient. Then, the receiving magnetic field data measured on the same imaging body is combined with the obtained electrical characteristics, and the magnetic characteristic distribution, i.e., the magnetic permeability distribution, is inverted based on the convection-reaction equation containing the characteristic gradient to improve the accuracy of the result.

[0305] The core equation form for inversion is derived below, and the assumptions used are statistically analyzed. The electromagnetic wave equation in time-harmonic vector form is:

[0306]

[0307] Here, both the electrical characteristic gradient term and the field differential are included. Combined with the boundary conditions, the electromagnetic characteristic distribution can be solved. Currently, no assumptions or approximations are made except that the field is in the form of a simple harmonic wave.

[0308] For the first type of characteristic inversion method, we first assume that the magnetic characteristics are homogeneous to solve the electrical characteristics, that is, ①: the magnetic homogeneity in the imaging body has Gauss's magnetic law for the magnetic field intensity, that is, μ = μ0, And replace H with the measured quantity B, we have:

[0309]

[0310] Referring to the convection-reaction partial differential equation form in fluid mechanics, the core inversion equations of magnetic field differential and complex permittivity are further derived from the above equations, and a sparse linear equation system based on the coefficient matrix of the voxel array is constructed in combination with the single-shot magnetic field measurement data. With the necessary boundary condition settings, the distribution of electrical characteristic parameters can be reconstructed using the least squares method with regularization.

[0311] After obtaining the inversion results of the electrical characteristics, another set of independent magnetic field measurements, such as the received magnetic field data in the same excitation and detection, is combined to further invert the distribution of magnetic characteristic parameters. Still starting from equation (57) without assumptions, assumption ②: The imaging volume is magnetically homogeneous, then according to Gauss's magnetic law Eliminate the second term on the left side of the equation, and write the other terms on both sides as measured quantities about the magnetic flux density, and extract the homogeneous magnetic permeability from the differential operator, we have:

[0312]

[0313] In the formula, to distinguish it from the second type of characteristic inversion method, is the electrical characteristic distribution used to calculate magnetic characteristics obtained by inversion of existing electrical characteristic imaging in the first method; the magnetic flux density of biological imaging body B=μ b H, μ b is the magnetic permeability. In this way, the complexity of the solution process is reduced, and the constitutive equation for the unknown magnetic permeability is explicitly derived. Continue with counter assumption ③: The magnetic permeability μ b Inhomogeneous, that is, it changes with the change of tissue. The core equation is about the known quantity and B and the unknown quantity μ b The linear equation of magnetic properties can be solved by the least square method.

[0314] In summary, in the above electromagnetic characteristic inversion calculation, the electric characteristic parameters are solved using the assumption of magnetic homogeneity once, and the magnetic characteristic parameters are solved using the assumptions of magnetic homogeneity and magnetic inhomogeneity once. So far, the known magnetic field measurement sensitive field amplitude and phase data have been used to complete the inversion of the distribution of the conductivity and relative permittivity and magnetic permeability parameters of the imaging body, that is, the relevant image reconstruction has been achieved.

[0315] According to an embodiment of the present invention, the electromagnetic characteristics of human tissue are inverted from the measured magnetic field, that is, the magnetic permeability parameter distribution is calculated by the second type of characteristic inversion method. Figure 7 Figure (b) shows the process of synchronously inverting the distribution of electric and magnetic properties from the measured and processed in-vivo transmitted and received magnetic field data. In order to reduce the complexity of the core equations for electromagnetic property inversion, the distribution of the gradients of electric and magnetic properties is assumed to be zero, that is, the imaging volume is considered to be homogeneous, thereby ignoring the boundaries between different tissues in the body, and deriving the approximate Helmholtz wave equation. The distribution of magnetic permeability, electrical conductivity and permittivity parameters is synchronously inverted and calculated. Among them, electrical conductivity and permittivity constitute complex permittivity, which are the imaginary part and the real part respectively; after calculating the complex magnetic permeability of the voxel, the real part is taken to obtain the real magnetic permeability.

[0316] The following is a specific derivation example of the core equation for inverting electromagnetic properties from the measured magnetic field, and analyzes the inversion error, that is, the use of statistical assumptions. Starting from the Faraday law and Ampere's law of the Maxwell equations describing the electromagnetic field behavior in matter, under the time-harmonic propagation conditions of nuclear magnetic resonance, formula (13) is rewritten as follows:

[0317]

[0318] Depend on have Substituting into the first equation of formula (60), and B = μ b H, has:

[0319]

[0320]

[0321]

[0322]

[0323] So far, the known quantity measured is B, and the unknown quantity to be determined is the inhomogeneous magnetic permeability μ of human tissue. b and complex permittivity The subscript 'b' denotes human body. No other assumptions are made except the harmonic precondition of the magnetic resonance RF magnetic field.

[0324] In order to reduce the complexity of the equation, assumption ①: The magnetic properties of the imaging body are homogeneous, that is, the magnetic permeability is constant. According to Gauss's magnetic law, Assumption ②: The electrical properties in the body are homogeneous, that is, the complex permittivity is a constant, so As a price, the second type of inversion method of the present invention sacrifices the accuracy of the solution of the equation, especially systematic calculation errors will occur at the tissue boundary. Let v = 1 / μ b , then:

[0325]

[0326] Here, we make the counter assumption ③: imaging the magnetic inhomogeneity in the body, that is, Combining the relevant vector identities, we can write equation (62) in component form:

[0327]

[0328] So far, the equations for the nonlinear relationship between electrical and magnetic properties have been derived. A single scan of the same imaging body can measure a set of independent transmission magnetic field and reception magnetic field distribution data, that is, the complex value known quantity and The conditions necessary for solving a system of equations containing two unknowns and two equations are basically met. By using two or more sets of transmitting and receiving magnetic field measurements, such as changing sequence parameters for additional excitation and detection, the real permeability and complex permittivity can be appropriately solved according to the complex constitutive equation set (63).

[0329] In summary, the above second type of inversion contains one complex permittivity homogeneous assumption and one magnetic permeability homogeneous assumption, as well as one magnetic permeability inhomogeneous counter-assumption.

[0330] In order to illustrate the feasibility and effectiveness of the magnetic resonance magnetic field measurement and electrical and magnetic properties imaging method of the present invention, a numerical model is designed based on the two characteristics to simulate the radio frequency electromagnetic field distribution, and the inversion results of the conductivity and permittivity distribution are obtained and the display image is reconstructed, as well as the original data images of the density, proton density, target electromagnetic properties, etc. of the attached model itself.

[0331] In a simulation embodiment of a homogeneous water phantom model containing an array of abnormal bodies of the present invention, small cylindrical areas of different diameters are created, giving different conductivity and relative permittivity values. Under the main magnetic field conditions of 1.5T magnetic resonance imaging, i.e., 63.864MHz resonance frequency, an orthogonal birdcage type radio frequency coil is used to obtain the excitation emission magnetic field sensitive field amplitude and the emission-receiving continuous phase data in the detection signal. The cube model is 200mm long, 200mm wide, and 11mm high. The internal main effective area is a large cylinder. The data lattice size is [200, 200, 11]. The original target conductivity distribution range is [0.1, 2] S / m, with a background value of 0.5S / m; the original target relative permittivity distribution range is [50, 100], with a background value of 80, which is equal to the relative permittivity of water.

[0332] The distribution of target conductivity in the fault simulated by water containing a circular anomaly array is as follows: Figure 8 As shown, the corresponding target relative permittivity distribution is as follows Fig. 9 As shown in the figure, the smallest characteristic difference that can be reflected is 0.1S / m and 1 relative unit, and 1 voxel can represent the smallest square area, which is 1mm*1mm for a fault, i.e., one pixel.

[0333] The amplitude distribution of the transmitting magnetic field sensitivity field obtained by forward simulation is reconstructed as follows: Fig.10 As shown, the continuous phase distribution of the received magnetic field and the transmitted magnetic field is reconstructed as Fig.11 As shown. Among them, the amplitude range is (1,1.1)*10 -7 T, phase change range (-1.2,-0.4) rad.

[0334] Based on the forward modeled transmit sensitive field amplitude and receive-transmit continuous phase data, the conductivity distribution is inverted using the standard Helmholtz equation method without characteristic gradients as follows: Fig.12 As shown, the permittivity distribution is Fig.13 As shown. Among them, due to the significant dispersion of the actual inversion result numerical range, the reference model original target conductivity and permittivity values ​​are used to limit the display range to [0,2.5]S / m and [40,100]. It can be seen that although the standard Helmholtz equation method is based on the assumption that the electromagnetic properties of the imaging body are homogeneous, the algorithm simulation results can still clearly distinguish the boundaries between the background and the imaging body, and the preset circular array area and its boundaries also have high sensitivity and specificity. Note that the amplitude fluctuation at the edge of a single circular area is caused by the homogeneous assumption. Further use of low-pass filtering with a standard deviation per unit pixel can improve the influence of this spatial harmonic effect.

[0335] Another simulation embodiment of the present invention is to use the forward modeling radio frequency magnetic field of a human head model for the verification of the electrical property inversion algorithm. Based on 7T magnetic resonance imaging conditions, a 16-channel transverse electromagnetic transmitting and receiving shared radio frequency coil is constructed, and the magnetic field distribution is calculated in a numerical head model derived from real tissue characteristic parameters. The transmission-receiving dual source Dual method is used to invert the distribution of conductivity and relative permittivity parameters, and a series of image results are reconstructed. The model contains 41 different tissue types.

[0336] The proton density distribution related to nuclear magnetic resonance set by the human head model is as follows Fig.14 As shown in the figure, the composition and structure are rich, and various tissues such as gray matter, white matter, cerebrospinal fluid, cerebellum, eyes, and subarachnoid space can be seen. The proton density is normalized. The target conductivity distribution is set as Fig.15 As shown, the range of variation is [0,2.5]S / m. The target relative permittivity distribution is as follows Fig.16 As shown, the range of variation is [1,70].

[0337] Assuming that the magnetic properties of the head model are homogeneous, that is, the magnetic permeability is equal to the magnetic permeability of vacuum, pulsed time-harmonic continuous wave voltage excitation is implemented in the multi-channel transverse electromagnetic coil, and the amplitude and phase distribution of the sensitive field of the generated transmitting magnetic field intensity are shown in Figures a and b of Figure 17, respectively, with a range of (0, 25) A / m and (-3, 3) rad, respectively. The amplitude and phase distribution of the sensitive field of the receiving magnetic field intensity based on the reversed propagation direction of the transmitting magnetic field are shown in Figures a and b of Figure 18, respectively, with a range of (0, 25) A / m and (-3, 3) rad, respectively. It can be seen that the amplitude is consistent with the amplitude range of the transmitting magnetic field intensity, and the contrast of the phase distribution is opposite to the phase of the transmitting magnetic field.

[0338] The core inversion equation of the dual-source method has three unknowns: complex permittivity, complex permittivity transverse gradient, and complex permittivity longitudinal gradient. In order to reconstruct the distribution of electrical characteristic parameters by calculating the radio frequency magnetic field, this embodiment divides the coil excitation and detection single channels into four channel groups (1, 5, 9, 13), (2, 6, 10, 14), (3, 7, 11, 15), and (4, 8, 12, 16), and inverts the complex permittivity and its gradient based on the linear algebraic equations using four independent sets of transmission magnetic field data. According to the correspondence between the electrical characteristics and the real and imaginary parts of the solution, the conductivity distribution of the human head model is reconstructed as follows: Fig.19 As shown, the value range is [0,1.5] S / m; the reconstructed relative permittivity distribution is as follows Fig. 20 As shown, the value range is [1,70).

[0339] In another embodiment of the present invention, the distribution of the transmitting magnetic field and the receiving magnetic field in the human tissue is measured based on the existing clinical routine magnetic resonance imaging scanning, and the electromagnetic characteristic parameter distribution image is obtained by inversion and reconstruction using the Helmholtz equation method and the convection-reaction equation method. The specific scanning conditions are as follows: 1.5T magnetic resonance system 2D imaging mode, pulse sequence geometry parameters are field of view FOV 256*256mm 2 , the frequency encoding and phase encoding steps are both 128, the slice thickness is 4 mm, the inter-slice distance is 1 mm, the number of slices is 9, and the number of echo times is 1. Using a proton density weighted fast spin echo sequence (Proton Density Weighted Imaging, PDWI), a series of scans were performed on a homogeneous water phantom, a homogeneous water phantom containing an abnormal shape, a homogeneous water phantom containing the same abnormal shape with 2.5 g NaCl added, a homogeneous water phantom containing the same abnormal shape with 2.5 g NaCl and 0.25 mL of 377 mg / mL gadoteric acid dimeglumine (Gd-DOTA) added, and a volunteer's human brain tissue, and 24 groups of original magnetic resonance amplitude images and angle image data were obtained, and the results of the inversion of the electrical characteristic distribution by the magnetic resonance imaging radio frequency magnetic field and the inversion of the magnetic characteristic distribution by the present invention were compared. Taking advantage of the fast imaging speed of the fast spin echo sequence, the parameters of the three scans based on the proton density-weighted low T1 and T2 contrast sequence were (120°, 40ms, 10000ms), (60°, 40ms, 10000ms) and (60°, 20ms, 10000ms), respectively. The first parameter was the excitation deflection angle, the second parameter was the echo time, and the third parameter was the repetition time, so as to meet the TR>>3-4xT1 condition of eliminating the contrast factor term in the calculation of the magnetic field amplitude from the signal intensity as much as possible.

[0340] The spherical homogeneous water phantom data is used to assist in calculating the magnetic field measurement data of other water phantoms and subjects, including the correction of the non-uniformity of the coil magnetic field. Figures a and b of Figure 21 are respectively the weighted amplitude and angular images of the original magnetic resonance proton density obtained by scanning the homogeneous water phantom containing an abnormal shape. Figures a and b of Figure 22 are respectively the weighted amplitude and angular images of the original magnetic resonance proton density obtained by adding 2.5g NaCl to the homogeneous water phantom with the same abnormal shape. Figures a and b of Figure 23 are respectively the weighted amplitude and angular images of the original magnetic resonance proton density obtained by scanning the homogeneous water phantom with the same abnormal shape with 2.5g NaCl and 0.25mL of 377mg / mL gadoteric acid meglumine. Figures a and b of Figure 24 are respectively the weighted amplitude and angular images of the original magnetic resonance proton density obtained by scanning the human brain tissue.

[0341] The magnetic resonance amplitude and angle images of the spherical homogeneous water phantom and the target imaging body, three homogeneous water phantoms with abnormal shapes and human brain tissue acquired by scanning are processed to obtain magnetic field distribution data. The specific processing includes the following process.

[0342] (1) The amplitude data of the imaged body o scanned by the PDWI auxiliary sequence (120°, 40ms, 10000ms) is divided by the amplitude data of the imaged body scanned by the main sequence (60°, 40ms, 10000ms) to obtain the ratio data m o ; The amplitude data of the spherical homogeneous water model b scanned by the PDWI sequence (120°, 40ms, 10000ms) is divided by the amplitude of the homogeneous water model scanned by the sequence (60°, 40ms, 10000ms) to obtain the ratio data m b ; m o Divide by m b , and obtain the unit receiving magnetic field sensitivity field measurement data corresponding to the main sequence (60°, 40ms, 10000ms) scanning imaging body The magnetic field is related to the set echo time, and in addition to eliminating proton density coupling, it also eliminates the non-uniformity of the sensitive field amplitude of the transmitting coil magnetic field and the sensitive field amplitude of the receiving coil magnetic field.

[0343] (2) The amplitude data of the imaged body scanned by the PDWI auxiliary sequence (120°, 40ms, 10000ms) and the main sequence (60°, 40ms, 10000ms) are divided by the obtained unit receiving magnetic field sensitivity field measurement data. Get the ratio data t1 and t2; then divide t1 by t2 to get the cosine term data related only to the emission magnetic field with the tissue proton density weight eliminated. It is known that the width of the radio frequency pulse setting of the system excitation coil transmitting magnetic field is 4000μs. After calculating the inverse cosine of the cosine term data and dividing it by the gyromagnetic ratio and the pulse width, the transmitting magnetic field sensitive field amplitude measurement data corresponding to the main sequence (60°, 40ms, 10000ms) scanning imaging body is obtained. Note that this measurement is related to the excitation deflection angle.

[0344] (3) The phase data of the imaged body scanned by the PDWI main sequence (60°, 40ms, 10000ms) is subtracted from the phase data of the imaged body scanned by the auxiliary sequence (60°, 20ms, 10000ms) to obtain the phase difference data Δp o The phase data of the homogeneous water model scanned by the PDWI sequence (60°, 40ms, 10000ms) was subtracted from the phase data of the homogeneous water model scanned by the sequence (60°, 20ms, 10000ms) to obtain the phase difference data Δp b , which only includes the inhomogeneity contribution of the receiving coil magnetic field phase; then the imaging volume phase difference data Δp o Subtract the phase difference Δp of the homogeneous water model b , and considering that the echo time parameters of the two sets of sequences are 2 times, the received magnetic field phase measurement data of the main sequence (60°, 40ms, 10000ms) scanning imaging body is obtained as follows: Since the intermediate processing step includes subtracting the transmit magnetic field phase from the original magnetic resonance phase, the receive magnetic field phase does not contain the contribution of the transmit coil magnetic field phase inhomogeneity.

[0345] (4) The received magnetic field phase of the PDWI main sequence (60°, 40ms, 10000ms) scanning the imaging body has been calculated and processed. Where a_6_o is the original magnetic resonance phase of the main sequence (60°, 40ms, 10000ms), and the emission magnetic field phase measurement data corresponding to the main sequence can be obtained.

[0346] For each group of fast spin echo sequence scans implemented on each water phantom and cranial brain tissue, in addition to scanning a set of data with normal or default gradient polarity, another set of data is scanned by changing the gradient polarity without changing the sequence group parameters. The positive and negative polarity phase data of the main sequence (60°, 40ms, 10000ms) and the auxiliary sequence (60°, 20ms, 10000ms) are averaged respectively as the original magnetic resonance phase data of the corresponding sequence scanning subject, which is used to calculate the measured receiving magnetic field and transmitting magnetic field phase distribution of the corresponding subject. The refocusing flip angle of all the sequences must be 180°.

[0347] In addition, preferably, when the sequences (120°, 40ms, 10000ms) and (60°, 40ms, 10000ms) are used, the receiving bandwidth is adjusted so that the relative signal-to-noise ratio of the two scans of the same subject is equal or nearly equal, which is conducive to the results obtained in the calculation of the corresponding original magnetic resonance amplitude data of the measured receiving magnetic field and the transmitting magnetic field sensitive field of the corresponding subject to be consistent in amplitude. Under the condition that the imaging geometric parameters remain unchanged, the influence of changing the bandwidth and thus changing the sampling rate and the acquisition time on the spatial resolution can be compensated by the automatic magnetic field gradient intensity control function of the magnetic resonance equipment. For the sequences (60°, 40ms, 10000ms) and (60°, 20ms, 10000ms), considering that the echo interval of the former is longer, the total receiving bandwidth is set to 150kHz and 200kHz respectively, then the unit sampling time is 3.3μs and 2.5μs respectively, and the echo acquisition time is 427μs and 320μs respectively. According to the pixel resolution R RO With gradient G RO According to the relationship formula (55) between the acquisition time and the acquisition time, under the same resolution conditions, the gradients of the sequences (60°, 40ms, 10000ms) are calculated to be 27.5mT / m and 36.7mT / m, respectively, which are both less than the rated value of 45mT / m under the general magnetic resonance medical device standard.

[0348] In the embodiment, Figures a and b of Figure 25 are respectively the amplitude and phase distribution images of the transmitting magnetic field sensitive field obtained by scanning a homogeneous water model containing an abnormal shape, and Figures c and d of Figure 25 are respectively the corresponding amplitude and phase images of the receiving magnetic field sensitive field; in the figure, the water model contains a long slope organic glass body, so the corresponding area shows a noise distribution.

[0349] Figures a and b of Figure 26 are respectively the amplitude and phase distribution images of the transmitting magnetic field sensitive field obtained by scanning a homogeneous water model containing the same abnormal shape with 2.5g NaCl added, and Figures c and d of Figure 26 are respectively the corresponding amplitude and phase images of the receiving magnetic field sensitive field; in the figures, the gradual distribution of the magnetic field phase reflects the changes in the electromagnetic properties of the contents of the water model.

[0350] Figures a and b of Figure 27 are respectively the amplitude and phase distribution images of the transmitting magnetic field sensitivity field obtained by scanning the homogeneous water phantom containing the same abnormal shape with 2.5g NaCl and 0.25mL of 377mg / mL gadoterate meglumine, and Figures c and d of Figure 27 are respectively the corresponding amplitude and phase images of the receiving magnetic field sensitivity field; in the figure and the above-mentioned transmitting magnetic field and receiving magnetic field distribution diagrams reconstructed by actual scanning measurement calculation, in order to effectively assist in solving the transmitting magnetic field and receiving magnetic field distribution data in the homogeneous water phantom containing abnormal shape and the aforementioned brain tissue imaging based on the spherical homogeneous water phantom scanning data, the reconstructed image results generated are all the intersection of the spherical homogeneous water phantom and the target imaging body signal area.

[0351] Using the three proton density weighted fast spin echo sequences with different parameters, Figures a and b of Figure 28 are the excitation emission magnetic field sensitivity field amplitude and phase distribution images obtained by scanning the cranial brain tissue, and Figures c and d of Figure 28 are the corresponding receiving magnetic field sensitivity field amplitude and phase distribution images, respectively.

[0352] Based on the basic Helmholtz equation method that ignores the gradient of electromagnetic characteristics, the transmitting magnetic field and receiving magnetic field distribution data of the homogeneous water model containing abnormal shape scanned by the original magnetic resonance equivalent proton density measurement are substituted, and the conductivity and permittivity inverted and reconstructed are shown in Figures a, b, c and d of Figure 29 respectively. When calculating the magnetic field gradient and second-order derivative, the Savitzky-Golay filter design principle is used to solve the method of quadratic parabola fitting coefficient based on least squares, avoiding the decrease in accuracy caused by direct correlation division calculation. In the figure, since no regularization and normalization are performed, the displayed voxel electrical characteristic parameters are the original values ​​obtained by inversion. In the figure, the change in brightness reflects the change in the local CuSO4 concentration.

[0353] Based on the basic Helmholtz equation method that ignores the gradient of electromagnetic characteristics, the transmitted magnetic field and received magnetic field distribution data of the cranial brain tissue scan calculated by the original magnetic resonance equivalent proton density measurement are substituted, and the inverted and reconstructed conductivity and permittivity are shown in Figures a, b, c and d of Figure 30, respectively. In Figure c of Figure 30, the anatomical structure of the tissue such as the corpus callosum can be relatively clearly seen; Fig.32 The obvious salt and pepper noise in Figures (a) and (b) can be further processed by median filtering.

[0354] In another embodiment of the imaging scan of the magnetic resonance radio frequency magnetic field measurement of the present invention, in order to minimize T1 and T2 weighting and maximize PD weighting, the fast spin echo pulse sequence group parameters used are I (120°, 24ms, 9000ms), II (120°, 24ms, 9000ms) and III (120°, 32ms, 9000ms). For sequences I and II, assuming the receiving bandwidth is 62.5kHz and the frequency encoding step number is 128, the single point sampling time is 8μs and the single echo acquisition time is 1.024ms. Assuming the echo chain length is 8, considering the transmitting coil time constant (dead time) and the radio frequency power amplifier entering the working pre-start control time, the equal-interval echo interval is 6ms. According to the echo acquisition time, the frequency encoding readout gradient strength is calculated to be 11.46mT / m; and according to the definition of relative signal-to-noise ratio, the Rel.SNR is calculated from the bandwidth to be 28%.

[0355] For sequence III, the echo time is 32ms, and the echo interval is 8ms under the condition that the echo chain length is consistent with that of I and II. Since the echo chain is relatively backward on the transverse relaxation decay curve, it means that the signal-to-noise ratio decreases, and the bandwidth needs to be reduced to maintain amplitude consistency (although the phase calculation only involves the angular data of II and III scans). Assuming that the receiving bandwidth is halved to 31.25kHz, the sampling time and acquisition time are 16μs and 2.048ms respectively. On the whole, the single echo interval of 8ms is appropriate, and the gradient intensity is reduced to 5.73mT / m. The gradient slew rate of sequences I, II and III is small, which is conducive to avoiding the influence of eddy current effects, especially in angular scanning.

[0356] Although the above sequences I, II and III adopt the shortest possible echo time, achieving the purpose of PDW imaging and reasonably adapting to the inherent parameter constraints of magnetic resonance scanning, it will not help shorten the imaging time. Since TR is very long, the acquisition time of a fast spin echo chain under one excitation is much shorter than TR, and only one cycle acquisition (#Acquisition) is required for limited slice scanning. The phase encoding step of the sequence group in this embodiment is 128, so the time of sequence II and III is 9s*128 / 8=2:24s respectively, and the total scanning time is 7:12s.

[0357] The above description is only an embodiment of the present invention and is not intended to limit the present invention. For those skilled in the art, the present invention may have various modifications and variations. Any modification, equivalent substitution, improvement, etc. made within the spirit and principle of the present invention should be included in the scope of the claims of the present invention.

Claims

1. A method for measuring radio frequency magnetic field distribution in magnetic resonance imaging and inverting human tissue data, characterized in that: The method comprises the following steps: Step (1) Scanning and imaging body radio frequency magnetic field measurement data: Step (1.1) performs spin echo scanning on a homogeneous water phantom, and the sequence excitation deflection angle, echo time and repetition time parameters are (90°, TE, TR) respectively, to obtain amplitude image and angular image data; Spin echo scanning was performed on the homogeneous water phantom, and the parameters of the sequence excitation deflection angle, echo time and repetition time were (45°, TE, TR) respectively, to obtain the amplitude image and the angle image data; Spin echo scanning was performed on the homogeneous water phantom, and the parameters of the sequence excitation deflection angle, echo time and repetition time were (45°, TE / 2, TR) respectively, to obtain the amplitude image and the angle image data; Spin echo scanning is performed on the imaging human body part, and the parameters of the sequence excitation deflection angle, echo time and repetition time are (90°, TE, TR) respectively, to obtain amplitude image and angular image data; Spin echo scanning is performed on the imaging human body part, and the parameters of the sequence excitation deflection angle, echo time and repetition time are (45°, TE, TR) respectively, to obtain amplitude image and angular image data; Spin echo scanning is performed on the imaging human body part, and the parameters of the sequence excitation deflection angle, echo time and repetition time are (45°, TE / 2, TR) respectively, to obtain amplitude image and angular image data; In each of the above spin echo scans, the geometric parameters such as field of view, slice thickness, slice spacing, frequency encoding number, phase encoding step and number of slices were the same and kept consistent; Step (1.2) processes the homogeneous water model and the scan data of the imaged human body part, and obtains the radio frequency receiving magnetic field sensitivity field distribution data and the transmitting magnetic field sensitivity field distribution data of the human body tissue; Step (1.3) processes to obtain radio frequency receiving magnetic field phase distribution data and transmitting magnetic field phase distribution data of human tissue; Step (2) Invert the conductivity and permittivity from the magnetic field measurement data: Assuming that the magnetic permeability of human tissue is constant and equal to the magnetic permeability of vacuum, the distribution of conductivity and relative permittivity in the tissue is obtained by inverting the transmitted magnetic field or received magnetic field data induced by the human body using the magnetic resonance electrical characteristic imaging method; Step (3) Invert the magnetic permeability or magnetic susceptibility from the magnetic field measurement data and the existing electrical properties: The recovery assumes that the magnetic permeability of human tissue varies with space. The human body receives or transmits a magnetic field and combines the obtained electrical characteristic data to invert the distribution of magnetic permeability in the tissue based on the partial differential equation of the magnetic field; the electrical characteristic data includes electrical conductivity and relative permittivity.

2. The method for measuring the distribution of radio frequency magnetic field in magnetic resonance imaging and inverting human tissue data according to claim 1, characterized in that: The spin echo sequence adopts a 2D imaging scanning mode, i.e., a frequency encoding gradient is used as a slice selection gradient, or adopts a 3D imaging scanning mode, i.e., a phase encoding gradient is used as a slice selection gradient; When the 2D imaging scanning mode is adopted, it cannot be guaranteed that the sensitive fields and phases of the magnetic field measured between different layers are continuous in the longitudinal direction. When the electrical and magnetic properties are inverted based on the measured magnetic field, a 2D approximate algorithm is selected, that is, the longitudinal spatial partial derivative of the electromagnetic properties is considered to be equal to 0; When a 3D imaging scanning mode is adopted, it can ensure that the sensitive fields and phases of the magnetic field measured between different layers are continuous in the longitudinal direction, and a 2D approximation algorithm or a 3D volume algorithm is selected when inverting the electrical and magnetic properties based on the measured magnetic field.

3. The method for measuring the distribution of radio frequency magnetic field in magnetic resonance imaging and inverting human tissue data according to claim 1, characterized in that: The sequences are all based on the signal formation mechanism of fast spin echo rather than gradient echo, which is conducive to eliminating the inconsistency distortion of proton precession distribution caused by the weak inhomogeneity of the main static magnetic field. Therefore, the measured radio frequency magnetic field is only related to the inhomogeneity of the magnetic fields of the transmitting and receiving coils and the non-single material composition of the tissue, and is affected by the eddy current effect caused by the gradient coil.

4. A method for measuring radio frequency magnetic field distribution and inverting human tissue data by magnetic resonance imaging according to claim 1 or 3, characterized in that: The eddy current effect in the imaging body caused by the gradient coil is eliminated by implementing a double scanning method with opposite polarities of the gradient pulses and unchanged other sequence parameters. Averaging the two phase image data significantly reduces the spatial domain second harmonic interference in the phase distribution of the receiving magnetic field and the transmitting magnetic field.

5. The method for measuring the distribution of radio frequency magnetic field in magnetic resonance imaging and inverting human tissue data according to claim 1, characterized in that: In order to eliminate the inherent non-uniformity of the magnetic field sensitivity field distribution of the radio frequency excitation transmitting coil and the magnetic resonance signal detection receiving coil of the imaging scanning facility, a set of homogeneous water phantom scans are implemented in a sequence that is all used to scan the human body to measure the magnetic field distribution. The relevant coil field non-uniformity is eliminated in the data processing of calculating the receiving and transmitting magnetic field amplitudes.

6. The method for measuring the distribution of radio frequency magnetic field in magnetic resonance imaging and inverting human tissue data according to claim 1, characterized in that: The transverse relaxation time constant and longitudinal relaxation time constant of the imaging volume respectively control the rates of attenuation of the transverse component and recovery of the longitudinal component of the precession magnetization intensity vector. By keeping the echo time TE and repetition time TR of the two sequences the same but with different excitation deflection angles FA, during the echo signal acquisition, the magnetization intensity vector amplitude of the same voxel in the two scans is different but the direction is the same.

7. The method for measuring the distribution of radio frequency magnetic field in magnetic resonance imaging and inverting human tissue data according to claim 1, characterized in that: By keeping the echo time TE of the two scans at a fixed ratio and the deflection angle FA the same during the sequence execution, the purpose of achieving a constant phase difference between the transverse component of the precession magnetization intensity vector during the readout period in the two scans is achieved, thereby measuring the initial phase of the received magnetic field; the repetition time in the two scans is equal, so that the amplitude of the longitudinal magnetic vector is equal at the beginning of each sequence cycle, and the amplitude of the transverse magnetic vector remains relevant and consistent during the imaging signal acquisition.

8. A method for measuring radio frequency magnetic field distribution and inverting human tissue data in magnetic resonance imaging according to claim 1 or 6, characterized in that: When the imaging body is scanned with sequence parameters (90°, TE, TR) and (45°, TE, TR) respectively, TR is equal in the front and back scans to ensure that the amplitudes of the initial transverse component and the final longitudinal component of the magnetization intensity vector of each k-line acquisition cycle are equal. The equal TE ensures that the direction of the transverse component of the vector is always consistent during the precession process of echo acquisition, so that the amplitude part of the precession vector of the two scans can be divided according to a scalar to eliminate proton density coupling.

9. A method for measuring radio frequency magnetic field distribution in magnetic resonance imaging and inverting human tissue data according to claim 1 or 7, characterized in that: When the imaging body is scanned with sequence parameters (45°, TE, TR) and (45°, TE / 2, TR) respectively, the equal excitation deflection angles in the front and rear scans determine the specific attenuation mode of the transverse magnetic vector within a single sequence cycle, that is, the two echoes have amplitude correlation, and the relaxation echo time ratio is controlled to 1 / 2, so that the phase difference of the transverse component of the precession magnetic vector remains constant during the echo acquisition refocusing period, thereby achieving the elimination of the same excitation radio frequency pulse initial phase distribution contained in the received magnetic resonance signal phase, and the remaining phase amount is equal to 1 / 2 of the received magnetic field phase of the target scanning sequence; the received magnetic field phase is subtracted from the original image signal phase and then subtracted from the theoretical pi / 2 constant offset phase amount to obtain the transmitting magnetic field phase.

10. The method for measuring the distribution of radio frequency magnetic field in magnetic resonance imaging and inverting human tissue data according to claim 1, characterized in that: In order to adapt to the partial saturation phenomenon caused by insufficient recovery of the steady-state longitudinal magnetization intensity vector due to the fact that the measurement sequence group repetition time TR parameter does not meet the condition of being much greater than the tissue longitudinal relaxation time, a TR greater than 5 times the transverse relaxation time T2 is set according to the situation of the target imaging tissue to ensure that there is no residual transverse magnetic vector coherence at the end of a pulse sequence TR cycle; For imaging scans based on fast spin echo sequences, within one TR cycle, it is not possible to avoid the connection of the front and rear signals of adjacent echoes, that is, the echo time TE is required to be set greater than 5 times the magnetic sensitive transverse relaxation time T2* to allow the transverse magnetic vector to be fully decoherent.

11. A method for measuring radio frequency magnetic field distribution in magnetic resonance imaging and inverting human tissue data according to claim 1, 6 or 8, characterized in that: First, the amplitude data m_9_p collected by the scanning water phantom sequence (90°, TE, TR) is divided by the amplitude data m_4_p collected by the sequence (45°, TE, TR), and the intermediate process data m_9_p containing the sine ratio of the transmitting magnetic field amplitude and the normalized receiving magnetic field amplitude is obtained. p , where m represents the amplitude and p represents the water model; Then, the amplitude data m_9_h collected by the human body scanning sequence (90°, TE, TR) is divided by the amplitude data m_4_h collected by the sequence (45°, TE, TR), and the intermediate process data m_9_h containing the sine ratio of the transmitting magnetic field amplitude and the normalized receiving magnetic field is obtained, which eliminates the tissue proton density weight. h , where h represents the human body; Because the water model is homogeneous, the process data m h The inhomogeneity of the magnetic field sensitivity of the receiving coil in m p Correction, that is, the human body receives the magnetic field sensitive field It also contains the assumption or approximation that the magnetic permeability of the human body is homogeneous and equal to the magnetic permeability of the water model.

12. A method for measuring radio frequency magnetic field distribution in magnetic resonance imaging and inverting human tissue data according to claim 8, characterized in that: The human body's emitted magnetic field sensitivity field is calculated by the following steps: Step (1) divides the original amplitude data m_9_h collected by scanning the human body sequence (90°, TE, TR) by the human body receiving magnetic field sensitivity field Calculate the intermediate process value amplitude data t1; m represents the amplitude, and h represents the human body; Step (2) divides the original amplitude data m_4_h collected by scanning the human body sequence (45°, TE, TR) by the human body receiving magnetic field sensitivity field Calculate the intermediate process amplitude data t2; Both data t1 and t2 contain contributions from the macroscopic magnetization intensity, i.e., the proton density, and the sine value of the deflection angle related to the amplitude of the emission magnetic field; Step (3) Divide the intermediate quantity data t1 and t2 to obtain The inverse cosine of this ratio is used to obtain the target excitation emission magnetic field distribution induced inside the imaging human tissue. Where γ is the gyromagnetic ratio of the target nuclear isotope in magnetic resonance imaging, τ is the temporal width of the excitation radio frequency pulse, In order to transmit the magnetic field amplitude uniformly, the superscript 45 represents the excitation deflection angle implemented.

13. A method for measuring radio frequency magnetic field distribution in magnetic resonance imaging and inverting human tissue data according to claim 11 or 12, characterized in that: The receiving magnetic field sensitivity field is the magnetic field distribution generated when a unit current flows through the receiving coil, and its amplitude range is within the range of (0.5, 1.5) Tesla; the calculated amplitude of the transmitting magnetic field sensitivity field is on the order of 10 - 6 Tesla.

14. A method for measuring radio frequency magnetic field distribution in magnetic resonance imaging and inverting human tissue data according to claim 1 or 12, characterized in that: The implemented parameters are (45°, TE, TR) and the excitation deflection angle selected for the sequence is between 0° and 90°, ensuring that the time-quantity ratio t1 / t2 of the calculated arccosine function value is a positive number, and the emission magnetic field sensitivity field value falls within the range of (0,π / 2) / (γτ), γ is the gyromagnetic ratio of the target nuclear isotope of magnetic resonance imaging, τ is the time domain width of the excitation radio frequency pulse, and t1 and t2 are intermediate quantity data.

15. A method for measuring radio frequency magnetic field distribution in magnetic resonance imaging and inverting human tissue data according to claim 11 or 12, characterized in that: When the sequence group parameters (90°, TE, TR) and (45°, TE, TR) are used to scan the human body, the center value of the voxel amplitude histogram distribution of the received magnetic field sensitivity field may seriously deviate from 1, and the center value of the cosine value histogram distribution of the transmitted magnetic field sensitivity field amplitude may seriously deviate from 0.707 and approach or even exceed 1. Therefore, the sequence group parameters (120°, TE, TR) and (60°, TE, TR) are used to improve and adjust the calculated measurement magnetic field amplitude distribution to a reasonable statistical range without affecting the applicability and rationality of the magnetic field amplitude distribution measurement method, and without affecting the measurement of the magnetic field phase distribution.

16. A method for measuring radio frequency magnetic field distribution in magnetic resonance imaging and inverting human tissue data according to claim 1, 7 or 9, characterized in that: The phase of the human body RF magnetic field of the receiving coil is calculated by the following steps: Step (1) Subtract the angular data a_4_p collected by the sequence (45°, TE, TR) from the angular data a_t_p collected by the sequence (45°, TE / 2, TR) to obtain the phase difference Δp p ; a represents the argument, p represents the water model; Step (2) The phase difference Δp is obtained by subtracting the argument data a_4_h collected by the sequence (45°, TE, TR) from the argument data a_t_h collected by the sequence (45°, TE / 2, TR). h ; Step (3) calculates the phase of the induced receiving magnetic field in the human tissue for the target sequence (45°, TE, TR) as Wherein, the argument data a represents the angle; different from the sequences (90°, TE, TR) and (45°, TE, TR), t represents the sequence of short echo time TE / 2; the phase data of the scanned human body is subtracted from the phase data of the scanned homogeneous water phantom to correct the non-uniformity of the phase distribution of the receiving coil magnetic field.

17. A method for measuring radio frequency magnetic field distribution in magnetic resonance imaging and inverting human tissue data according to claim 1, 7 or 9, characterized in that: The phase of the human body induced RF transmission magnetic field is calculated by the following steps: Step (1) The phase of the induced magnetic field received in the human tissue corresponding to the target sequence (45°, TE, TR) is Subtract the original angle data a_4_h collected during the target scanning sequence to obtain the intermediate process phase data p M , which is the initial transverse component of the magnetization intensity vector M 0,⊥ The phase of Step (2) Set the phase p M Subtract π / 2 to get the phase of the magnetic field induced by the human body 18. A method for measuring radio frequency magnetic field distribution in magnetic resonance imaging and inverting human tissue data according to claim 6 or 7 or 8 or 9 or 11 or 12 or 16 or 17, characterized in that: The flip angle of the refocusing radio frequency pulse in the magnetic field measurement sequence is 180°, which ensures that the position of the precession magnetic vector component is symmetrical around the reference average speed magnetic vector in the transverse plane before and after flipping.

19. A method for measuring radio frequency magnetic field distribution in magnetic resonance imaging and inverting human tissue data according to claim 11, 12, 16 or 17, characterized in that: When using a sequence scan with an echo time parameter of TE, set the sampling data interval of each echo to td, collect N data for each echo, and the echo receiving time tacq is N*td. Assume that the received signal-to-noise ratio at this time is rSNR; When using a sequence scan with an echo time parameter of TE / 2, if the sampling data interval of each echo is set to td / 2 accordingly, and each echo collects N data, then the echo receiving time tacq is N*td / 2, and the receiving signal-to-noise ratio will change significantly, that is, decrease; In order to keep the signal-to-noise ratio of the two scans consistent, that is, the transverse components of the two magnetic vectors decay from the same amplitude and along the same curve under the action of the same excitation deflection angle pulse, td is adjusted in disguise by adjusting the bandwidth BW of the sequence (45°, TE / 2, TR); and because the single echo reception time tacq is related to the image resolution, without changing the imaging field of view and resolution, the magnetic resonance facility will automatically adjust the size of the magnetic field gradient G according to the set echo sampling data points N and tacq.

20. The method for measuring the distribution of radio frequency magnetic field in magnetic resonance imaging and inverting human tissue data according to claim 1, characterized in that: Based on the electromagnetic field and wave Maxwell equations, the magnetic characteristic parameter distribution is inverted by two different methods: Type I magnetic property inversion method: first assume that the magnetic permeability of the human body is homogeneous and equal to the magnetic permeability of vacuum μ0, use the existing magnetic resonance electrical property imaging inversion conductivity and permittivity algorithm, and use the measured transmitted magnetic field or received magnetic field to calculate the complex permittivity distribution of the tissue; Then, the permeability of the same imaged human body is considered to be non-homogeneous, and the remaining measured magnetic field data is used, or new radio frequency magnetic field data is measured again with different pulse sequence scanning parameters, and the permeability is inverted under the condition that the distribution of tissue electrical characteristic parameter values ​​is known; Type II magnetic property inversion method: Under the condition that the magnetic permeability of the human body is also non-uniform, a partial differential equation is constructed with multiple sets of magnetic field measurements as known quantities and magnetic permeability, electrical conductivity and relative permittivity as unknown quantities to simultaneously invert the relevant magnetic and electrical properties.

21. The method for measuring the distribution of radio frequency magnetic field in magnetic resonance imaging and inverting human tissue data according to claim 20, characterized in that: The first type of magnetic property inversion method is to use the remaining standby magnetic field measurement data, or change the sequence parameters to measure the same imaging body again to obtain a new available radio frequency magnetic field after obtaining the distribution parameters of the human body conductivity and relative permittivity, and then combine the inverted electrical property parameters to invert the human body tissue magnetic permeability μ according to the following equation: Where ω is the resonant Larmor angular frequency, is the obtained complex permittivity of the imaging body, μ is the magnetic permeability of the imaging body to be determined, B is the measured magnetic field vector, is the Laplace operator, is the gradient operator, '×' is the curl operator; it contains the assumption twice that the magnetic permeability of the human body is equal to the magnetic permeability of the vacuum to simplify the scalar and vector derivation by applying Gauss's magnetic law, and obtains the core equation for inverting the magnetic permeability by measuring the magnetic field.

22. The method for measuring the distribution of radio frequency magnetic field in magnetic resonance imaging and inverting human tissue data according to claim 20, characterized in that: The type II magnetic property inversion method is to invert the electric and magnetic properties synchronously based on the Maxwell equations, that is, to invert the three unknowns of magnetic permeability, electrical conductivity and relative permittivity with known magnetic field measurements, and derive the following equations: In the formula, is the Laplace operator, 'v' is the inverse of magnetic permeability, is the complex permittivity, including conductivity and relative permittivity; To measure the transmitted magnetic field excitation amplitude, To measure the receiving magnetic field, a conjugate is taken; Further investigation Item, B is or The fluid mechanics model can be expanded into convection terms and diffusion terms to obtain the core calculation equations for the magnetic field components: By changing the pulse sequence parameters and performing two scans, two sets of independent known field quantity data are obtained. A highly well-posed equation group containing four equations is constructed based on the core equation form to solve the unknown electromagnetic characteristic parameter distribution data.

23. The method for measuring radio frequency magnetic field distribution and inverting human tissue data in magnetic resonance imaging according to claim 20 or 21, characterized in that: When deriving the first type of equation for calculating magnetic permeability under 2D magnetic resonance tomography scanning conditions, the z component B of the magnetic field transmitted by the excitation coil and the magnetic field induced by the detection coil parallel to the main static magnetic field direction is ignored. z , and when solving the derivative of the electrical properties in a specific layer, it is assumed that its z-direction gradient is 0, that is, B z = 0, and the electrical characteristics between adjacent layers are equal, there is 24. A method for measuring radio frequency magnetic field distribution in magnetic resonance imaging and inverting human tissue data according to claim 21 or 22, characterized in that: The first-order partial derivatives and second-order partial derivatives in each direction of the radio frequency magnetic field measurement are obtained by least squares operation fitting based on the Savitzky-Golay filter construction principle method; for the data point matrix of the 3x3 operation kernel, the complex value measurement magnetic field B at the middle position, i.e., the coordinate point (x, y), is the fitting target quantity, and the new value obtained by fitting is: Where c1 to c6 are fitting coefficients, (x i ,y i ) is the coordinate of the adjacent field point, i = 1, 2, 3, 4, 5, 6, 7, 8; after implementing the quadratic parabola numerical fitting, the magnetic field gradient and second-order partial derivative at the target point (x5, y5) are calculated by the following formula: In the formula, It means to find partial derivatives in space.

25. The method for measuring radio frequency magnetic field distribution and inverting human tissue data in magnetic resonance imaging according to claim 21 or 22, characterized in that: The repetition time TR parameter of the measurement sequence group is greater than or equal to 4 times the longest longitudinal relaxation time T1 of the tissue component, so as to satisfy the calculation of the signal intensity SI based on the spin echo formation mechanism. SE The conditions under which the tissue and sequence parameter contrast can be factored are: SI SE The expression is: Where M0 is the steady-state magnetization, α is the excitation deflection angle, β is the refocusing flip angle, TE is the echo time, T2 is the transverse relaxation time, and S is the receiving sensitive field; when TR ≥ 4T1 and β ≡ 180°, the signal intensity is reduced to: In this way, the contrast is achieved when calculating the measured magnetic field amplitude. The parameters of the item sequence are factored to achieve the emission magnetic field and receiving magnetic field for purposes only related to sinα(r) and S(r), respectively.

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