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

Abstract

The present application relates to a kind of magnetic resonance imaging radio frequency magnetic field distribution measurement and human tissue data inversion method, comprising the following steps:1, respectively in spin echo sequence (90 °, TE, TR), (45 °, TE, TR) and (45 °, TE / 2, TR) scanning homogeneous water model and imaging body, obtain amplitude and argument data;2, by the amplitude data of sequence (90 °, TE, TR), (45 °, TE, TR) is calculated to eliminate the sensitive field distribution of the emission magnetic field and receiving magnetic field of imaging body coupling with proton density, the phase distribution of the emission and receiving magnetic field of imaging body is calculated by the argument data of sequence (45 °, TE, TR) and (45 °, TE / 2, TR). Based on the magnetic resonance electric characteristic imaging, the complex permittivity of the imaging body is inverted from the emission magnetic field measurement quantity, and then the magnetic permeability distribution of the tissue is further inverted from the receiving magnetic field and the existing complex permittivity.

Description

Technical Field

[0001] This invention belongs to the field of medical electromagnetics research, specifically relating to a method for measuring the radio frequency magnetic field distribution in magnetic resonance imaging and for inverting human tissue data. Background Technology

[0002] Based on their magnetic properties, substances in nature are mainly classified into three categories: diamagnetic, paramagnetic, and ferromagnetic. The main components of human tissue are water, proteins and fats composed of elements such as carbon, hydrogen, and oxygen, and trace amounts of metallic elements. Macroscopically, human tissue exhibits weak diamagnetic properties, expressed as a real negative magnetic susceptibility, characterizing the degree of magnetization in an external magnetic field. When magnetic resonance imaging (MRI) scans are involved, human tissue is exposed to the main static magnetic field B0 and the magnetic field gradient B... G The effect of magnetization in χ DC and in alternating excitation magnetic field The effect of magnetization in χ RF The former χ DC The impact on the proton density image is mainly manifested in the non-uniformity of the resonant Larmor frequency distribution, and consequently, the spatial fluctuation of the precession phase. The latter χ RF The primary influence is on the uniformity of the amplitude and argument distribution of the excitation magnetic field, which in turn affects the consistency of the distribution of the transverse vector formed by the steady-state magnetization intensity deflection within the fault. This, combined with longitudinal and transverse relaxation effects, results in different contrast-weighting mechanisms on the image. The effects of static magnetization can be effectively eliminated using radio frequency spin echo scanning techniques, while alternating (radio frequency) magnetization is one of the objects of electromagnetic characteristic imaging magnetic field measurement and numerical inversion described in this invention. According to Maxwell's equations, alternating excitation and reception magnetic fields generate conduction and displacement currents in tissues related to the induced electric field. The relevant characteristic parameters are denoted by conductivity σ and permittivity ε, which are the other two measurement and inversion objects of this invention.

[0003] Magnetic susceptibility χ m Both magnetization and permeability μ are used to characterize the magnetic properties of materials, and the following derivation relationship exists between them. For linear media, magnetization is defined as M = χ. m H is the intensity of the applied magnetic field. Expressing the magnetic field intensity as the magnetic flux density formed in the medium, we have H = (B / μ0) - M, where μ0 is the permeability of free space. 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 It is the relative permeability of the medium. The target parameter for inversion in this invention is μ. r =1+χ m , χ mAlso called volume magnetic susceptibility, as distinguished from mass magnetic susceptibility and molar magnetic susceptibility. For electromagnetic property imaging based on nuclear spin radio frequency magnetic field resonance, the contributions of B0 and B G Related μ DC is eliminated, the measured excitation transmit and detection receive magnetic fields are inverted to the property distribution, respectively, with:

[0006]

[0007] where, and are the amplitudes of the transmit and receive magnetic induction vectors, respectively, is the equivalent externally applied magnetic field strength, and the'' indicates that the field quantity is related to the effect of a unit current flowing through the receive coil. Since there is no external magnetic field present during the signal reception period, is directly proportional to only the magnetization M developed inside the medium, which has the same unit as H.

[0008] A special case is in the field of medical electromagnetics, where the magnetic susceptibility is usually used instead of the magnetic permeability to characterize the magnetic properties of human tissues. The reason is that most kinds of biological tissues have weak magnetic susceptibility, approximately at the level of -5.0 x 10 -6 -6, so it is more convenient to study the electromagnetic effects in ppm units. For inorganic ferromagnetic materials, r = 1 + χ m ≈ χ m , the magnetic permeability is normally used as the property analysis parameter.

[0009] The interaction of a medium with an electromagnetic field or a time-harmonic wave can be described by three parameters, the magnetic permeability, the electric conductivity, and the permittivity (μ, σ, ε). In analyzing the electrical properties of weakly conductive biological tissues, the effects of temperature, electromagnetic frequency response, and material physical state are usually ignored, and the tissue is assumed to be a homogeneous anisotropic medium. During the magnetic resonance scanning process, the time-harmonic magnetic flux density induces weak interference induced currents and displacement currents:

[0010]

[0011] where, in the rectangular coordinate system, is the vector of three orthogonal partial derivative components, 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 amplitudes of the two currents is J C / J D = σ / ωε. In the magnetic resonance radio frequency band, the human body has J C and J DThe magnitude levels are of the same order of magnitude, while J D increases, for example, for fat, gray matter, muscle and cerebrospinal fluid at 7.0T (300MHz) J C D are approximately 0.4, 0.6, 0.7 and 1.7, respectively.

[0013] Prior art magnetic resonance electrical property imaging and the present magnetic resonance electromagnetic property imaging rely on the measurement of the distribution of the radio frequency magnetic field. In the coordinate system rotating at the resonance Larmor frequency, the process of magnetic resonance imaging involves right-handed excitation magnetic field (propagating from the transmit coil to the imaging body) and left-handed detection magnetic field (propagating from the imaging body to the receive coil) in the acquisition of signals with radio frequency and gradient pulse sequences scanning the human tissue. In the Cartesian coordinate system, the magnetic field strength or the magnetization strength equivalent magnetic field strength H x and H y are expressed as:

[0014]

[0015] where (B + ,B - ) and (H x ,H y ) are complex numbers defined on the complex plane, and '*' represents taking the conjugate. Essentially, B + and B - or H + and H - are equal in status, i.e. in the inversion process of electromagnetic properties, different magnetic field distribution data sets measured by different excitation parameters for the same imaging body should be inverted to 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 the imaging of anatomical structure of human tissue in morphology, which includes voxel-based spatial position encoding and image reconstruction. In the conventional magnetic resonance imaging process, image reconstruction is completed based on Fourier inverse transform from wave number domain to image domain. When the human body is placed in a static magnetic field B0, the hydrogen proton spin system in the body forms a macroscopic magnetization vector where 0 represents the initial state before the start of resonance excitation, M0 is the amplitude and is greater than 0, M is negative along the z-axis, and the negative sign means that the spin system is in a low-energy level steady state that is not excited, and indicates the diamagnetism exhibited by the hydrogen-containing proton substance. For a nuclear spin system in thermal equilibrium state in the sample, the amplitude of the macroscopic magnetization is related to the spin density, static magnetic field strength and system temperature T s , etc., and for a homogeneous sample:

[0017]

[0018] where N​s γ is gyromagnetic ratio, γ = 42.576 MHz / T for hydrogen proton, h is the reduced Planck constant, I is the spin quantum number, k is the Boltzmann constant, T s T is the temperature of the imaged body. For imaging with hydrogen nuclei 1 H is the imaged nucleus, The part is an invariable constant. The number or density of hydrogen protons contained in different organs and tissues of different parts of the human body is not the same. The applied static magnetic field B0(r) = μ(r)H0(r) generated by the superconducting or permanent magnet of the magnetic resonance scanner exists not only a small change related to the spatial position r, but also an actual distribution changes with time, which can be improved or even ignored the influence of field inhomogeneity through the shimming and locking system. The magnetic resonance imaging is also affected by certain temperature changes of the human body, including physiological or pathological and specific absorption rate effects of radio frequency during scanning, temperature T s which is also a quantity varying with space and time. In summary, under strict conditions, the magnetization vector has

[0019] The radio frequency excitation transmission magnetic field for imaging is perpendicular to the B0 direction, and ‘_’ represents a complex value, i.e. has periodicity, and the frequency is the same as the precession frequency of M, wherein represents the amplitude of the circularly polarized magnetic field, represents the initial phase. M and are in the same direction, and form a flip angle after interaction τ is the radio frequency pulse width, and the generated transverse component can be expressed by the following formula:

[0020]

[0021] In the formula, i is the imaginary unit, which means that there is a phase increment i.e. if is applied along the positive direction of the x axis, for the right-handed precession M, the rotation will be overturned at the end of the rotation and will be towards the positive direction of the y axis.

[0022] The electromagnetic characteristic imaging is carried out by using the whole-body large-size birdcage type body excitation transmission coil commonly used in clinic and the head and neck separate receiving coil, based on specific imaging scan sequence parameters, i.e. setting specific (α, TE, TR), wherein α is the excitation deflection angle, TE is the echo time, and TR is the repetition time, two-dimensional tomography is implemented to measure the sensitivity field amplitude and phase of the radio frequency magnetic field distribution. And a certain frequency or phase encoding selection layer gradient, frequency encoding gradient and phase encoding gradient are set to determine the imaging field of view and voxel size, and the wave number domain but the time domain complex value represented by the lumped signal is:

[0023]

[0024] where r is the voxel position vector; M ⊥ and M 0,⊥ The initial phase of and is ω0t and The received magnetic field 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 the magnetic resonance signal; B0and the magnetic field gradient B G The initial phase generated by the joint action of the nuclear spin is B0, and if there is only B0 V is the volume of the imaging region.

[0025] For the image domain data reconstructed from the scan k-space line array original data, it contains both amplitude and phase data, and the equivalent spin density m(r) is defined as follows:

[0026]

[0027] where, that is, the amplitude image commonly displayed on medical imaging film, is the phase image. The prior art "radio frequency field map" B1map is a measurement of the amplitude of the radio frequency excitation transmission magnetic field, and the distribution of B1map is obtained through a special scanning sequence and data post-processing , but it cannot be measured because the end of the transmission phase and the beginning of the reception phase cannot be defined in the transceive phase and needs to be "found another way out". Because the human eye has special biological structure and function that is sensitive to amplitude ratio and insensitive to phase difference, general radiology departments do not provide phase image films as a basis for differential diagnosis. In the gradient echo-based imaging method, B0and B G interact with non-homogeneous tissues to cause distortion in the distribution of the phase image related to the magnetic susceptibility, so the 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, which is restricted by the imaging principle itself and the difficulty of calculating the mapping from phase to dipole field source is very large, and the clinical application level of this kind of imaging technology needs to be improved.

[0028] Since there is a coupling relationship between magnetization and the unit received magnetic field in the magnetic resonance detection signal, existing technologies mainly focus on the measurement method of the emitted magnetic field distribution. According to the principle of nuclear magnetic resonance imaging, the amplitude spatial distribution of an ideal emitted magnetic field should be uniform, thereby exciting the magnetization vector to form a uniformly distributed deflection angle in the transverse direction. Furthermore, since the permeability of the human body can generally be considered to be approximately constant and equal to the permeability of vacuum, the emitted magnetic field in equation (6) is different from the received magnetic field. The symbol does not have a wavy line, and the phase angle is... It is also approximately uniformly distributed, and its actual non-uniformity depends only on the geometry of the transmitting coil.

[0029] Existing technology also proposes a dual-angle method to measure the amplitude distribution of the emitted magnetic field. The same imaging body is scanned twice with gradient echo sequences having excitation deflection angles of α and 2α, and the resulting image amplitude data are divided to obtain:

[0030]

[0031] Because the flip angle 2α is set to not exceed 2π, the excitation emission magnetic field is finally obtained by dividing the image data and taking the inverse cosine. The amplitude distribution. Although according to the reciprocity theorem, the received magnetic field amplitude in the equivalent spin density m(r) expression of formula (8) is... 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 a continuous scanning sequence, directly dividing the magnitudes of the magnetic vectors in pixels with different orientations without considering the (TE, TR) condition does not conform to physical principles. In magnetic resonance imaging scans, images obtained by setting different echo times and repetition times show specific contrasts between tissues with different longitudinal and transverse relaxation times. This is essentially an adjustment of the longitudinal magnetic vector M. || Recovering steady state and initial transverse magnetic vector M 0,⊥ The process of reducing control is crucial. Therefore, it is incorrect to forcibly treat the amplitude image, i.e., the reconstructed complex equivalent spin density amplitude, as the transverse magnetic vector amplitude to calculate the emission magnetic field amplitude distribution.

[0032] To increase the sensitivity of measuring the amplitude distribution of the emitted magnetic field, Smith et al. proposed a dual-angle method where the refocusing flip angle β is equal to twice the excitation deflection angle α. For spin-echo SE scanning sequences, 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] In the formula, x represents the voxel position, and C is a contrast variable related to tissue characteristics and sequence parameters. When β equals twice α, the signal amplitude is proportional to the cube of sinα. Furthermore, the quotient λ of the amplitude image data obtained from two scans can be used to further express this characteristic. SE The distribution of the emitted magnetic field amplitude can be calculated using the following formula.

[0035]

[0036] Although theoretically this method has the optimal measurement sensitivity, the function sin 3 The slope is greatest when θ = 60° and 120°. However, due to the refocusing flip angle not being equal to 180°, insufficient deflection of the lateral component of the magnetic vector, uneven distribution of the magnetic vector on both sides of the average velocity position, and the tilt of the refocusing plane (i.e., the lateral component is sparser on one side and denser on the other), the echo signal is distorted. Combined with the effect of the spatial coding gradient pulse group of each sequence period, this causes severe interference artifacts in the original magnetic resonance image.

[0037] Direct measurement or indirect assessment of the phase component distribution of the transmitted magnetic field is more difficult. Current technology suggests that on conventional clinical MRI scanners, the transmitted phase can only be measured using a shared-transmit / receive RF coil, such as a birdcage coil. Compared to a single-transmit / receive RF excitation-detection method, a single-channel shared-transmit / receive coil, because it swaps the signal-ground polarity at its port when acting as a receiving coil to detect the MRI induced signal, creates an advantageous condition for processing the phase of the equivalent spin density, allowing the transmitted and received magnetic field phases to be considered approximately equal in value but opposite in sign. The following equation gives the equivalent spin density under this "transmit / receive phase assumption."

[0038]

[0039] In the formula, That is, circularly polarized single-rotation emission magnetic field initial phase Although the amplitude during imaging excitation The distribution is approximately uniform, but because it includes the permeability parameter, its distribution is closely related to the actual properties of the tissue. Similarly, the phase component... It must also be related to the location of the voxels.

[0040] The classic Bloch equations completely describe the longitudinal and transverse relaxation processes of a spin system in a magnetic field. Considering the macroscopic magnetization vector in each voxel as a precessing arrow, after a radio frequency excitation magnetic field pulse, its horizontal component undergoes circular motion at the Larmor frequency while its amplitude gradually decreases. From the perspective of voxels, the horizontal component of the vector gradually decouples; while the longitudinal component, without considering phase angle factors, gradually recovers its amplitude from zero to approach its original steady state. Combining the principles of pulse excitation and echo signal detection in magnetic resonance imaging (MRI) scan sequences, how to calculate the transmitted and received magnetic field data from the amplitude and phase angle 2D tomographic complex numerical matrix reconstructed from the image, based on and utilizing the basic processing methods in the laws of electromagnetic wave propagation and interaction—amplitude ratio and phase difference—and combining the characteristics of vectors and their evolution, to obtain the corresponding sensitive field (amplitude) and phase (unwinding phase angle) distributions, is the foundation for electromagnetic imaging and other related biomedical technologies and applications derived from MRI, including EEG, MRI, ECG, and ECG.

[0041] This invention focuses on the measurement of the transmitting and receiving magnetic fields separately, based on the conventional coil facilities and established sequence techniques used in clinical magnetic resonance imaging (MRI). This is used to reconstruct the inversion distribution of the electrical and magnetic properties of human tissues, thereby further assisting in the differential diagnosis of diseases and identifying and assessing problems that cannot be detected by existing morphological anatomical imaging. Regarding electrical properties, human tissues exhibit both conductive and insulating dielectric properties; while magnetically, the human body as a whole can be considered to have weak diamagnetism, i.e., a permeability slightly less than 1 and a negative magnetic susceptibility. To describe the behavior of the radiofrequency magnetic field in MRI, under time-harmonic conditions, according to Faraday's law and the Maxwell-Ampère law in Maxwell's equations:

[0042]

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

[0044]

[0045] In the formula, Complex capacitance Taking the curl of the second equation in equation (14) and substituting the first equation into the result, or substituting the second equation into the first equation to eliminate the electric field quantity, and assuming that the magnetic permeability of the human body is equal to the magnetic permeability of vacuum, 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 for measuring radio frequency magnetic field by using magnetic resonance imaging to invert the distribution of electrical property parameters of human tissue. In the imaging operation and data processing flow, it is mainly divided into implementing pulse sequence scanning to collect the initial amplitude and phase angle data of magnetic resonance imaging, substituting the data into the operation formula to obtain the magnetic field sensitive 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 real electromagnetic properties of human tissue are very complex, with anisotropy and frequency dispersion effects, it is generally necessary to develop property inversion method research under the assumption of homogeneous tissue. According to whether there is a gradient assumption of ignoring properties and the method of measuring magnetic field sensitive field and phase, magnetic resonance electrical property tomography 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 the complex permittivity, i.e. there will be systematic errors at the interface of different tissues in the inversion result. Taking the transmitting magnetic field as an example, the core equation is obtained as:

[0048]

[0049] According to the correspondence between the real part and the imaginary part of the complex solution of the equation, the conductivity and permittivity values of each voxel are finally calculated. However, the main problem of this method is that when the transmitting magnetic field is divided into amplitude and phase parts to calculate the permittivity and conductivity respectively, the expression for calculating the relative permittivity from the 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 transmitting initial phase, but there is currently no reasonable measurement technology for the real phase distribution.

[0050] A method of electrical property inversion considering the gradient factor of complex permittivity is given in the prior art, Gradient Electrical Properties Tomography (gEPT). Since the complex permittivity also changes with space like the field itself, the derivation of the core equation composed of scalars and vectors of the method increases the difficulty and complexity. Compared with the Helmholtz equation method, the applicable conditions of the gradient electrical property imaging method have many differences from clinical magnetic resonance imaging, and the magnetic field measurement work needs to be completed by combining a special multi-channel independently controllable pulse transmitting and receiving common coil, and is divided into two steps of first inverting the electrical property gradient distribution from the magnetic field data, and then inverting the electrical property by using the least square method from the gradient. The gradient term is defined as g Still taking the calculation of the electrical property from the transmitting magnetic field as an example, the related core equation is:

[0051]

[0052] In the formula, represents the magnetic field sensitive field of the i th channel transmitting and the rest of the channels receiving when the phase of the r th channel is taken as a reference; is the transmitting reference phase of the r th channel; the right-handed component g + of the gradient vector is defined as g + = g x + ig y . Although this method improves the distortion problem at the tissue interface caused by the electrical property homogeneity assumption, it needs to overcome the ill-posed problem of the solution in the calculation process, and the rationality of the method still needs to be verified, especially whether the many unconventional excitation and detection methods used in magnetic resonance are reasonable is still controversial.

[0053] To say goodbye to the problems of indirect solution from the gradient and multi-channel receiving and transmitting common coil which are not generally used in clinical 3T and below, the prior art proposes a convection-reaction Electrical Properties Tomography (crEPT). The method is similar to solving typical partial differential equations in fluid mechanics in form, and more importantly, the inversion core equation of the complex permittivity and the magnetic field satisfies a linear relationship. Still starting from the formula (14) of the Maxwell equation set, the core equation of the convection-reaction method containing the gradient factor of the electrical property which is homologous to the gradient electrical property imaging is derived:

[0054]

[0055] In the formula, u is the reciprocal of the complex permittivity. Further introduce the auxiliary vector The above core equation is simplified to the equation about u:

[0056]

[0057] where a is an arbitrary set diffusion coefficient to control the stability of the equation solution. With the above equation, after setting the initial values of the electrical properties on the boundary of the solution domain, i.e. Dirichlet boundary condition, the complex permittivity distribution of the electrical properties imaging can be solved iteratively. In addition to solving the electrical properties with the measured transmit magnetic field, the electrical properties can also be solved with the received magnetic field, from the physical point of view, is propagated from the boundary of the solution domain to the inside, is propagated from the inside of the solution domain to the boundary. In theory, although the amplitude and phase distribution of the transmit magnetic field and the received magnetic field used may be different, the electrical properties results obtained should tend to be the same. Although the above convection-reaction equation method does not ignore the electrical property gradient near different tissue interfaces, only one measured known quantity is needed to invert a unique unknown quantity, but 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 property features are relatively significant, and even can be reflected and applied in the auxiliary differential diagnosis of clinical imaging. The magnetic evolution process of hemoglobin molecules at different stages in the hematoma accompanying cerebral hemorrhage has obvious contrast differences in the magnetic resonance imaging image. The magnetic contrast agent injected in the enhanced scan of the magnetic resonance imaging of brain tumor patients can control the contrast of the image to some extent to see the lesion and help the doctor to distinguish the radioactive necrosis, false progression and early diagnosis of tumor. Therefore, including the existing susceptibility weighted imaging, quantitative susceptibility imaging, perfusion imaging and blood oxygen level dependent functional magnetic resonance imaging and other potential magnetic property imaging methods, they all have important application value in clinical medicine.

[0059] Magnetic resonance electrical property imaging and related magnetic property imaging can measure the electromagnetic property parameter distribution of human tissues. Some physiological and pathological changes sometimes cannot be reflected by structural imaging, and electromagnetic property imaging as potential functional imaging is expected to play a role in the diagnosis of intractable and difficult-to-find tumors, cerebral vascular diseases, breast cancer, prostate cancer and other malignant diseases. However, in existing magnetic resonance electrical property imaging, it is difficult to effectively, accurately and reliably measure the radio frequency transmit magnetic field and received magnetic field distribution. Most measurement methods are based on single-channel birdcage type transmit-receive shared coils that are no longer used in clinical practice, or rely on super-high field transverse electromagnetic multi-channel independent receiving and transmitting controllable coils that only exist on research magnetic resonance systems. Moreover, there are generally homogeneous assumptions and magnetic permeability vacuum approximations in electrical property inversion, and it is often difficult to completely measure the amplitude and phase information of the magnetic field, to seek more complex calculation formulas and physical quantity processing methods. SUMMARY

[0060] In order to solve the problem that the received magnetic field distribution of the radio frequency detection coil cannot be measured and the transmit magnetic field sensitive field and phase distribution cannot be measured simultaneously in the existing magnetic resonance electrical property imaging, the present application provides a magnetic resonance imaging radio frequency magnetic field distribution measurement and human tissue data inversion method, which calibrates the sensitive field and phase distribution data of the measured magnetic field by means of a homogeneous water phantom, and relates to a method for independently measuring two kinds of magnetic field distributions respectively, and the measurement results include two sets of data of the sensitive field amplitude component and the phase component of the magnetic field. The present application also relates to a human tissue data inversion method, which is a method for inverting the electrical property and magnetic property parameter distribution in the human tissue based on the magnetic resonance radio frequency magnetic field inversion imaging, and particularly relates to a method for first inverting the complex permittivity by using one set of measured magnetic field, and further inverting the magnetic permeability by using another set of magnetic field combined with the complex permittivity.

[0061] In order to solve the problem that the received magnetic field amplitude and the proton density are not decoupled in the magnetic resonance signal in the existing magnetic resonance electrical property imaging, and the birdcage type body coil excitation and multi-channel coil receiving configuration mode of the clinical magnetic resonance cannot meet the demand of the multi-component discrete magnetic field measurement in a single scan in the electrical property imaging, the present application can measure the sensitive field and the phase distribution of the received magnetic field and the transmit magnetic field. The core inversion equation is derived based on the elimination of the electric field of the Maxwell equation set, two complex value known quantities of the received magnetic field and the transmit magnetic field are used to solve two unknown quantities of the electrical property and the magnetic property parameter, wherein the electrical property is the complex permittivity, and the magnetic property is the real value magnetic permeability.

[0062] It should be noted that, like the existing magnetic resonance electrical property imaging method, the present application does not consider the anisotropy of the human tissue in the calculation process of inverting the electromagnetic property parameter, that is, for a given imaging target fault, it is assumed that the measurement in the direction other than the radio frequency circular polarization direction related to the excitation and detection magnetic field will invert the same electromagnetic property. It is assumed that the electromagnetic property of the human tissue has no frequency dispersion effect in the magnetic resonance radio frequency magnetic field frequency band. It is assumed that the human tissue is a magnetic linear medium, that is, when a 90° deflection angle pulse sequence is excited and a 45° pulse sequence is used for scanning imaging, a magnetic flux density B proportional to the magnetic field strength corresponding to the pulse is generated in the human tissue, and the related proportional coefficient is equal to the magnetic permeability μ. In addition, the human body has weak diamagnetism, and the typical magnetic hysteresis loop, magnetic memory effect and other effects of ferromagnetic substances are not considered. Unlike the permittivity, the magnetic 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. Otherwise, the change of the magnetic permeability and the complex permittivity with temperature is also ignored.

[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, and the eddy current caused by the rising edge and the falling edge of the gradient magnetic field can be reduced by changing the polarity twice and averaging the phase image data. In the magnetic resonance imaging gradient echo technology, because the signal reading process retains the static magnetization difference in the divergent phase relationship, the aforementioned magnetization weighting and quantitative magnetic susceptibility imaging are used to reproduce the low-frequency magnetic properties of the tissue, and in the latter, the dipole field source model inversion is the target. In the magnetic resonance electrical property imaging and the magnetic resonance electromagnetic property imaging of the application, a spin echo or fast spin echo sequence is used to collect the original magnetic resonance image data, the static magnetization difference is compensated by a high amplitude radio frequency refocusing pulse, including the contributions of B0 non-uniformity and gradient magnetic field spatial encoding non-uniformity, the transmitting magnetic field and the receiving magnetic field distribution containing the influence of radio frequency coil non-uniformity and human tissue electromagnetic property difference are measured and calculated, and the electromagnetic property parameter distribution in the radio frequency band is obtained by inversion.

[0064] To achieve the above object, the application adopts the following technical scheme:

[0065] A magnetic resonance imaging radio frequency magnetic field distribution measurement and human tissue data inversion method, the method comprises the following steps:

[0066] Step (1) scanning the radio frequency magnetic field measurement data of the imaging body:

[0067] Step (1.1) spin echo scanning is performed on the homogeneous water phantom, the sequence excitation deflection angle, echo time and repetition time parameters are (90°, TE, TR) respectively, and the amplitude image and the argument image data are obtained;

[0068] Spin echo scanning is performed on the homogeneous water phantom, the sequence excitation deflection angle, echo time and repetition time parameters are (45°, TE, TR) respectively, and the amplitude image and the argument image data are obtained;

[0069] Spin echo scanning is performed on the homogeneous water phantom, the sequence excitation deflection angle, echo time and repetition time parameters are (45°, TE / 2, TR) respectively, and the amplitude image and the argument image data are obtained;

[0070] Step (2) spin echo scanning is performed on the imaging human body part, the sequence excitation deflection angle, echo time and repetition time parameters are (90°, TE, TR) respectively, and the amplitude image and the argument image data are obtained;

[0071] Step (2) spin echo scanning is performed on the imaging human body part, the sequence excitation deflection angle, echo time and repetition time parameters are (45°, TE, TR) respectively, and the amplitude image and the argument image data are obtained;

[0072] The spin echo scanning is implemented on the imaged human body part, and the sequence excitation deflection angle, echo time and repetition time parameters are (45°, TE / 2, TR) respectively to obtain amplitude image and argument image data.

[0073] In each spin echo scanning above, the field of view, layer thickness, layer spacing, frequency encoding number, phase encoding step and layer number are the same and consistent;

[0074] The step (1.2) processes the homogeneous water phantom and the imaged human body part scanning data to obtain the radio frequency receiving magnetic field sensitive field distribution data and the transmitting magnetic field sensitive field distribution data of the human tissue in sequence;

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

[0076] The step (2) inverses the conductivity and the permittivity from the magnetic field measurement data:

[0077] Assuming that the magnetic permeability of the human tissue is constant and equal to the vacuum magnetic permeability, the radio frequency receiving magnetic field or the transmitting magnetic field data of the human body is used to inverse the conductivity and the relative permittivity distribution in the tissue by using the magnetic resonance electrical property imaging method;

[0078] The step (3) inverses the magnetic permeability or the magnetic susceptibility from the magnetic field measurement data and the existing electrical property:

[0079] The magnetic permeability of the human tissue is recovered to change with space, the radio frequency receiving magnetic field or the transmitting magnetic field of the human body is combined with the obtained electrical property data to inverse the distribution of the magnetic permeability in the tissue based on the magnetic field partial differential equation.

[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 the condition for the generation of the magnetic resonance signal, the applied magnetic field gradient encodes the distribution of the magnetization intensity vector in the wave number domain in frequency and phase, and finally the amplitude and phase distribution of the magnetic vector in the spatial domain are obtained through the magnetic resonance imaging Fourier inverse transform.

[0081] The present application reconstructs the image of the equivalent magnetic vector distribution signal based on the spin echo and fast spin echo scanning sequence, i.e. the equivalent spin density distribution information, and realizes the manipulation of the magnetic vector amplitude ratio and phase difference by combining the sequence parameter flip angle, echo time and repetition time control. After the parameterized image data is processed, the radio frequency receiving magnetic field and transmitting magnetic field distribution data are obtained.

[0082] Specifically, taking fast spin echo imaging as an example, in the rotating coordinate system, the magnetic vector system makes right-handed precession, assuming that the initial free induction decay signal after a radio frequency excitation pulse with a 90-degree deflection angle and a 0-degree phase is at t=0 moment, at this time, the magnetic vectors all point to the positive direction of the y' axis, i.e., are in the same phase. After a half echo interval TS / 2, the magnetic vector system has a certain degree of dephasing, assuming that the magnetic vector with an average speed precesses along the -x' axis direction at t=TS / 2, at this time, a 180-degree deflection angle refocusing pulse is applied, the magnetic vector system is changed to point to the x' axis direction, and the average speed magnetic vector therein is along the x' axis direction, and other magnetic vectors are arranged in reverse order according to the precession speed, i.e., are subject to the spin echo refocusing pulse principle. After continuing for TS / 2, the magnetic vector system again has phase coherence (the first time is at the end of the 90-degree pulse), and the echo signal reaches the peak value, at this time, the whole magnetic vector system again points to the positive direction of the y' axis. If the echo chain length is n, it means that n times of magnetic vector system refocusing will occur, i.e., n times of echo signals with different amplitudes will be generated.

[0083] According to the common setting method of the echo time TE of the fast spin echo scanning sequence, the present application takes an even number n, sets TE=TS*n / 2, and fills the sampling data of the k=n / 2th echo to the central position of the k-space. In the above description example, the multi-echo generation process of the magnetic vector system initially along the positive direction of the y' axis, dephased at the -x' axis at TS / 2, and again along the positive direction of the y' axis after TS / 2, the average speed magnetic vector does not necessarily need to turn through 90 degrees, and the multi-dephasing and refocusing condition after a certain specific rotation angle value is also applicable to the magnetic field measurement of the present application. In the process of each dephasing 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 received magnetic field measurement object, rather than the total phase accumulated in a sequence cycle by the rotation of the transverse magnetic vector in the complex plane. The phase difference between the transverse magnetic vectors corresponding to the maximum degree of dephasing at the half echo time each time is the opposite number of the phase difference at the beginning of the signal collection during the phase coherence period, which is the initial phase of the equivalent proton density physical quantity in the magnetic resonance imaging. The present application further derives the phases of the received magnetic field and the transmitted 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 first do phase unwrapping after obtaining the spatial distribution of the equivalent proton density initial phase, and then implement the calculation and processing of the phases of the received magnetic field and the transmitted magnetic field.

[0084] The parameter group (α, TE, TR) of the scanning pulse sequence implemented in one imaging examination controls the precession process of the magnetic vector system, and the imaging scanning based on the spin echo sequence also needs to consider the influence of the radio frequency refocusing pulse on the system precession. Referring to the change rule of the magnetic vector amplitude and phase after the radio frequency excitation pulse acts on the system given in 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 vector is:

[0085] Mrf,⊥ = M i,⊥ e iπ (19)

[0086] In the formula, the subscript 'rf' represents refocusing, and 'i' represents initial before inversion. That is, for the transverse component of the magnetic vector, there is no change in amplitude, but there is an inverse order of equal phase difference between the precession fast and slow magnetic vectors; for the longitudinal component, the 180° refocusing pulse does not change its amplitude. Therefore, unlike the effect of the excitation pulse deflection on the initial transverse component to produce coherence, the use of the refocusing pulse of the non-inclined inversion plane can satisfy the conditions of the magnetic vector precession process "conformal amplitude ratio" and "amplitude consistent angle constant difference" required by the RF magnetic field measurement scanning sequence of the present application. In addition, under the condition of constant phase encoding and frequency encoding gradient polarity, the formation mechanism of the echo signal determines that during each 180° refocusing pulse, it is a strictly symmetrical process of the specific fast precession magnetic vector initially lagging behind the specific slow precession magnetic vector, being coherent, and then leading the specific slow precession magnetic vector, which has the advantage that the equivalent spin density amplitude and phase distribution based on voxel measurement obtained by the reconstructed magnetic vector image after Fourier inversion has a high degree of consistency between different sequences and the same sequence scanning, simplifying the parameter-dependent Bloch equation describing the precession of the magnetic vector.

[0087] For a normal or fast spin echo sequence, due to the limited repetition time TR, the target human tissue longitudinal relaxation time T1 in general imaging cannot reach more than 4 to 5 times, so the steady-state longitudinal magnetization vector Mss will reach saturation after a number of equal-interval radio frequency excitation pulses, that is, less than the maximum environmental thermal steady-state longitudinal magnetization M0 before the start of the scan. In order to adapt to the conditions of measuring the magnetic field amplitude and phase distribution of the present application, 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 period is minimized. It is known that in the brain, the T1 and T2 reference values of cerebrospinal fluid, gray matter and white matter are 3050 ms and 323 ms, 1300 ms and 99 ms, and 830 ms and 82 ms, respectively, so the sequence TR can be set to 2000 ms, although it cannot completely restore the longitudinal magnetic vector, but it can achieve the effect of reaching a new steady state with uniform amplitude of the longitudinal magnetic vector after a number of excitations, thereby ensuring the consistency of the precession longitudinal magnetic vector at the beginning and end during the TR period, to effectively measure the magnetic field distribution. Regarding the echo time TE of the sequence, it is not necessary to make the coherence between the tissue transverse magnetic vectors fully decay before the refocusing pulse is implemented, that is, TE≥T2*, T2* is the magnetic sensitive transverse relaxation time, which is shorter than the inherent spin-spin transverse relaxation time T2; because the first and last signals of adjacent echoes before and after the TE time in a period overlap, it does not affect the amplitude consistency between consecutive echoes by keeping the phase difference unchanged after the refocusing pulse inverts the transverse vector. Figure 32The figure shows a schematic diagram of the formation of new steady states of longitudinal and transverse vectors under equally spaced radiofrequency excitation. In Figure 32 Figure (a), four radiofrequency excitation pulses are continuously applied at equal time intervals TR. Since TR is shorter than five 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 magnitude of the thermal equilibrium steady state M0. For the second excitation and so on, the longitudinal component deflected after the same TR time has a lower magnitude 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 Figure 32 Figure (b), the formation of residual transverse steady state free precession during the pulse sequence scan is described. The same four equally spaced pulses act on the target object. Since the subsequent excitations are not carried out until the transverse magnetic vector is fully coherent, 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 imaging (Magnetic Resonance Electromagnetic Properties Tomography, MREMPT) of the present invention, namely the radiofrequency excitation magnetic field and the receiving magnetic field, and their 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 detector, and digital filter. Although there are many links, their core purpose is to perform spectral shift, remove the carrier frequency, and demodulate 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 deviation of the precession resonance frequency of the nuclear spin system, that is, the frequency offset for voxel spatial encoding in imaging, is the initial phase of the lumped or voxel contribution signal, that is, in Equation (8). After demodulation, the in-phase and quadrature components S I = Acos(δωt + φ) and S Q= - Asin(δωt + φ). Because the physical meaning of the magnitude and phase data information of the magnetic resonance image obtained by the inverse Fourier transform is the amplitude and phase of the transverse component of the rotating precessing magnetization vector, the magnitude and phase data in the image file can be used to process the distribution of the RF transmit and receive magnetic field measurements. Figure 31 A typical functional diagram of the cascade units of the receiver of a magnetic resonance imaging system. The raw magnetic resonance signal from the detection coil is first amplified by a preamplifier to improve the signal-to-noise ratio in the resonance frequency band, then adjusted by a variable gain amplifier to match the dynamic range of the signal to the range of the analog-to-digital converter, and then passed through a bandpass filter to further suppress the noise outside the frequency band, and to prepare for the mixing anti-aliasing in the digital domain. After the magnetic resonance signal is filtered by the bandpass filter, it is converted into a 14-bit digital signal by the analog-to-digital converter, and the signal is split into two paths to implement quadrature phase detection, and the local oscillator signals used are cos(2πfIFt) and sin(2πfIFt), respectively. The two signals after detection are filtered by a digital low-pass filter, and finally the real and imaginary parts of the baseband k-space data are obtained, and the data after inverse Fourier transform obtains the amplitude and argument distribution of the transverse component of the precessing magnetic vector in the image domain, i.e. the original magnetic resonance measurement equivalent proton density image data. Based on the proton density data, the amplitude and phase distribution of the RF transmit and receive magnetic fields can be calculated.

[0091] Based on the fast spin echo signal acquisition scan pulse sequence, the human tissue magnetic resonance imaging is performed to measure the spatial and temporal evolution of the magnetization vector induced by the detection coil to generate a received magnetic field, and then based on the received magnetic field, the distribution of the transmit magnetic field is calculated, for example, to perform brain, breast or prostate imaging. To measure the magnetic field distribution and correct the amplitude and phase image distortion caused by the non-uniformity of the coil magnetic field in the tissue signal, a homogeneous water phantom is scanned with the same sequence parameters to obtain the corresponding signal amplitude and phase image data. Specifically, the human tissue is scanned three times with pulse sequence parameters set to (90°, TE, TR), (45°, TE, TR) and (45°, TE / 2, TR) to obtain amplitude and argument image data. The water phantom is scanned in the same way to obtain three sets of reference amplitude and argument image data, which are collectively listed in Table 1 of the magnetic resonance imaging acquisition raw image data.

[0092] Table 1

[0093] Sequence Amplitude data Phase data Tissue 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 phantom 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 water phantom.

[0095] For ease of processing the magnetic field quantity data, formula (8) is rewritten as follows:

[0096]

[0097] where for M0 according to its definition, the present invention assumes that M0 is determined only by Ns(r) in the measured magnetic field and inversion properties, i.e. that the main magnetic field distribution is uniform B0(r) = B0 and the temperature distribution of the imaged body is uniform T0 before the measured magnetic field. s (r) = T s Considering the unavoidable influence of tissue contrast, sequence time parameters, etc. on imaging in actual scanning, the measured data of the magnetization field, i.e. the magnetic resonance signal intensity SI, can be written in the following form according to the Bloch equation for a spin echo sequence SE:

[0098]

[0099] In the formula, the parameters related to the properties of the voxel are distinguished as the magnetization vector, the excitation deflection angle and the refocusing flip angle, the longitudinal and transverse relaxation times, and the receiving sensitivity field S(r) i.e. the 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 two sequence parameters, in particular TE, TR and β, are given to be equal or consistent with each other, the molecular term and The factor term can be eliminated by dividing the image data, like the same weight M0(r) term, for example, in the present invention β≡180°. However, the nonlinear coupling relationship exists in the denominator term In order for the magnetic field measurement method of the present invention to be effective and feasible, TR>>T1 is required so that the denominator is always approximately equal to 1. Taking the cerebrospinal fluid tissue with the most serious potential error T1=3050ms at 3.0T as an example, when TR=10000, If α takes 60° or 120°, the corresponding term value is further halved, and C(r) is approximated to:

[0102]

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

[0104] It should be noted that the deflection angle, echo time and repetition time in the table 1 are set to have the following meanings: in a synchronous cycle of two different deflection angles, same TE and TR sequences, the magnetic vector transverse component in the slice keeps the same angle and different amplitude, which can be divided by the difference; in a cycle of two same angles, doubled TE and same TR sequences, the transverse component keeps the same angle difference, and the target received magnetic field phase is obtained after the simple multiplication of the single phase difference. In the sequences, for the T2 weighted imaging magnetic field measurement sequence, the same TR is set to ensure that the quasi-approximate longitudinal steady state of the magnetic vector system is the same at the beginning of each cycle, regardless of the sequence flip angle, i.e. the longitudinal magnetization is saturated, and the transverse coherent magnetization in the quasi-approximate steady state can naturally decay to almost nothing. Except for 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 has low signal-to-noise ratio, and the original high-noise pixel result in the tissue region becomes an uncertain value after the received magnetic field division operation. Using the processing method of "amplitude ratio and phase difference", a set of received magnetic field and emitted magnetic field distribution measurement results can be obtained for a set of deflection angle, echo time and repetition time parameter settings combined with the homogeneous water phantom measurement. By changing the sequence parameters, a plurality of sets of magnetic field distribution measurement data of the same imaging body can be obtained, which are mathematically different measurement magnetic fields but correspond to the same electromagnetic characteristic distribution.

[0105] In the process of performing magnetic resonance imaging scanning to measure the magnetic field to collect data, the special feature of the fast spin echo sequence compared to the ordinary spin echo sequence is to consider the number of echo times and the position of the selected echo time in the echo sequence, because the k-space filling of the line group type echo signal sampling data is in order, thereby affecting the accuracy and consistency of the magnetic field phase measurement. In an embodiment of the present application, when scanning the image with the fast spin echo sequence, the echo time number of sequences II and V is set to 1, the echo chain length is 32, and the echo time is 120 ms. The echo time number of sequences III and VI is also set to 1, and the echo chain length is still 32, but the echo time is set to 60 ms. In this way, the echo time of sequences III and VI is half of that of II and V, although the echo amplitude of sequences III and VI is higher than that of II and V, but because the deflection angles are the same, the two sequences follow the same specific e exponential T2 decay law, so that the measurement and calculation of the received and emitted magnetic fields have the prerequisite of "amplitude consistency and constant phase difference".

[0106] Regarding the setting of the receive bandwidth in the sequence parameters of the magnetic resonance imaging scan, the receive bandwidth is related to the frequency encoding steps. In one embodiment of the present application, the sampling frequency encoding steps are 128, if the total sampling time of the complex samples is 5.12 ms in the readout gradient time, then the single sample time is 40 μs, the total receive bandwidth is 25 kHz, and the bandwidth given to each voxel is about 195 Hz; while in the k-space domain, the time domain frequency encoding steps 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 256 mm, then the pixel width is 2 mm, and the corresponding k-domain field of view k PE_FOV equals 0.5 mm -1 , and Δk PE = 1 / 256 mm -1 equals. It should be pointed out that the setting of the image domain field of view size does not affect the measurement of the magnetic field, and because the receive bandwidth affects the signal-to-noise ratio of the pixel signal acquisition, the Bandwidth needs to be adjusted to ensure that the Rel.SNR of different sequences is consistent, which is beneficial to the calculation of the magnetic field amplitude under the same conditions.

[0107] Regarding the use of spatially encoded magnetic field gradients, as with the above-mentioned image geometry setting parameters, only the size of the field of view and the resolution are affected, and the measurement of the magnetic field based on the pixel signal is not affected. In the two-dimensional frequency encoding slice selection gradient mode, the slice thickness is proportional to the derivative of the product of the gradient and the pulse width, and the spatial resolution in the frequency encoding readout gradient direction corresponding to a single scan slice and the cyclically executed phase encoding gradient direction is also proportional to the inverse of the product of the gradient and the acquisition time (for the 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 application has nothing to do with the setting of the main static magnetic field and the magnetic field gradient, and the resulting specific size and resolution of the image spatial characteristics.

[0108] According to Faraday's law of electromagnetic induction, a time-varying magnetic field produces a spatial gradient distribution of eddy current electric field, which is a different kind of electromotive force from the Coulomb electric field, and Lenz's law shows that the change in the resistance to magnetic flux density is the source of the gradient pulse execution period coil collision noise. Due to the permittivity of human tissues, the rising edge and falling edge of the magnetic field gradient pulse will also induce eddy currents, which will present as spatial domain second harmonic interference in the phase image. The correction method can be to perform a set of scan with changed slice, phase and frequency encoding gradient polarities on the basis of sequences I to VI, and before calculating the received magnetic field phase, the phase time of each sequence normal scan and reverse polarity scan 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 slice selection gradient, the existing 2D and 3D imaging modes of magnetic resonance imaging provide good methods for the radio frequency magnetic field measurement of the magnetic resonance electromagnetic property imaging of the present application. Based on the frequency encoding slice selection gradient and the selective excitation radio frequency pulse, the obtained tomographic voxel array original equivalent spin density measurement data have no certain correlation in the layer direction, and therefore the corresponding calculated receiving magnetic field and transmitting magnetic field are only applicable to the existing various 2D approximate electric property imaging inversion algorithms and the electromagnetic property imaging algorithms of the present application. Based on the volume imaging mode of the magnetic resonance phase encoding slice selection gradient and the non-selective radio frequency pulse, the 3D pixel matrix spin density measurement data obtained after reconstruction have natural correlation between layers, and the correlation is also transferred to the calculated receiving magnetic field and transmitting magnetic field distribution data, so that the electromagnetic property imaging inversion method which is more in line with the general physical meaning can be applied.

[0110] At present, the 2D imaging modes of the medical magnetic resonance equipment mainly include gradient echo, ordinary spin echo, echo planar and fast spin echo technologies, and the 3D imaging modes include fast spin echo and gradient echo technologies. In the 2D imaging technology, the scanning time of the ordinary spin echo sequence is too long. Although the echo planar imaging sequence provided by the manufacturer has the so-called spin echo based sequence and the phase encoding gradient polarity modification function, the spin echo-echo planar imaging is essentially based on the gradient echo, and if the excitation number is not 1, the serious band data combination artifact on the reconstructed radial image related to the excitation number. Under the condition of no gradient polarity modification function provided by the manufacturer, although the fast spin echo sequence inevitably has eddy current artifact on the scanning radial image, by setting a lower sequence parameter receiving bandwidth, a smaller readout gradient strength can be used, which can compensate for the lack of eddy current interference to a certain extent. Therefore, the present application preferably uses the fast spin echo sequence technology as the imaging technology for scanning the equivalent spin density distribution. In the 3D fast spin echo technology, there are various 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 because the radio frequency deflection angle in the pulse period has potential changes, it is not suitable for the imaging of the present application. For the 3D scanning which maintains the correlation of the measurement data between layers, the present application preferably uses the basic fast spin echo sequence to measure the equivalent spin density and the magnetic field distribution.

[0111] In the following data processing of the received magnetic field, the relative reasonable assumption of the tissue magnetic permeability is equal to the standard homogeneous water phantom solution magnetic permeability is made to meet the condition that the transmitted magnetic field sine function value can be approximated. Therefore, in order to maintain consistency, in the electromagnetic characteristic inversion process, the electrical characteristic parameters of the assumed constant magnetic permeability can be inverted by the received magnetic field distribution measurement, and the magnetic characteristics can be inverted by the transmitted magnetic field. In addition, in the measurement of the magnetic field data generated by the transmitting coil and the magnetic field induced by the receiving coil, there are two different sources of differentiation contributions, one is the natural distribution of the electromagnetic characteristics of the human body itself, and the other is the non-uniform distribution of the magnetic field generated by the non-ideal coil. Respectively, B o (r) and B c (r) are recorded, where o represents object, i.e. the subject, and c represents coil, i.e. the coil.

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

[0113] 1. Received magnetic field measurement data calculation method

[0114] (1) Amplitude component

[0115] First, the calculation process of the sequence I and II data of the tissue to obtain its received magnetic field sensitive field distribution is as follows, and the equivalent spin density equal to M0(r) is obtained:

[0116]

[0117] In the formula, r is the voxel position coordinate, γ is the gyromagnetic ratio, the superscript '90' represents the excitation deflection angle, the subscript 'h' represents the human body tissue, μ h is the magnetic permeability of the human body tissue, τ is the radio frequency excitation pulse width, the upper left superscript 'TE' represents is the received magnetic field under a particular TE setting, the top superscript'' represents the alternating quantity, the right superscript '+' and '-' represent transmission and reception respectively, and the right subscript '1' of the symbol B represents the transmitted magnetic field, which is different from the subscript '0' in B0, which represents the main magnetic field. In theory, the amplitude spatial distribution of the radio frequency excitation coil transmitted magnetic field is uniform, i.e. The same as The received magnetic field The subscript '_' represents a complex number, which is also The reason for not using the top mark with a wavy line. In addition, the excitation radio frequency pulse in the sequence of the present application takes the same pulse width, and the deflection angle is modulated by changing the pulse amplitude, which ensures that the spin system makes synchronous precession during the scanning process of different α, same TE and TR sequences, that is, it ensures that the precession phase is always equal, which is a prerequisite.

[0118] Magnetic resonance imaging requires Uniform distribution, according to B(r) = μ(r)H(r), its non-uniformity from the possible μ(r) and H(r) two parts. The following will be based on this relationship from the transmit magnetic field inversion of magnetic permeability. For further processing to obtain the received magnetic field sensitive field distribution The following derivation maintains Expanded as μ and H two parts.

[0119] Then, also in the measured sequence IV parameters (90°, TE, TR) and sequence V parameters (45°, TE, TR) scan homogeneous water phantom, processing to obtain the same conditions under the unit received magnetic field distribution expression containing the specific trigonometric function weight of the transmit magnetic field:

[0120]

[0121] In the formula, μ p,0 The magnetic permeability of the phantom water phantom, subscript 'p' represents the water phantom, subscript '0' represents the magnetic permeability constant, because the water phantom is homogeneous Only related to the received coil magnetic field c non-uniformity, c represents the coil coil. By using the characteristics of the water phantom, the transmit magnetic field related specific trigonometric function weight image distribution contained in the tissue scanning data is obtained.

[0122] In order to further eliminate the excitation coil transmit magnetic field non-uniformity contained, and the received magnetic field non-uniformity related to the detection coil, the ratio of the equivalent proton density distribution of the imaged human tissue and the phantom is:

[0123]

[0124] When considering that the magnetic permeability μ h (r) of human tissue is approximately equal to the magnetic permeability μ p,0 of the homogeneous content of the phantom, the contribution of the transmit coil magnetic field non-uniformity in the above formula can be eliminated, and the received magnetic field sensitive field distribution under the scanning condition of specific sequence parameters (TE, TR) is:

[0125]

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

[0127] (2) Phase component

[0128] First, the precession process of a single spin in the tissue or water phantom is analyzed, especially how the received magnetic field is formed during the spin echo sequence imaging process The magnetization vector in the received magnetic field precesses at the Larmor frequency, and the signal produces an initial transverse magnetization component M 0,⊥ The initial phase of The initial phase of the excitation transmit magnetic field is π / 2 different from the initial phase , and then a free induction decay signal is generated. During the scanning process, the echo signal formed each time contains spatial encoding, which is the object of signal acquisition by the detection coil, i.e. the received magnetic field of the application. It should be noted that the instantaneous phase during the echo formation, including the front and rear coherent and incoherent phases, is the relative phase distribution displayed by the image domain phase image obtained by the inverse Fourier transform. Although the relative phase may sometimes exceed the 2π range due to the long acquisition time, it needs to be unwrapped, but the imaging measurement phase has nothing to do with how much phase a single spin precesses during the entire TR period, nor does it have anything to do with the phase at nω0TE.

[0129] In order to measure the received magnetic field phase distribution when imaging the human body, the contribution of the transmit magnetic field phase part needs to be eliminated from the equivalent spin density phase distribution in the original magnetic resonance signal data. Similar to the measurement method of the radio frequency magnetic field amplitude, two scans are performed on the same imaging body, i.e. involving main sequences II and V and auxiliary sequences III and VI. Among them, the same excitation deflection angle of 45° is implemented to obtain the same M 0,⊥ The initial phase of The echo times of the two sequences with the same echo signal type are proportionally proportional during the corresponding cyclic execution cycle. Since the transverse relaxation time constant of the voxel determines the relaxation rate, and the precession angular frequency is the same under the same gradient geometry, the two scans have a constant phase difference related to TE at the same position. The same repetition time TR for the two sequences ensures that the magnetization vector tends to end with a decrease in the transverse component in the current sequence cycle and a steady state in the recovery of the longitudinal component in the next sequence cycle, with nearly the same amplitude and azimuth angle. In this invention, according to the sequence implementation scheme in Table 1, the goal of phase correlation data processing is to obtain the phase distribution of the radio frequency magnetic field of the human body scanned by the master parameter sequence (45, TE, TR). The specific data processing process is as follows: The phase difference is obtained by subtracting the argument data of sequence III from the argument data of sequence II:

[0130]

[0131] In the formula, and These represent the phases of the receiving and transmitting magnetic fields, 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. This difference removes the constant of the transmitting magnetic field phase distribution and includes information about the non-homogeneity of the test subject and the non-homogeneity of the receiving coil's magnetic field.

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

[0133]

[0134] In the formula, the subscript p denotes the phantom of the water model. Here, because Δp p The value only reflects the non-uniformity of the phase distribution of the magnetic field in the receiving coil, but this distribution also exists in the phase difference of the human body, so it can be used to measure Δp. h Further revisions are made, including:

[0135]

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

[0137]

[0138] It is worth mentioning that, compared with the received magnetic field sensitive field derivation calculation, the nonlinear sin function acting on the transmitted magnetic field is assumed to be homogeneous and water model of human body magnetic permeability, and the calculation of the single received magnetic field phase quantity Δp of human tissue h The calculation of the single received magnetic field phase quantity Δp of human tissue Irrelevant to the magnetic properties of the tissue, and then do the relevant subtraction operation.

[0139] 2, the method for calculating the transmitted magnetic field measurement data

[0140] (1) Amplitude component

[0141] Still use the amplitude measurement main sequence II and auxiliary sequence I magnetic resonance signal, that is, the equivalent proton density, and the calculated received magnetic field sensitive field data corresponding to the unit current generated in the detection coil Under the sequence time parameter (TE, TR) condition, the transmitted magnetic field sensitive field distribution data of the imaged human tissue can be calculated by the following steps. According to the formula (6) of the macroscopic magnetization vector, define two groups of ratio type intermediate quantities:

[0142]

[0143]

[0144] In the formula, M0 is the amplitude of the macroscopic magnetization vector of the imaged human tissue, and the superscript '90' represents the sequence excitation deflection angle, and 'TE' represents the field B measured at a certain echo time TE.

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

[0146]

[0147] Compared with formula (9), further can obtain:

[0148]

[0149] In the formula, t1 and t2 are intermediate quantities of formulas (33) and (34). It can be seen that, similar to the existing double-angle method, the present application also uses the inverse cosine function to solve the transmitted magnetic field sensitive field distribution. And in order to meet the relationship of the trigonometric identity sin 2θ = 2sinθcosθ, the sensitive field measurement main sequence and auxiliary sequence deflection angle are specially taken as typical 90° and 45° respectively.

[0150] Regarding the excitation radio frequency pulse width τ, the magnetic resonance vendors usually take the way of changing the transmit coil driving current amplitude, instead of the pulse width, to control the deflection angle size, so as to facilitate the calculation and adjustment of the echo time and repetition time to keep the consistency between different sequences when performing the multi-echo scan sequence. In practice, considering the saturation effect of the magnetic moment, the current amplitude ratio of 2 times the deflection angle is about 1.9, and if the transmit magnetic field sensitive field distribution data of the amplitude sequence is wanted which can be calculated by the main sequence is multiplied by 2 to obtain the approximate value. Both the imaging scan times are reduced, and the need for multiple magnetic field measurement quantities for certain electromagnetic characteristic imaging. Due to the unit constraints of the formula relationship between the radian and the gyromagnetic ratio and the pulse width magnitude, the order of magnitude of the transmit magnetic field sensitive field is in μT.

[0151] (2) Phase component

[0152] Like the calculation of the transmit magnetic field sensitive field distribution , the transmit magnetic field phase distribution also needs to use the calculated received magnetic field phase distribution data information. The received magnetic field phase distribution of the human tissue scan is calculated by using the main sequences II and V and the auxiliary sequences III and VI The measured transmit magnetic field phase distribution based on the sequence parameters (45°, TE, TR) is calculated as follows by the Fourier inverse transform related equivalent proton density formula (21):

[0153]

[0154] In the formula, the phase difference π / 2 is derived from the phase of the transverse component of the magnetic moment M 0,⊥ which is equal to the phase of the transmit magnetic field of the radio frequency excitation pulse plus a constant π / 2, which is equivalent to M0 multiplied by the imaginary unit i.

[0155] So far, using the imaging scan pulse sequence parameters (45°, TE, TR) as the radio frequency magnetic field measurement means, based on the original equivalent spin density distribution of magnetic resonance imaging to collect image data, a complete excitation transmit magnetic field and detection received magnetic field sensitive field and phase information derivation and calculation method is realized. Except for the assumption that the human tissue is a homogeneous magnetic permeability in the received magnetic field sensitive field amplitude calculation process, the formula derivation is strictly in accordance with the amplitude and phase law of electromagnetic wave propagation in the calculation of the transmit magnetic field sensitive field and the transmit and receive phase. And in the calculation process of using the measurement data of the homogeneous water phantom to assist the human tissue magnetic field measurement, the effect of eliminating the magnetic field inhomogeneity of the receive coil and the transmit coil is additionally obtained.

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

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

[0158] Due to the limitation of the independent variable of the inverse cosine function being in the range (0,1), beg The dynamic range is relatively small; and there is a trend that the smaller the target single flip angle, the smaller the dynamic range. Therefore, in actual imaging scanning to measure the magnetic field, according to the value rules of trigonometric functions, 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 in a normal spin echo (SE) sequence, the formula (22) for the signal intensity generated at a specific position r in the magnetic resonance receiving coil is rewritten according to the Bloch equation describing the spin precession process:

[0161]

[0162] In the formula, M0 is the initial magnetization intensity in the ambient 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, i.e., the unit receiving magnetic field amplitude of the present invention. From the formula, it can be calculated that when α is equal to 120° and 60° respectively and β is equal to 180°, the intensity SI is respectively 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 the same, although implementing double excitation deflection angle exceeds the transverse plane, there is a certain power waste from the perspective of signal-to-noise ratio. The excitation schemes with deflection angles of 120° and 60° in the sequence group produce the same amount of initial transverse magnetization component. However, from the comparison of signal strength, the latter is greater than the former. That is, the relationship that the intensity of double angle is greater than that of single angle continues in the 90° and 45° schemes.

[0166] 3. Magnetic Resonance Electromagnetic Property Imaging Inversion Method

[0167] While the Gauss law and the Gauss magnetic law describe the interaction of matter with the electric field or the magnetic field of a single field source, the Faraday's law of electromagnetic induction and the Maxwell-Ampere law establish the interaction of matter with electromagnetic fields, including electromagnetic waves. In order to use the radio frequency magnetic field distribution measurement quantity under the magnetic resonance condition to inverse the electrical characteristic and magnetic characteristic parameters of matter, the present application starts from the Faraday's law and the Maxwell-Ampere law containing electric field and magnetic field, establishes the core calculation equation to solve the electromagnetic characteristic from the known magnetic field quantity, and the electric field quantity is expressed by the magnetic field quantity. In addition, considering that there is one complex value magnetic flux density vector in one equation, and two scalar unknown quantities, i.e. the electrical characteristic complex permittivity and the magnetic characteristic permeability, the complex permittivity is divided into conductivity and permittivity.

[0168] Starting from the principle of magnetic resonance imaging, the present application proposes a method for measuring the radio frequency transmitting magnetic field and receiving magnetic field distribution based on the same imaging body in the two continuous links of magnetization vector excitation and echo signal detection. The two magnetic field propagation directions are opposite, and the physical meanings are also different. For imaging scanning, different measurement transmitting magnetic field and receiving magnetic field distribution can be obtained by changing the pulse sequence parameters, and different core calculation equations can be derived according to the precondition of the electromagnetic characteristic (σ, ε, μ) inversion algorithm, so as to ensure that there are enough magnetic field measurement known quantities in the case of multiple electromagnetic characteristic parameter unknown quantities, and the field quantities are linearly independent with each other and also originate from the imaging body. In addition, at the calculation execution level of the inversion algorithm, multiple sets of independent magnetic field measurement data are beneficial to reduce the occurrence of rank deficiency of the coefficient matrix. First, the related wave propagation basic equation expressed by the magnetic field under the assumption of no matter homogeneity is derived:

[0169]

[0170] In the formula, for non-homogeneous matter And μ are related to the spatial position, so Note that the field quantity H is not a known quantity, and it needs to be replaced by (B / μ) before solving

[0171] The complex value magnetic field vector B can be orthogonally decomposed into (B x ,B y ,B z ) in the rectangular coordinate system, and can also be decomposed into (B + ,B - ,B z ) in the cylindrical coordinate system, wherein B + is the right-handed component, and B - is the left-handed component. In magnetic resonance imaging, the transmitting magnetic field measurement quantity is represented as The receiving magnetic field is represented as ​The following derivation process is used to inverse the target electromagnetic characteristic parameters with the magnetic field rotation component in cylindrical coordinate system. In the component form of equation (43), the formula is written as, are equal in status.

[0172] Record is the conjugate of the received magnetic field Rewrite equation (4) and directly express the magnetic flux density in rectangular coordinate system components B x and B y represent the rotating transmitting magnetic field and the received magnetic field

[0173]

[0174] In the formula, B x and B y are the complex-valued amplitudes of the vector, i is the imaginary unit, the superscript ‘*’ represents the conjugate of the complex number, and the subscripts ‘x’ and ‘y’ represent the x and y direction components of the field, respectively. Regardless of the magnetic resonance electrical characteristic imaging inversion method, such as the standard Helmholtz equation method about the magnetic field vector which ignores the electrical characteristic gradient, the corresponding scalar core equation related to the rotation component needs to be derived, so that the unknown electrical characteristic parameter distribution can be solved with the known component sensitivity field amplitude or phase. Continue to take the Helmholtz equation as an example to derive the equation in the form of the rotating coordinate system component from the orthogonal field components in the rectangular coordinate system. The derivation is as follows:

[0175]

[0176] Linear combination of the equations has:

[0177]

[0178] Further, according to the definition expression of the magnetic resonance transmitting and receiving magnetic field:

[0179]

[0180] In the formula, the complex permittivity is the target parameter, which can be obtained by or Based on the voxel data matrix, it can be inverted by the least square method.

[0181] In order to reduce the difficulty of solving equations and obtain relatively accurate electromagnetic characteristics, two different technical routes are used to implement the inversion of complex permittivity and permeability in combination with necessary assumptions. In the existing magnetic resonance electrical property imaging, the assumption of permeability close to vacuum is used to derive the conductivity and permittivity from the transmitted magnetic field distribution. As described above, the mainstream three inversion methods, Helmholtz equation method has reconstruction system error at the tissue interface due to ignoring the gradient of electrical properties; the gradient EPT method improves the accuracy of the results by introducing the intermediate calculation process of the property gradient, but the field measurement method has problems; in the algorithm without ignoring the gradient of electrical properties, the convection-reaction EPT method derives linear equations in combination with necessary Dirichlet boundary conditions under the condition of 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 application is based on the existing crEPT, which first calculates the complex permittivity parameter distribution of the imaged human tissue from the transmitted magnetic field distribution data, and then completes the inversion of all electromagnetic properties from the received magnetic field data measured in the same scan according to another crEPT equation about permeability under the condition that the complex permittivity is known.

[0183] The second type of method of the application considers that adding an unknown quantity in the partial differential equation in the complex domain of the vector necessarily increases the difficulty of the solving process, therefore, in combination with necessary several homogeneous assumptions of electrical properties and magnetic properties, a simple Helmholtz type core equation is derived, which explicitly gives the relationship between the magnetic field and the electromagnetic properties, and the least square method is used to fit and solve the complex permittivity and permeability, and finally the conductivity, relative permittivity and permeability distribution based on the voxel are obtained according to the corresponding relationship between the real part and the imaginary part. The magnetic property inversion method of magnetic resonance electromagnetic property imaging is shown in Table 2.

[0184] Table 2

[0185]

[0186] (1) Core equation derivation of the first type of property inversion method

[0187] Because the measurement of the received magnetic field distribution contains the assumption that the magnetic property of the imaged human tissue is similar to the magnetic property of the homogeneous water phantom in the derivation and calculation, the transmitted magnetic field distribution data is used to solve the complex permittivity distribution based on the existing magnetic resonance electrical property imaging inversion method.

[0188] Under the condition that the conductivity and permittivity distribution and one or several groups of received magnetic field data are known, according to the Maxwell equation set, the wave equation containing the unknown quantity of permeability and only containing the magnetic field is:

[0189]

[0190] where is the complex permittivity for distinguishing from the magnetic resonance electrical property imaging. If the assumption of homogeneous magnetic permeability is not made, i.e. the magnetic permeability is not homogeneous, the derivation process is very complicated. The final result of the equation for the received magnetic field is as follows:

[0191]

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

[0193] However, if the assumption of homogeneous magnetic permeability is made, the derivation process can be greatly simplified, and the program can be written, the error rate can be reduced, and the stability of the solution can be maintained. First, if the 1st magnetic homogeneity assumption is made for the divergence term of the magnetic field strength, we have:

[0194]

[0195] The advantage of the above equation is that it leaves the unknown μ. The 2nd magnetic homogeneity assumption is made to remove μ and multiply it on the right side term, and we have:

[0196]

[0197] It should be noted that after multiplying the right side term, μ returns to the non-homogeneous property. In the Helmholtz equation method of electrical property imaging, when the equation component form is derived, attention should be paid to the difference between the received magnetic field conjugate and the transmitted magnetic field itself when the received magnetic field is used to determine the property.

[0198] (2) The core equation derivation of the second type of property inversion method

[0199] Compared with the first type of inversion method, the second type of property inversion method uses a certain number of homogeneous assumption conditions 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 the scalar complex permittivity and the magnetic permeability. Starting from the formula (43) without any approximation processing, the specific process is as follows. Make 1st human tissue magnetic property homogeneity assumption and 1st electrical property homogeneity assumption, then from the Gauss magnetic law we have and Let v = μ -1 , we get:

[0200]

[0201] In the present invention of magnetic resonance electromagnetic property imaging, the right-handed motion of the radio frequency transmit magnetic field and the left-handed motion of the receive magnetic field are the measured quantities, i.e. the known quantities in the above equation. Therefore, the above equation can be written in the orthogonal component form using equation (44) as:

[0202]

[0203] So far, the goal is to solve the unknown quantities in the above equation set Although the two equations in the equation set are in the same form, the two magnetic field measured known quantities are different, i.e. the response functions under different field excitations of the same inhomogeneous property source in the space, so that the two unknowns, i.e. the electrical and magnetic property parameter distributions, can be solved independently from the two complex known quantities. If there are multiple sets of magnetic field measured known quantities obtained from the parameter scan with variation, then the solvability of the solution can be increased.

[0204] Further examining term, it can be written as:

[0205]

[0206] At this time, the equation is suitable for applying the method for solving the convection-diffusion equation.

[0207] Based on the condition of 3.0T magnetic resonance imaging system, the present application uses fast spin echo sequence to scan the human tissue equivalent proton density image data, further calculates the distribution of radio frequency transmitting magnetic field and receiving magnetic field, thereby inverts the distribution of electrical and magnetic characteristics parameters of the tissue and reconstructs. Under the condition of 3.0T cranium brain imaging, the longitudinal relaxation time and transverse relaxation time reference values of cerebrospinal fluid, grey matter and white matter tissue are 3050ms, 1300ms, 830ms and 323ms, 99ms, 82ms respectively, in order to basically meet the measurement scanning requirements of TR>3*T1 and TR>>T2, the preferred parameters of the fast spin echo sequence used in the present application are (120°, 30ms, 9000ms), (60°, 30ms, 9000ms) and (60°, 15ms, 9000ms). Assuming that the imaging geometry parameters of frequency encoding and phase encoding direction are both 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 echo time TE=30ms, and referring to the echo train length etl 16 in conventional imaging, the calculation echo time interval is 3.75ms, and considering the magnetization field non-uniform effect, the interval can fully accommodate tacq; and the total acquisition time of 1 excitation is 60ms, i.e. 2 times of TE, referring to the conventional magnetic resonance imaging examination, the image contrast obtained by using the sequence scanning is biased to proton density weighting. Generally, for cranium 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 application (120°, 30ms, 9000ms), (60°, 30ms, 9000ms) is within 3*T2_M, M represents grey 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 is the readout gradient in the readout direction. Under the condition of bandwidth 62.5kHz and field of view 256*256mm 2 , the readout gradient strength G_RO is only 5.7mT / m.

[0210] For TE = 15 ms of (60°, 15 ms, 9000 ms), the echo time interval is reduced to 1.875 ms under the same etl condition, when Bw is unchanged, the degree of the first and last echo connection phenomenon may occur to the transverse relaxation time T2* value of the tissue depending on the magnetic susceptibility, but does not affect the refocusing pulse flip transverse magnetic vector precession phase to form the echo and complete the signal acquisition, and further calculate the radio frequency magnetic field distribution from the measurement quantity. The total acquisition time is reduced to 60 ms, and the signal-to-noise ratio is increased as a whole due to the echo chain time moving forward since the sampling time is unchanged. In order to maintain the consistency of the signal-to-noise ratio of the acquired signal, according to:

[0211]

[0212] In the formula, t represents the time domain, ω0 is the angular frequency, β is the coefficient related to the deflection angle, B 1u is the magnetic field generated by the unit current through the receiving coil, that is, the magnetic field in the present application dV s is the volume of the voxel, k is the Boltzmann constant, T is the coil temperature, R c is the resistance, B is the bandwidth, that is, Bw, N RO is the number of matrix points in the readout direction, N PE is the number of matrix points in the phase direction, N a is the number of excitations; that is, under the condition that 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 reduction of TE. Adjusting the bandwidth of the (60°, 60 ms, 9000 ms) sequence, that is, reducing td, can reduce the time domain signal-to-noise ratio. Note that when td and tacq are reduced, the frequency encoding gradient strength will automatically increase due to the unchanged voxel spatial resolution, which will bring a double superposition effect of enhanced gradient slew rate-vortex effect, especially in the calculation of precession phase, that is, the original magnetic resonance radian data used in the calculation of the emission and reception magnetic field phase of the sequences (60°, 120 ms, 9000 ms) and (60°, 60 ms, 9000 ms) are sensitive to eddy current. Assuming that Bw is increased to 62.5 kHz, td = 8 μs, tacq = 1.024 ms, and the voxel spatial resolution remains unchanged, G_RO = 11.5 mT / m; at this time, the echo interval of 1.875 ms can still accommodate tacq. In summary, in order to facilitate the accuracy of the magnetic field measurement, the preferred embodiment of the present application is to set the bandwidth to a relatively narrow value of 31.25 kHz for the sequences (120°, 30 ms, 9000 ms) and (60°, 30 ms, 9000 ms), leaving a margin for adjusting the short TE sequence to reduce the eddy current interference effect, and the bandwidth of (60°, 15 ms, 9000 ms) can be adjusted according to the actual value displayed by the system after setting the short TE to obtain the same relative signal-to-noise ratio Rel.SNR as the long TE sequence.

[0213] In the longitudinal direction, the set TR is greater than 3 times the longest longitudinal relaxation time of the tissue cerebrospinal fluid T1, which can ensure that the magnetization vector in most tissues is fully recovered, and the influence of the longitudinal partial saturation effect is basically avoided.

[0214] Beneficial effects:

[0215] The present application can measure the amplitude and phase components of the time-harmonic radio frequency transmit magnetic field excitation in nuclear magnetic resonance imaging, and the amplitude and phase components of the received magnetic field induced detection signal. The present application overcomes the problem that the existing technology cannot remove the coupling between the imaging body spin density and the amplitude of the received magnetic field sensitive field in the detection signal, and also overcomes the problem that the detection signal cannot distinguish the phase of the transmit magnetic field and the phase of the received magnetic field. In turn, the spin density distribution of the human tissue can be calculated after obtaining the amplitude of the received magnetic field, thereby having a profound technical impact on the existing transverse relaxation time weighting, longitudinal relaxation time weighting and proton density weighting imaging principle system.

[0216] Although the magnetic permeability of human tissue is relatively homogeneous and close to the magnetic permeability of vacuum, the present application breaks the unrealistic premise assumption of the magnetic permeability of tissue in vacuum in the existing magnetic resonance electrical property imaging, and proposes magnetic resonance electrical property and magnetic property imaging. In medical clinical practice, for the related pathological state of brain tissue micro-hemorrhagic mass, breast calcification and other magnetic property changes, magnetic property imaging can play an auxiliary role in imaging differential diagnosis.

[0217] In order to evaluate the local microcirculation of the human body, the existing magnetic resonance perfusion imaging needs to inject exogenous gadolinium-based contrast agent to change the magnetic properties of the target tissue and the relaxation time, thereby reflecting the image specificity. However, it may have some side effects due to incomplete metabolism and deposition in the deep brain tissue. Like arterial spin labeling imaging and blood oxygenation level dependent functional magnetic resonance imaging, the present application provides an imaging method based on endogenous magnetic property differences.

[0218] Unlike the existing susceptibility weighted imaging and quantitative susceptibility imaging, which estimate the static magnetic susceptibility distribution based on the precession phase difference, the present application of magnetic resonance radio frequency magnetic property and electrical property imaging uses spin echo sequence instead of gradient echo sequence to acquire magnetic field measurement signals, which can effectively compensate for the influence of the main static magnetic field and gradient magnetic field non-uniformity in the signal. Combined with the pre-scan homogenization 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 property imaging of the present application, the influence of the transmit and receive coil magnetic field non-uniformity on the measurement results is effectively reduced or even eliminated.

[0219] In combination with the method for measuring the distribution of the transmitting and receiving magnetic fields of the magnetic resonance provided by the application, the magnetic property parameters of the human tissue can be inversed by two technical routes. The first method is to solve the electrical property by using the existing magnetic resonance electrical property imaging gradient inversion algorithm, and then to inverse the magnetic property by using the receiving magnetic field as the measurement and the magnetic inversion core equation, wherein the magnetic inversion has different core equations according to the homogeneous condition. In the second method, the solving difficulty of the core equation is reduced by ignoring the property gradient condition, and the advantages are that the electrical property and the magnetic property can be inversed at the same time, and the disadvantages are that the property results have errors at different tissue interfaces. Compared with the first method, the second method does not assume the magnetic homogeneity, and the calculated electromagnetic property parameter distribution is closer to the real situation.

[0220] The measurement feasibility and accuracy of the radio frequency magnetic field distribution of the magnetic resonance imaging provide the possibility for the real evaluation of the specific absorption rate (SAR) related to the safety of the human body, make the quantification of various weighted imaging such as T1, T2, T1 / T2 and PD possible, and can be used for more accurately correcting the magnetic field inhomogeneity in the images of various radio frequency coils such as the transmit-receive separation, the transmit-receive integration, the phased array and the parallel imaging application. BRIEF DESCRIPTION OF DRAWINGS

[0221] The application will be further described below in combination with the drawings and examples.

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

[0223] Figure 2 Fig. 2 is a flow chart of the human magnetic resonance imaging radio frequency transmitting and receiving magnetic field measurement and the inversion of the electrical and magnetic property parameters of the tissue of the application.

[0224] The relative precession of the magnetization vector and the timing of the resonance echo signal generation related to the imaging scanning pulse sequence affect the magnetic field measurement, Figure 3 (a) is a precession process schematic diagram when the transverse component of the magnetic vector rotates with equal phase and unequal amplitude, Figure 3 (b) is a precession process schematic diagram when the transverse component of the magnetic vector has equal amplitude decay rate and constant phase difference, Figure 3 (c) is a timing relationship diagram of the echo signal acquisition in two scans with equal sequence excitation deflection angle and half echo time difference.

[0225] The relative change of the precession phase of the magnetization vector when the spin echo sequence is used to acquire the magnetic resonance signal affects the magnetic field measurement, Figure 4 (a) is when the specific echo interval time makes the transverse component of the magnetic vector refocus on the specific coordinate axis each time, Figure 4(b) is when the specific echo time interval causes the lateral component to refocus on a different coordinate semiaxis each time.

[0226] Figure 5 is a flow chart of the procedure of receiving magnetic field distribution data from magnetic resonance acquisition and reconstructed image processing.

[0227] Figure 6 is a flow chart of the procedure of transmitting magnetic field distribution data from magnetic resonance acquisition and reconstructed image and existing receiving magnetic field processing.

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

[0229] Figure 8 is the target conductivity distribution map of the water phantom simulation containing an array of abnormal bodies for verifying the inversion algorithm of electromagnetic characteristic imaging of the measured radio frequency magnetic field of magnetic resonance.

[0230] Figure 9 is the target relative permittivity distribution map of the water phantom simulation containing an array of abnormal bodies.

[0231] Figure 10 is the excitation transmitting magnetic field sensitive field distribution map obtained by inversion from the water phantom simulation containing an array of abnormal bodies.

[0232] Figure 11 is the excitation transmitting-receiving magnetic field continuous phase distribution map obtained by inversion from the water phantom simulation containing an array of abnormal bodies.

[0233] Figure 12 is the conductivity reconstruction distribution map obtained by inversion based on the Helmholtz equation method from the water phantom simulation containing an array of abnormal bodies.

[0234] Figure 13 is the relative permittivity reconstruction distribution map obtained by inversion based on the Helmholtz equation method from the water phantom simulation containing an array of abnormal bodies.

[0235] Figure 14 is the target proton density distribution map of the numerical simulation of the human head model for verifying the inversion algorithm of electromagnetic characteristic imaging of the measured radio frequency magnetic field of magnetic resonance.

[0236] Figure 15 is the target conductivity distribution map of the numerical simulation of the human head model for verifying the inversion algorithm of electromagnetic characteristic imaging of the measured radio frequency magnetic field of magnetic resonance.

[0237] Figure 16 is the target relative permittivity distribution map of the numerical simulation of the human head model for verifying the inversion algorithm of electromagnetic characteristic imaging of the measured radio frequency magnetic field of magnetic resonance.

[0238] is the simulated excitation magnetic field sensitivity map without applying magnetic resonance spatial encoding gradient magnetic field, Figure 17a is the simulated excitation magnetic field sensitivity map without applying magnetic resonance spatial encoding gradient magnetic field, Figure 17b is the corresponding phase map.

[0239] is the simulated excitation magnetic field sensitivity map without applying magnetic resonance spatial encoding gradient magnetic field, Figure 18a is the simulated excitation magnetic field sensitivity map without applying magnetic resonance spatial encoding gradient magnetic field, Figure 18b is the corresponding phase map.

[0240] Figure 19 is the reconstructed conductivity distribution map based on the Dual method with characteristic gradient term.

[0241] Figure 20 is the reconstructed relative permittivity distribution map based on the Dual method with characteristic gradient term.

[0242] Figure 21a is the original proton density weighted equivalent proton density amplitude distribution image, Figure 21b is the corresponding equivalent proton density argument distribution image (in one practical phantom measurement and imaging example of the present invention of magnetic resonance magnetic field measurement and electromagnetic property imaging, a homogeneous water phantom with abnormal shape was scanned under 1.5T magnetic resonance condition).

[0243] Figure 22a is the original proton density weighted equivalent proton density amplitude distribution image, Figure 22b is the corresponding equivalent proton density argument distribution image to investigate the influence of conductivity change on electromagnetic property imaging (in one practical phantom measurement and imaging example of the present invention of magnetic resonance magnetic field measurement and electromagnetic property imaging, a homogeneous water phantom with same abnormal shape, dissolved with 2.5g NaCl was scanned under 1.5T magnetic resonance condition).

[0244] Figure 23a is the original proton density weighted equivalent proton density amplitude distribution image, Figure 23b is the corresponding equivalent proton density argument distribution image to investigate the influence of conductivity and magnetic permeability change on electromagnetic property imaging (in one practical phantom measurement and imaging example of the present invention of magnetic resonance magnetic field measurement and electromagnetic property imaging, a homogeneous water phantom with same abnormal shape, dissolved with 2.5g NaCl and 0.25ml Gd-DTPA with concentration of 377mg / ml was scanned under 1.5T magnetic resonance condition).

[0245] Figure 24a is the original proton density weighted equivalent proton density amplitude distribution image, Figure 24bis the corresponding equivalent proton density phase angle distribution image (in a practical imaging volume measurement and imaging example of the present invention magnetic resonance magnetic field measurement and electromagnetic property imaging, human cranial tissue scanned under 1.5T magnetic resonance conditions).

[0246] Figure 25a is the measured transmit magnetic field susceptibility distribution, Figure 25b is the transmit magnetic field phase distribution, and Figure 25c is the corresponding receive magnetic field susceptibility distribution, Figure 25d is the receive magnetic field phase distribution (in the described magnetic field measurement embodiment of scanning a homogeneous water phantom containing an abnormal shape, three fast spin echo sequences (120°, 40 ms, 10000 ms), (60°, 40 ms, 10000 ms), and (60°, 20 ms, 10000 ms) were utilized for proton density weighted imaging).

[0247] Figure 26a is the measured transmit magnetic field susceptibility distribution, Figure 26b is the transmit magnetic field phase distribution, and Figure 26c is the corresponding receive magnetic field susceptibility distribution, Figure 26d is the receive magnetic field phase distribution (in the described magnetic field measurement embodiment of scanning a homogeneous water phantom containing the same abnormal shape, dissolved with 2.5 g NaCl, three fast spin echo sequences (120°, 40 ms, 10000 ms), (60°, 40 ms, 10000 ms), and (60°, 20 ms, 10000 ms) were utilized for proton density weighted imaging).

[0248] Figure 27a is the measured transmit magnetic field susceptibility distribution, Figure 27b is the transmit magnetic field phase distribution, and Figure 27c is the corresponding receive magnetic field susceptibility distribution, Figure 27d is the receive magnetic field phase distribution (in the described magnetic field measurement embodiment of scanning a homogeneous water phantom containing the same abnormal shape, dissolved with 2.5 g NaCl and 0.25 ml of gadopentetic acid meglumine at a concentration of 377 mg / ml, three fast spin echo sequences (120°, 40 ms, 10000 ms), (60°, 40 ms, 10000 ms), and (60°, 20 ms, 10000 ms) were utilized for proton density weighted imaging).

[0249] Figure 28a is the measured transmit magnetic field susceptibility distribution, Figure 28b is the transmit magnetic field phase distribution, and Figure 28c is the corresponding receive magnetic field susceptibility distribution, Figure 28dThe phase distribution of the received magnetic field (in the described embodiment of the measurement of the magnetic field of the scanned human brain tissue, the 3 fast spin echo sequences of proton density weighting (120°, 40 ms, 10000 ms), (60°, 40 ms, 10000 ms) and (60°, 20 ms, 10000 ms) are used).

[0250] Figure 29a is the conductivity distribution map of the abnormal shape homogeneous water phantom obtained by the inversion of the calculated amount of the transmitted magnetic field, Figure 29b is the permittivity distribution map of the abnormal shape homogeneous water phantom obtained by the inversion of the calculated amount of the transmitted magnetic field, Figure 29c is the conductivity distribution map of the abnormal shape homogeneous water phantom obtained by the inversion of the calculated amount of the received magnetic field, Figure 29d is the permittivity distribution map of the abnormal shape homogeneous water phantom obtained by the inversion of the calculated amount of the received magnetic field.

[0251] Figure 30a is the tissue conductivity distribution map obtained by the inversion of the calculated amount of the transmitted magnetic field, Figure 30b is the tissue permittivity distribution map obtained by the inversion of the calculated amount of the transmitted magnetic field, Figure 30c is the tissue conductivity distribution map obtained by the inversion of the calculated amount of the received magnetic field, Figure 30d is the tissue permittivity distribution map obtained by the inversion of the calculated amount of the received magnetic field.

[0252] Figure 31 is a schematic diagram of the typical cascade unit function of the receiver of the magnetic resonance imaging system.

[0253] Figure 32 is a schematic diagram of the formation of the longitudinal magnetization vector and the transverse magnetization vector, wherein Fig. (a) is a schematic diagram of the formation of the longitudinal magnetization, showing that the longitudinal magnetization vector Mss is less than the initial magnetization M0 in the environment at thermal equilibrium; Fig. (b) is a schematic diagram of the formation of the transverse magnetization vector, showing that the transverse magnetization vector M’ss is greater than 0 during the period (the new steady-state magnetization in the imaging body will be formed under the excitation of multiple equal-interval radio frequency pulses, which is different from the pre-scan).

[0254] In the figure: 1. The main magnet for nuclear magnetic resonance imaging, 2. The imaging body, 3. The excitation radio frequency transmitting coil, 4. The detection of the received signal receiving coil. DETAILED DESCRIPTION

[0255] The present application will be described with respect to the drawings in which the various embodiments of the application are illustrated. The detailed description set forth below in connection with the appended drawings is intended as a description of various embodiments of the present application and is not intended to represent the only embodiments in which the present application can be practiced. Each embodiment described represents only a typical embodiment and is not meant to exclude other embodiments from the scope of the present application. The following detailed description is not meant to limit the present application to the embodiments described, as alternatives will become apparent to those of ordinary skill in the art. For the purposes of clarity, not every feature of the embodiments described can be shown in each of the figures and some of the figures can be simplified for ease of illustration.

[0256] In addition, the descriptions such as "first", "second", etc. in the present application are only for the purpose of description, and are not intended to mean the order or sequence, nor to limit the present application, but merely to distinguish components or operations described by the same technical terms, and cannot be understood as indicating or implying the relative importance of the indicated technical features or implicitly indicating the number of the indicated technical features. Therefore, the features defined as "first", "second" can explicitly or implicitly include at least one of the features. In addition, the technical solutions of various embodiments can be combined with each other, but it must be based on the realization of ordinary skilled in the art, when the combination of technical solutions appears contradictory or unachievable, it should be considered that the combination of technical solutions does not exist, nor within the scope of protection required by the present application.

[0257] The present application measures the distribution of excitation transmission magnetic field and detection receiving magnetic field to inverse the distribution of magnetic permeability, conductivity and relative permittivity characteristic parameters, which is suitable for various magnetic resonance imaging facilities based on superconducting magnet, permanent magnet and electromagnetic magnet, etc., and is not limited by the strength of the main magnetic field. On the contrary, for facilities with field strength above 3.0T, the present application can effectively evaluate the specific absorption rate (SAR), the coil radio frequency magnetic field uniformity, etc. through the measurement of transmission magnetic field and receiving magnetic field. The measurement of magnetic resonance transmission magnetic field and receiving magnetic field is suitable for radio frequency configuration of full-body birdcage coil transmission-local coil receiving, such as head and neck joint receiving coil, breast coil, etc., and is also suitable for transverse electromagnetic multi-channel transceiving integrated coil or parallel imaging surface coil array configuration. Within the scanning capability range of human tissue magnetic resonance imaging, the present application can check a wide range of tissue organs such as cranium, pelvic cavity and limbs, etc. through magnetic field measurement and electromagnetic characteristic imaging. The scanning pulse sequence is based on spin echo or fast spin echo signal formation mechanism, and implements two-dimensional frequency encoding selected layer tomography or three-dimensional phase encoding selected layer volume imaging mode. In electromagnetic characteristic inversion, the former involves two-dimensional approximation method assuming that the characteristic gradient is zero, and the latter is not restricted by the boundary condition of thin layer characteristics. Since the electromagnetic characteristic parameter value varies with frequency, the inversion of magnetic permeability and complex permittivity is only the electromagnetic characteristic value at the specific field strength corresponding to the Larmor resonance frequency, such as 21.288MHz, 63.86MHz and 127.73MHz frequency response corresponding to 0.5T, 1.5T and 3.0T systems respectively.

[0258] Figure 1 The device is shown for measuring magnetic resonance radio frequency transmit and receive magnetic field and imaging tissue electrical and magnetic properties. Wherein, the main magnet 1 of the nuclear magnetic resonance imaging generates a uniform distribution of the magnetic polarization static main magnetic field acting on the imaging body 2, the excitation radio frequency transmit coil 3 generates a uniform distribution of the alternating magnetic field acting on the spin system in the imaging body 2, and the equivalent macroscopic magnetization vector is appropriately deflected to create the condition for the generation of the nuclear magnetic resonance signal. The transverse component amplitude formed by the deflection of the magnetization vector is linearly related to the transmit magnetic field strength, and the component phase is 90° out of phase with the magnetic field wave. The conventional magnetic resonance imaging is a process of using the magnetic field gradient to modulate the spin precession frequency and phase in space, which corresponds to the frequency and phase of the k-space harmonic in the Fourier transform relationship. The wave number domain discrete sample data in the form of the time domain signal containing the spatial encoding information is collected by the receiving coil 4, and the distribution image of the equivalent spin density precession amplitude and phase is obtained after the inverse transform. Starting from the spin precession amplitude and phase information measured by the magnetic resonance imaging, the proton density is decoupled to obtain the receive magnetic field sensitive field distribution related to the magnetization vector of the imaging body 2, and the phase distribution of the receive magnetic field is obtained by processing the data of the sequence scanning with the same deflection angle twice, different echo time and the same repetition time parameters. Then, the transmit magnetic field sensitive field distribution under the target condition is obtained by using the existing receive magnetic field sensitive field and combining the original two magnetic resonance image amplitude data. Finally, the transmit magnetic field phase distribution is obtained by subtracting the existing receive magnetic field phase from the original magnetic resonance image phase. Based on the measured transmit magnetic field and receive magnetic field data, the standard Helmholtz wave equation or the magnetic field-property core equation containing the gradient factor can be solved to inverse the distribution of the magnetic permeability, electrical conductivity and relative permittivity parameters in the imaging body 2.

[0259] The workflow of the present application is shown as Figure 2 The steps include the following steps,

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

[0261] Step 2: Perform 3 times of fast spin echo sequence imaging scanning, set different specific flip angles, echo times and repetition times and other sequence parameters, such as (90°, TE, TR), (45°, TE, TR) and (45°, TE / 2, TR); in addition, change the polarity of the selected layer, phase encoding and frequency encoding gradient, and repeat the 3 times of scanning; keep the geometric parameters such as field of view, voxel, layer thickness and layer number consistent in the 6 times of scanning;

[0262] Step 3: Store the scanning image data;

[0263] Step 4: Process the image data, compare and verify that the ratio of the maximum amplitude of different flip angle sequences meets the theoretical expectation, and implement phase unwinding measures;

[0264] Step 5: Replace the homogeneous water phantom and repeat steps 1-4. Perform imaging at the same facility gradient center position, and keep the field of view, voxel, slice level, slice thickness and other sequence parameters consistent with the parameters used in the previous patient imaging.

[0265] Step 6: Using the patient amplitude acquisition data with different flip angles, echo times, and repeated time-series scans, combined with the corresponding data acquired by the homogeneous water phantom, the amplitude distribution of the receiving magnetic field sensitive field of the patient image is obtained; using the patient phase acquisition data with the flip angle, proportional echo time, and repeated time-series scans, combined with the corresponding data acquired by the homogeneous water phantom, the phase distribution of the receiving magnetic field of the patient image is obtained.

[0266] Step 7: Still based on the patient amplitude acquisition data with different flip angles, echo times, and repeated time-series scans, and combined with the obtained received magnetic field sensitive field amplitude, process to obtain the emitted magnetic field sensitive field amplitude of the patient image; still using the patient phase acquisition data with the flip angles, echo times, and repeated time-series scans, and combined with the obtained received magnetic field phase, process to obtain the emitted magnetic field phase distribution of the patient image;

[0267] Step 8: Using the first type of inversion method of electromagnetic property imaging of the present invention, firstly, the electrical properties are obtained by inverting the transmitted magnetic field data using existing magnetic resonance electrical property imaging, and then the magnetic properties are inverted based on the convection-response equation or the standard Helmholtz equation; or, using the second type of inversion method of the present invention, the distribution of electrical and magnetic property parameters are simultaneously inverted from the transmitted magnetic field and received magnetic field data based on the Helmholtz equation that ignores the property gradient.

[0268] Based on the precession process of the macroscopic magnetization intensity vector during imaging scanning, this invention, combined with the general characteristics of wave vector amplitude and phase changes, designs a method for measuring the distribution of the radio frequency transmitting magnetic field and the receiving magnetic field. For the scanning target image 2, the lumped longitudinal relaxation time and transverse relaxation time of each voxel are given, meaning that the rates of recovery of its longitudinal magnetic vector component and decay of its transverse magnetic vector component in space are also given. Figure 3 In Figure (a), the equivalent spin angular momentum A1 and A2 at a certain voxel position in two scans are examined. Under different excitation magnetic field amplitudes, different initial magnetic vector chamfers are generated, resulting in different magnitudes of the transverse magnetic vector components. Here, A1 is longer than A2. If the signal acquisition echo time and magnetic vector recovery repetition time of the two scan sequences are consistent, based on the phase angle... In each echo formation, A1 and A2 always maintain equal phase, i.e., the magnetic vectors are in the same direction. The amplitude ratio of A1 and A2 depends on the initial transverse vector magnitude because the longitudinal relaxation time and transverse relaxation time of the same voxel are constant. The acquired signal of magnetic resonance imaging is essentially a magnetic field radio frequency wave in amplitude-phase form. The spatial decoding of the precession frequency and phase by the static magnetic field gradient defines the voxel array. Therefore, this invention provides the possibility of decoupling the proton density from the detection signal to measure the sensitive field amplitude of the received magnetic field by scanning the imaging body in two sequences with different deflection angles, consistent echo times, and repetition times.

[0269] For the voxel-based spin precession amplitude and phase information obtained from the inverse Fourier transform of conventional magnetic resonance imaging (MRI), the effect of the radio frequency pulse-excited transmitting magnetic field on the initial steady-state magnetic field vector not only deflects the absolute longitudinal magnetic vector in direction, generating a transverse component of the magnetic vector, but also the initial phase of this 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 MRI represents the sum of the continuous phases of the receiving magnetic field and the transmitting magnetic field. Since the transmission and reception propagation directions are opposite, the phase of the transmitting magnetic field is negative during the summation. Under certain conditions, such as the classic use of orthogonal dual-channel birdcage coils as a shared excitation and detection coil, the initial phase of the transmitting magnetic field is approximately equal to half the phase of the detection signal because the transmit / receive conversion is achieved by changing the port polarity. During the formation of the acquired echo signal, the relative dephasing of the precession between voxel arrays precisely represents the difference in the amplitude and phase of the transverse magnetic vector, and is reconstructed and displayed by the inverse Fourier transform. Similar to the aforementioned measurement of magnetic field amplitude, this invention requires the execution of two scanning sequences with special parameters to measure the magnetic field phase. The two sequences maintain the same deflection angle, constant echo time ratio, and consistent repetition time, for example, (45°, TE, TR) and (45°, TE / 2, TR). Figure 3 Figure (b) shows the operating amplitude and phase during two scan echoes at the same voxel position under the same deflection angle. Because the deflection angles are equal and the echo times differ by half, although the amplitudes of A3 and A4 are not strictly equal, they decrease according to the same exponential decay curve, while there is a constant phase difference between the relative astigmatism of the two precessions. Figure 3 Figure (c) shows two spin echo signals generated by the same voxel. With consistent longitudinal and transverse relaxation rates, conditions are created for measuring the phase of the received magnetic resonance radio frequency magnetic field, ensuring amplitude correlation and phase coherence of the transverse component precession. Because the phase constant of the transmitting magnetic field is transferred to the initial transverse component of the magnetic vector during excitation, subtracting the two original magnetic resonance phase distribution data yields a single received magnetic field phase related to the echo time ratio. Using the phase distribution data of the original reconstructed magnetic resonance image—i.e., the continuous phase measurement of the transmitting and receiving magnetic fields—the phase distribution of the transmitting magnetic field is further calculated from the received magnetic field phase.

[0270] When scanning with a fast spin echo sequence with echo time TE, such as (45°, TE, TR), in which the interval between adjacent echoes TS is determined by both TE and the echo train length ETL, and N data are acquired in each echo with a sampling data interval of td, the echo receiving time is N*td, and the received signal-to-noise ratio is assumed to be rSNR. When scanning with a sequence with echo time TE / 2, such as (45°, TE / 2, TR), and ETL is unchanged, the interval between adjacent echoes is TS / 2, and if the number of data acquired in each echo N is unchanged, the sampling data interval of each echo is correspondingly reduced to td / 2, and the echo receiving time is N*td / 2. Because the echo time is shorter, the received signal-to-noise ratio is reduced to about 0.707*rSNR. In order to keep the signal-to-noise ratio consistent between the two scans, i.e., under the action of the same excitation deflection angle pulse, the transverse component of the magnetic vector from the same amplitude decays according to the same curve, the present application adjusts the bandwidth BW of the sequence (45°, TE / 2, TR) to adjust td indirectly. Again, because tacq is related to 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 number of echo sampling points N and tacq.

[0271] The physical process of imaging can be regarded as a reciprocating motion process of the magnetic vector in the rotating coordinate system. Figure 3 Figure (c) of the present application analyzes the process of forming signals in the imaging body using the same flip angle and different echo time scanning sequences to measure the radio frequency magnetic field phase, i.e., the precession phase of the transverse component of the magnetic vector. Figure 4 From the spatial dimension, the process of the transverse component of the magnetic vector precession is further analyzed to explain the principle of measuring the magnetic field phase of the present application. Among them, Figure 4 Figure (a) of the present application shows the case where the average precession magnetic vector is located on the coordinate axis and each magnetic vector system back convergence peak is located on the same coordinate half axis under the action of the refocusing pulse, Figure 4Fig. (b) shows the case that the average precession magnetic moment is not located on the coordinate axis, but each magnetic moment system re-focuses the peak on the coordinate axis. In the rotating coordinate system x'y'z, using the fast spin echo sequence with echo time TE, such as (45°, TE, TR), scanning the imaging body (2) to produce the magnetic moment partial transverse component at the t=0 moment, the maximum relative phase difference π / 2 is formed between the transverse components in the TS / 2 time, and it is assumed that the medium rotating magnetic moment is along the +y' axis direction at this time, the first re-focus pulse is implemented. At the TS moment, the first echo signal peak is formed. The equal TS is selected so that the echo signal peak is formed in the same time and in the same direction, and the maximum degree of each cycle is also consistent, which is the phase difference π / 2. When the excitation pulse with the same deflection angle, the fast spin echo sequence with echo time TE / 2, such as (45°, TE / 2, TR), scans the imaging body 2 to produce the same size magnetic moment partial transverse component at the t=0 moment, and the maximum phase difference π / 4 is formed between the components in the TS / 4 time, it is assumed that the medium rotating magnetic moment is at an angle of 45° with the +x' axis, and the first re-focus pulse is implemented. At the TS / 2 moment, the first signal peak is formed, and the magnetic moment transverse component points to the -y' axis negative direction. Subsequently, every TS / 2, the 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 echo amplitude of the (45°, TE / 2, TR) sequence is larger and the signal-to-noise ratio is higher because the transverse relaxation has not been relaxed for a long time. The present application measures the phase of the radio frequency magnetic field, that is, the precession phase of the magnetic moment transverse component, which is realized by controlling the constant difference between the maximum degree of phase difference formed by the precession dispersion according to the length of the echo interval time of different sequences, to measure the receiving phase under the same excitation phase, and inversely combining the existing original magnetic resonance phase and receiving phase to calculate the relatively uniform transmitting phase. In the fast spin echo sequence, if one echo time parameter TE is set, it specifically refers to the time corresponding to the middle echo signal in the echo signal chain.

[0272] On the basis of the conventional magnetic resonance imaging examination, the present application needs to save and record the data of the argument image in addition to the amplitude image, and record the excitation magnetic field amplitude B1+RMS of each scan of the imaging device, so as to be compared with the magnetic field measurement value. Specifically, one embodiment is that the radio frequency magnetic field measurement and electromagnetic characteristic inversion of the cranium and breast are performed by using an 8-channel phased array parallel imaging coil, the measurement geometry parameters are set in combination with the magnetic field gradient, the field of view FOV is 256*256 mm 2 , the frequency encoding and phase encoding steps are 256*256, the layer thickness is 2 mm, the layer spacing is 2 mm, and the number of layers is 12.

[0273] Preferably, the deflection angles of the sequences I and II are set to 120° and 60°, respectively, in order to make the histogram distribution of the deflection angle cosine value image data fall into the approximate normal distribution range centered at 0.5 as much as possible, instead of cos45° = 0.7071, when calculating the transmit magnetic field sensitive field data. Unlike the existing double angle method of B1-mapping technique, the refocusing pulse deflection angles of the fast spin echo sequences I and II in the present application 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 added with a corresponding reverse polarity sequence, and the normal polarity argument data and the reverse polarity argument data obtained are averaged pixel 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 electrical and magnetic characteristic parameter distribution of the reconstructed cranium, the fluid-attenuated inversion recovery imaging (FLAIR) sequence is implemented on the human tissue and the water phantom in turn according to one embodiment. The sequence excitation and detection parameters of the magnetic resonance imaging original image data acquisition are shown in Table 3.

[0276] Table 3

[0277] Sequence T2-FLAIR (a, TE, TR) Amplitude data Phase data Antipole phase data Head I (120°, 110 ms, 9000 ms) m_2a_h - m_2a_h_po II (60°, 110 ms, 9000 ms) m_a_h a_a_h a_a_h_po III (60°, 55 ms, 9000 ms) - a_t_h a_t_h_po Water phantom IV (120°, 110 ms, 9000 ms) m_2a_p - m_2a_p_po V (60°, 110 ms, 9000 ms) m_a_p a_a_p a_a_p_po VI (60°, 55 ms, 9000 ms) - a_t_p a_t_p_po

[0278] In Figure 5 The radio frequency Magnetic field measurement method based on the amplitude ratio and phase difference principle is given in the magnetic resonance receiving magnetic field amplitude sensitive field and phase distribution measurement embodiment shown in the figure. First, the cranium tissue is scanned by the main sequence II with the parameters of deflection angle, echo time and repetition time being (60°, 110 ms, 9000 ms) to obtain the amplitude and phase angle data m_α_h and a_α_h, wherein m represents the amplitude, a represents the argument, α represents the deflection angle, and h represents the human head scanning part. Then, the auxiliary sequence I is scanned with the parameters of deflection angle, echo time and repetition time being (120°, 110 ms, 9000 ms) to obtain the amplitude data m_2α_h. Based on the pixel level, the obtained amplitude data m_2α_h is divided by the data m_α_h to obtain the cranium receiving magnetic field sensitive field amplitude distribution data m h (r) of the transmit magnetic field sine ratio weighted by the proton density decoupling, wherein m h (r) contains the contribution of the receiving coil magnetic field inhomogeneity and is related to the selected TE and TR. Since the transmit magnetic field weighting function is nonlinear, the homogeneous water phantom magnetic field data is measured with the same parameter sequence, and the brain tissue magnetic permeability is assumed to be the water magnetic permeability, and the influence of the transmit magnetic field is further eliminated after processing.

[0279] Continue to scan the homogeneous water phantom with the same imaging parameters of sequence V and IV, respectively, to obtain the amplitude and phase data m_α_p and a_α_p, and the amplitude m_2α_p, where p represents the water phantom. Divide the obtained amplitude m_2α_p by m_α_p to obtain the proton density decoupled transmit magnetic field sine ratio weighted water phantom receive magnetic field sensitive field amplitude distribution data m p (r). Since the water phantom is homogeneous, the m p (r) only contains the contribution of the receive coil magnetic field distribution fluctuation, i.e., non-uniformity.

[0280] Divide the transmit weighted receive magnetic field amplitude m h (r) obtained by processing the brain scan measurement by the receive magnetic field m p (r) of the water phantom to obtain the brain receive magnetic field sensitive field amplitude distribution under the condition of unit current, eliminating the transmit magnetic field sine amplitude ratio The superscript TE represents the condition under a specific echo time.

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

[0282] Subtract the phase angle data a_α_h' and a_t_h' obtained by the two brain scans with gradient polarity correction of eddy current interference to obtain the receive magnetic field phase Δp h (r) related to the selected echo time TE. Due to the existence of certain non-uniformity of the actual receive coil magnetic field, the brain phase distribution Δp h (r) contains the contribution of the receive coil field fluctuation in addition to the contribution of the electromagnetic properties of the tissue.

[0283] Similarly, continue to scan the water phantom with the same imaging parameters (60°, 110 ms, 9000 ms) sequence V to obtain the phase angle data a_α_p and a_α_p_po, and the auxiliary sequence VI with parameters (60°, 55 ms, 9000 ms) to obtain the data a_t_p and a_t_p_po. After correction of the eddy current effect, subtract a_t_p' from a_α_p' to obtain the phase difference Δp p (r) between the sequences. The data Δpp (r) contains the same receive coil intrinsic phase field inhomogeneity as data Δp h (r) in (r).

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

[0285] The sensitivity field amplitude and phase distribution data of the receive magnetic field under the unit current of the brain have been calculated, and the general processing method based on the amplitude ratio and the phase difference of the spatial domain periodic magnetic field wave is still used. Combined with the measurement scanning data of the foregoing sequence, the transmit magnetic field amplitude and phase under specific parameters are obtained through the following data operation, and the data processing process is as shown in Figure 6 .

[0286] In order to obtain the transmit magnetic field amplitude excited under the brain scanning main sequence II (60°, 110 ms, 9000 ms), first, the original magnetic resonance image amplitude data m_α_h is divided by the unit receive magnetic field amplitude sin quantity data t2 of the remaining proton density amplitude weighted transmit magnetic field sensitivity field. Similarly, under the condition that the same voxel position of the same imaging body, the image amplitude m_2α_h of the auxiliary sequence I with a 2-fold deflection angle, the same echo and repetition time is divided by the remaining transmit magnetic field sine quantity data t1 with the same proton density weight.

[0287] Considering that the transmit magnetic field amplitude is nested in the sine trigonometric function and the consistent proton density distribution contribution, the intermediate process quantity t1 is further divided by t2 by using the sine double-angle identity, and the transmit magnetic field amplitude of the brain related to the electromagnetic characteristics of itself and the coil magnetic field inhomogeneity is calculated , that is, the magnetic field under the excitation of the radio frequency pulse with a 60° deflection angle, where τ is the pulse width. Note that it is different from the principle that the magnetic resonance imaging requires the uniformity of the radio frequency excitation magnetic field amplitude without the superscript ‘~’ reflecting the principle.

[0288] In the foregoing, the phase distribution of the receive magnetic field of the brain scanning has been measured and calculated and processed Subtracting the eddy current corrected raw magnetic resonance image phase data a_α_h' of the main sequence II from it. Since the excitation field phase e in the magnetization expression is multiplied by 1 imaginary unit, i.e. the transverse component of the magnetic vector lags, subtracting π / 2 from the difference gives the transmit field phase distribution data where α denotes the deflection angle of 60°. Similarly to the amplitude of the transmit field, the phase distribution of the excitation field strength under unloaded conditions also tends to be uniformly distributed according to the principle of magnetic resonance imaging, in order to approach the ideal deflection angle uniform distribution condition.

[0289] In order to measure the electrical and magnetic property parameter distribution of the reconstructed imaging breast tissue, a multi-channel surface coil array is used to implement the transverse relaxation time weighted scan sequence (T2-Weighted Imaging, T2WI) of human tissue and phantom in the following embodiment. Considering the characteristics of breast tissue related to metabolites and mainly composed of fat, the repetition time TR is set to 4000 ms to save scanning time. The sequence excitation and detection parameters of the magnetic resonance imaging raw image data acquisition are shown in Table 4.

[0290] Table 4

[0291] Sequence T2WI (a, TE, TR) Amplitude data Phase data Antipole phase data Breast I (120°, 90 ms, 4000 ms) m_2a_b - - II (60°, 90 ms, 4000 ms) m_a_b a_a_b a_a_b_po III (60°, 45 ms, 4000 ms) - a_t_b a_t_b_po Phantom IV (120°, 90 ms, 4000 ms) m_2a_p - - V (60°, 90 ms, 4000 ms) m_a_p a_a_p a_a_p_po VI (60°, 45 ms, 4000 ms) - a_t_p a_t_p_po

[0292] In Figure 5 In the magnetic resonance received field amplitude sensitive field and phase distribution measurement embodiment shown in the figure, a radio frequency magnetic field measurement method based on amplitude ratio and phase difference principle is given. First, the main sequence II with the parameters of deflection angle, echo time and repetition time of (60°, 90 ms, 4000 ms) is used to scan the breast tissue, and the amplitude and phase angle data m_α_b and a_α_b are obtained, where b represents the scanned part of the human breast. Then, the auxiliary sequence I with the parameters of deflection angle, echo time and repetition time of (120°, 90 ms, 4000 ms) is used for scanning, and the amplitude data m_2α_b is obtained. Based on the pixel level, the obtained amplitude data m_2α_b is divided by the data m_α_b, and the breast received field sensitive field amplitude distribution data m b (r) is obtained, where m b (r) contains the contribution of the non-uniformity of the received coil magnetic field, and is related to the selected TE and TR. Since the transmit field weighting function nonlinear, the homogeneous phantom magnetic field data is measured with the same parameter sequence, and the influence of the transmit field is further eliminated after assuming that the magnetic permeability of the breast tissue is the magnetic permeability of the phantom.

[0293] The homogeneous phantom model is scanned again using sequences V and IV with consistent parameter sets, yielding amplitude and argument data m_α_p and a_α_p, as well as the amplitude m_2α_p. Dividing the obtained amplitude m_2α_p by m_α_p yields the amplitude distribution data m_p of the phantom model's receiving magnetic field sensitive field, which is decoupled from the proton density and weighted by the sinusoidal ratio of the emitted magnetic field. p (r). Due to the homogeneity of the phantom, the m p (r) contains only the contribution of the magnetic field distribution fluctuation of the receiving coil, i.e., the non-uniformity.

[0294] The received magnetic field amplitude m, which includes emission weighting, is obtained from the breast scan measurement. b (r) divided by the receiving magnetic field m of the phantom p (r) yields the amplitude distribution of the sensitive field of the breast receiving magnetic field under unit current conditions, eliminating the amplitude ratio of the sinusoidal value of the emitted magnetic field.

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

[0296] Subtracting the argument data a_α_b' and a_t_b' obtained from the two breast scans after gradient polarity correction for eddy current interference, yields the received magnetic field phase Δp, which does not include the transmitted magnetic field phase. b (r), which is related to the selected echo time TE. Due to the inherent non-uniformity of the magnetic field in the actual receiving coil, the phase distribution Δp of the breast tissue is... b In (r), in addition to the contribution of the organization's electromagnetic properties, there is also the contribution of the field fluctuation of the receiving coil.

[0297] Similarly, the phantom is scanned again using a sequence V with consistent imaging parameters (60°, 90ms, 4000ms) to obtain argument data a_α_p and a_α_p_po, and data a_t_p and a_t_p_po are obtained by scanning with an auxiliary sequence VI with parameters (60°, 45ms, 4000ms). After correction for eddy current effects, the phase difference Δp between the sequences is obtained by subtracting a_t_p' from a_α_p'. p (r). Data Δp p The (r) distribution contains data Δp b(r) the same receive coil intrinsic phase field inhomogeneity.

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

[0299] Having calculated the breast unit current received magnetic field sensitivity field amplitude and phase distribution data, the general processing method based on amplitude ratio and phase difference of spatial domain periodic magnetic field wave is still applicable, combined with the measurement scan data of the aforementioned sequences, the transmit magnetic field amplitude and phase under specific parameters are obtained through the following data operation, the data processing process is as shown in Figure 6 .

[0300] In order to obtain the breast scan main sequence II (60°, 90 ms, 4000 ms) under the excitation of the transmit magnetic field amplitude, first divide the original magnetic resonance image amplitude data m_α_b by the unit received magnetic field amplitude The remaining proton density amplitude weighted transmit magnetic field sensitivity field amplitude sine quantity data t2. Similarly, under the condition of the same voxel position of the same imaging body, divide the image amplitude m_2α_b of the 2 times deflection angle, the same echo and repetition time auxiliary sequence I by The remaining same proton density weighted transmit magnetic field sine quantity data t1.

[0301] Considering that the transmit magnetic field amplitude is nested in the sine trigonometric function And the consistent proton density distribution contribution, using the sine double angle identity, further divide the intermediate process quantity t1 by t2, calculate the breast transmit magnetic field amplitude related to its own electromagnetic characteristics and coil magnetic field inhomogeneity , which is the magnetic field under the excitation of the radio frequency pulse 60° deflection angle, where τ is the pulse width.

[0302] In the foregoing, the phase distribution of the breast scan received magnetic field is obtained by subtracting the eddy current corrected original magnetic resonance image argument data a_α_b' of the main sequence II from it. Since the excitation magnetic field phase e in the magnetization intensity expression is multiplied by 1 imaginary unit, that is, the transverse component of the magnetic vector lags, subtract π / 2 from the difference value to obtain the transmit magnetic field phase distribution data

[0303] Assume that the electromagnetic properties of the imaging body are isotropic, i.e. the circularly polarized right-handed direction of the transmitted magnetic field and the left-handed direction of the received magnetic field interact with the substance, and the scalar magnetic permeability and the complex permittivity amplitude measured under different magnetic field gradient encoding and decoding directions are consistent, and do not exhibit as a tensor. On the other hand, assume that the electromagnetic properties of the imaging body have linear properties, i.e. when the same voxel interacts with different amplitudes of the transmitted magnetic field and the received magnetic field, the internal magnetic flux density generated is proportional to the amplitude of the 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, i.e. the magnetic permeability is a real number.

[0304] According to one embodiment of the present application for inverting the electromagnetic properties of human tissue from the measured magnetic field, i.e. calculating the magnetic permeability parameter distribution from the first type of property inversion method, Figure 7 Fig. (a) shows the transmitted magnetic field measured during the imaging of the imaging body by magnetic resonance scanning, and under the assumption that the magnetic permeability in the body is uniformly distributed and equal to the magnetic permeability of vacuum, the electrical property distribution of the human tissue is first inverted by the convection-reaction equation method containing the permittivity gradient. Then, using the received magnetic field data measured by the same imaging body, combining the electrical properties obtained, and still based on the convection-reaction equation containing the property gradient, the magnetic property distribution, i.e. the magnetic permeability distribution, is inverted to improve the accuracy of the results.

[0305] The form of the core equation for inversion is derived below, and the use of the assumed conditions is also statistically analyzed. The electromagnetic wave equation in the form of a time-harmonic vector is:

[0306]

[0307] Here, both the electrical property gradient term and the field differential are included, and the electromagnetic property distribution can be solved by combining the boundary conditions. At present, except for the form of the field as a simple harmonic wave, there is no assumption or approximation processing.

[0308] For the first type of property inversion method, the electrical properties are first assumed to be homogeneous to solve the magnetic properties, i.e. ①: the magnetic homogeneity in the imaging body satisfies the Gauss magnetic law for the magnetic field strength, i.e. μ = μ0, and replacing H with the measured quantity B, we have:

[0309]

[0310] Referring to the form of the convection-reaction partial differential equation in fluid mechanics, the core inversion equation of the magnetic field differential and the complex permittivity is further derived from the above equation, and a sparse linear equation set based on the voxel array is constructed by combining the single-transmitted magnetic field measurement data. With the necessary setting of the boundary conditions, the electrical property parameter distribution can be reconstructed by using the least square 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, are combined to further invert the magnetic characteristic parameter distribution. Still starting from the assumption-free equation (57), assumption ②: the imaging body is magnetically homogeneous, then according to the Gauss magnetic law The second term on the left side of the equation is eliminated, and the other terms on the left and right sides are written as functions of the magnetic flux density, and the homogeneous permeability is removed from the differential operator, and we have:

[0312]

[0313] In the formula, is distinguished from the second type of characteristic inversion method, is the electrical characteristic distribution obtained by the existing electrical characteristic imaging inversion for calculating the magnetic characteristics in the first type of method; the magnetic flux density B = μ b H, μ b is the magnetic permeability. In this way, the complexity of the solving process is reduced, and the constitutive equation about the unknown magnetic permeability is explicitly derived. Continue to assume ③: the magnetic permeability μ b is non-homogeneous, i.e. changes with the change of the tissue, and the core equation is a linear equation about the known quantity and the unknown quantity μ b , and the magnetic characteristic distribution can be solved using the least squares method.

[0314] In summary, in the above electromagnetic characteristic inversion calculation, 1 magnetic homogeneity assumption is used to solve the electrical characteristic parameters, and 1 magnetic homogeneity and 1 magnetic non-homogeneity assumption are used to solve the magnetic characteristic parameters. At this point, using the known magnetic field measurement sensitive field amplitude and phase data, the inversion of the electrical conductivity and relative permittivity and magnetic permeability parameter distribution of the imaging body is completed, i.e. the related image reconstruction is realized.

[0315] According to one embodiment of the present application for inverting the electromagnetic characteristics of human tissues from measured magnetic fields, i.e. calculating the magnetic permeability parameter distribution by the second type of characteristic inversion method, in Figure 7 Fig. (b) shows the process of simultaneously inverting the electrical and magnetic characteristic distributions from the measured and processed transmission and reception magnetic field data in the imaging body. In order to reduce the complexity of the core equation of the electromagnetic characteristic inversion, it is assumed that the distributions of the electrical and magnetic characteristic gradients are both zero, i.e. the imaging body is homogeneous, so that the boundaries between different tissues in the body are ignored, and the approximate Helmholtz wave equation is derived to simultaneously invert the magnetic permeability, electrical conductivity and permittivity parameter distributions. Among them, the electrical conductivity and permittivity form the 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 example of deriving the core equation for inverting electromagnetic properties by measuring the magnetic field, and an analysis of the inversion error, i.e., the application of statistical assumptions. Starting from Maxwell's equations, Faraday's law, and Ampere's law, which describe the behavior of electromagnetic fields within 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 equation (60), and B = μ b H, has:

[0319]

[0320]

[0321]

[0322]

[0323] At this point, the known quantity is B, and the unknown quantity to be determined is the non-homogeneous magnetic permeability μ of human tissue. b and complex capacitance The subscript 'b' represents the human body. Apart from the harmonic preconditions for magnetic resonance radio frequency magnetic fields, no other assumptions are made.

[0324] To reduce the complexity of the equations, we assume ①: the magnetic properties within the imaging body are homogeneous, i.e., the magnetic permeability is constant. According to Gauss's magnetic law, we have... Assumption ②: The bulk electrical properties are homogeneous, i.e., the complex capacitance is constant, therefore... As a trade-off, the second type of inversion method in this invention sacrifices the accuracy of the equation's solution, especially exhibiting systematic computational errors at tissue boundaries. Let v = 1 / μ b Then we have:

[0325]

[0326] Here, we make the counter-assumption ③: the magnetic field within the imaging body is non-homogeneous, i.e. And combining the relevant vector identities, write out the component form of equation (62).

[0327]

[0328] This leads to a set of equations relating the nonlinear relationships between electrical and magnetic properties. A single scan of the same imaging object yields a set of independent data on the distribution of the emitted and received magnetic fields, i.e., complex-valued known quantities. and The conditions necessary for solving the equation set containing two unknowns and two equations are basically met. And by two or more groups of transmitting and receiving magnetic field measurement quantities, such as changing the sequence parameters for additional excitation and detection, the real magnetic permeability and complex permittivity can be solved according to the complex constitutive equation set (63).

[0329] In summary, the above second type of inversion contains a first-order complex permittivity homogeneous assumption and a first-order magnetic permeability homogeneous assumption, and a first-order magnetic permeability non-homogeneous inverse assumption.

[0330] In order to illustrate the feasibility and effectiveness of the magnetic resonance magnetic field measurement and electrical and magnetic property imaging method of the present application, the following numerical model based on two properties is designed to simulate the radio frequency electromagnetic field distribution, obtain the inversion results of the electrical conductivity and permittivity distribution, and reconstruct and display images, as well as the original data images of the density, proton density, target electromagnetic properties, etc. of the accompanying model itself.

[0331] In one simulation embodiment of the present application, a homogeneous water phantom model containing an array of abnormal bodies, small cylindrical regions of different diameters are created and given different electrical conductivity and relative permittivity values. Under the condition of a 1.5T magnetic resonance imaging main magnetic field, i.e. a resonance frequency of 63.864MHz, an orthogonal birdcage type radio frequency coil is used to obtain the excitation transmitting magnetic field sensitive field amplitude and the transmitting-receiving continuous phase data in the detection signal. The length of the cubic model is 200mm, the width is 200mm, and the height is 11mm, and the effective region inside the main body is a large cylindrical type, the data point array size is [200, 200, 11], the original target electrical conductivity distribution range is [0.1, 2]S / m, and the background value is 0.5S / m; the original target relative permittivity distribution range is [50, 100], and the background value is 80, which is equal to the relative permittivity of water.

[0332] The target electrical conductivity distribution in the water phantom simulation containing an array of circular abnormal bodies is shown in Figure 8 The corresponding target relative permittivity distribution is shown in Figure 9 In the figure, the minimum characteristic difference that can be reflected is 0.1S / m and 1 relative unit respectively, and 1 voxel can represent the smallest square region, and for a fault, i.e. a pixel, 1mm*1mm.

[0333] The reconstructed transmitting magnetic field sensitive field amplitude distribution obtained by forward simulation is shown in Figure 10 The reconstructed continuous phase distribution of the received magnetic field and the transmitting magnetic field is shown in Figure 11 The amplitude change range is (1, 1.1)*10 -7 T, and the phase change range is (-1.2, -0.4) rad.

[0334] Based on the forward emission sensitive field amplitude and the received-emission continuous phase data, the conductivity distribution is obtained by inversion using the standard Helmholtz equation method without characteristic gradient as shown in Figure 12 The permittivity distribution is shown in Figure 13 As the actual inversion result numerical range has significant dispersion, the reference model original target conductivity and permittivity values are taken as the display range is limited 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 imaging body electromagnetic properties are homogeneous, the boundary between the background and the imaging body can still be clearly distinguished from the algorithm simulation results, and the preset circular array region and its boundary also have high sensitivity and specificity. Note that the amplitude fluctuation at the edge of a single circular region is due to the homogeneous assumption, and further low-pass filtering with a standard deviation of unit pixels can improve the influence of this spatial harmonic effect.

[0335] Another simulation embodiment of the present application about the forward radio frequency magnetic field of the human head model for the verification of the electrical property inversion algorithm, based on the 7T magnetic resonance imaging condition, a 16-channel transverse electromagnetic transceiving common radio frequency coil is constructed, the magnetic field distribution is calculated in the numerical head model derived from the real tissue characteristic parameters, the conductivity and relative permittivity parameter distribution is obtained by inversion using the emission-receiving dual source Dual method, and a series of image results are reconstructed. The model contains 41 different tissue types.

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

[0337] Assuming that the magnetic properties of the head model are homogeneous, i.e. the magnetic permeability is equal to the vacuum magnetic permeability, the emission magnetic field strength sensitive field amplitude and phase distribution are shown in Figs. 17a and 17b respectively when the multi-channel transverse electromagnetic coil is implemented with pulse harmonic continuous wave voltage excitation, the change ranges are (0, 25) A / m and (-3, 3) rad respectively. The received magnetic field strength sensitive field amplitude and phase distribution based on the propagation direction adjustment of the emission magnetic field are shown in Figs. 18a and 18b respectively, the change ranges are (0, 25) A / m and (-3, 3) rad respectively, it can be seen that the amplitude range is consistent with the amplitude of the emission magnetic field strength, and the contrast of the phase distribution is opposite to that of the emission magnetic field.

[0338] The core inversion equation of the dual-source Dual method has three unknowns, i.e., complex permittivity, complex permittivity transverse gradient, and complex permittivity longitudinal gradient. In order to calculate and reconstruct the electric characteristic parameter distribution from the radio frequency magnetic field, the embodiment forms four channel groups, i.e., (1, 5, 9, 13), (2, 6, 10, 14), (3, 7, 11, 15), and (4, 8, 12, 16), from a single channel of coil excitation and detection, and inverses the complex permittivity and its gradient from four groups of independent transmission magnetic field data based on a linear algebra equation set. According to the corresponding relationship between the electric characteristics and the real part and the imaginary part of the solution, the reconstructed conductivity distribution of the human head model is as shown in Figure 19 , the value range is [0, 1.5] S / m; and the reconstructed relative permittivity distribution is as shown in Figure 20 , the value range is [1, 70).

[0339] In another embodiment of the application, the transmission magnetic field and the reception magnetic field distribution in the human body tissue are measured based on the existing clinical conventional magnetic resonance imaging scanning, and the electromagnetic characteristic parameter distribution image is inversely reconstructed by using the Helmholtz equation method and the convection-reaction equation method. The specific scanning conditions are as follows: a 1.5T magnetic resonance system, a 2D imaging mode, a pulse sequence geometry parameter of a field of view (FOV) of 256*256 mm 2 , a frequency encoding step of 128, a phase encoding step of 128, a slice thickness of 4 mm, an interlayer distance of 1 mm, a number of layers of 9, and an echo time of 1. A proton density weighted fast spin echo sequence (Proton Density Weighted Imaging, PDWI) is used to perform a series of scans on a homogeneous water phantom, a homogeneous water phantom containing an abnormal shape, a homogeneous water phantom containing the same abnormal shape and 2.5 g of NaCl, a homogeneous water phantom containing the same abnormal shape, 2.5 g of NaCl, and 0.25 mL of gadolinium acid portugal (Gd-DOTA) with a concentration of 377 mg / mL, and a volunteer's human cranial and brain tissue, to obtain 24 groups of original magnetic resonance amplitude image and argument image data, and to compare the results of the inversion of the electric characteristic distribution from the magnetic resonance imaging radio frequency magnetic field and the inversion of the magnetic characteristic distribution of the application. Based on the low T1 and T2 contrast sequence of the proton density weighted sequence, the imaging speed is fast, and the parameters of three scans are (120°, 40 ms, 10000 ms), (60°, 40 ms, 10000 ms), and (60°, 20 ms, 10000 ms), respectively, wherein the first parameter is the excitation deflection angle, the second parameter is the echo time, and the third parameter is the repetition time, so as to satisfy the TR>>3-4xT1 condition of eliminating the contrast factor term as much as possible in the calculation of the magnetic field amplitude from the signal intensity.

[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 correcting the non-uniformity of the coil magnetic field. Fig. 21, a and b are respectively the original magnetic resonance proton density weighted amplitude and phase angle images of the abnormal shape homogeneous water phantom scanning. Fig. 22, a and b are respectively the original magnetic resonance proton density weighted amplitude and phase angle images of the same abnormal shape homogeneous water phantom with 2.5g NaCl. Fig. 23, a and b are respectively the original magnetic resonance proton density weighted amplitude and phase angle images of the same abnormal shape homogeneous water phantom with 2.5g NaCl and 0.25mL gadopentetic acid meglumine with a concentration of 377mg / mL. Fig. 24, a and b are respectively the original magnetic resonance proton density weighted amplitude and phase angle images of the human brain tissue scanning.

[0341] The magnetic resonance amplitude and phase angle images of the spherical homogeneous water phantom and the three abnormal shape homogeneous water phantoms and the human brain tissue scanning are processed to obtain the magnetic field distribution data. The specific processing includes the following processes.

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

[0343] (2) The amplitude data of the imaging object scanned by the PDWI auxiliary sequence (120°, 40ms, 10000ms) and the main sequence (60°, 40ms, 10000ms) respectively is divided by the unit received magnetic field sensitive field measurement data obtained to obtain the ratio data t1 and t2; t1 is divided by t2 to obtain the cosine term data related only to the transmitting magnetic field eliminating the tissue proton density weight Given that the width of the radio frequency pulse for the emitted magnetic field of the system's excitation coil is set to 4000 μs, by taking the inverse cosine of the cosine term data and dividing it by the gyromagnetic ratio and pulse width, the amplitude measurement data of the emitted magnetic field sensitive field corresponding to the main sequence (60°, 40ms, 10000ms) scanning imaging body are obtained. Note that this measurement is related to the excitation deflection angle.

[0344] (3) Subtract the phase data of the auxiliary sequence (60°, 20ms, 10000ms) from the phase data of the image object scanned by the PDWI main sequence (60°, 40ms, 10000ms) to obtain the phase difference data Δp. o The phase difference data Δp is obtained by subtracting the phase data of the homogeneous water model scanned with the sequence (60°, 20ms, 10000ms) from the phase data of the homogeneous water model scanned with the PDWI sequence (60°, 40ms, 10000ms). b It only includes the contribution of non-uniformity in the phase of the receiving coil's magnetic field; then the phase difference data Δp of the imaging body is used. o Subtract the phase difference Δp of the homogeneous water model b Considering that the echo time parameters of the two sets of sequences are twice that of each other, the received magnetic field phase measurement data of the main sequence (60°, 40ms, 10000ms) scanning imaging body are obtained as follows: Because it includes an intermediate processing step of subtracting the transmitting magnetic field phase from the original magnetic resonance phase, the receiving magnetic field phase does not contain the contribution of the non-uniformity of the transmitting coil magnetic field phase.

[0345] (4) The received magnetic field phase of the PDWI master sequence (60°, 40ms, 10000ms) scanning image has been calculated and processed. Based on... Among them, a_6_o is the original magnetic resonance phase of the main sequence (60°, 40ms, 10000ms), from which the corresponding emission magnetic field phase measurement data of the main sequence can be obtained.

[0346] For each set of fast spin echo sequence scans performed on each water phantom and cranial tissue, in addition to scanning one set of data with normal or default gradient polarity, another set of data was scanned by changing the gradient polarity without altering 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) were averaged to obtain the original MRI phase data of the corresponding sequence scan of the subject, which was used to calculate the phase distribution of the measured received magnetic field and the emitted magnetic field of the corresponding subject. The refocusing flip angle of all 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 same subject in two scans is equal or close to equal, which is conducive to keeping the results consistent in amplitude in the calculation of the measured receiving magnetic field and the transmitted magnetic field sensitive field of the corresponding subject from the corresponding primary magnetic resonance amplitude data. Under the condition that the imaging geometry parameters remain unchanged, changing the bandwidth and thus the sampling rate and the acquisition time will affect the spatial resolution, which can be compensated by the automatic magnetic field gradient strength 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, and the unit sampling time is 3.3μs and 2.5μs respectively, and the echo acquisition time is 427μs and 320μs respectively. According to the formula (55) of the relationship between the pixel resolution R RO and the gradient G RO and the acquisition time, under the condition of the same resolution, the calculated gradients of the sequences (60°, 40ms, 10000ms) and (60°, 20ms, 10000ms) are 27.5mT / m and 36.7mT / m respectively, both of which are less than the rated value of 45mT / m under the standard of general magnetic resonance medical equipment.

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

[0349] Fig. 26a and Fig. 26b are respectively the transmitted magnetic field sensitive field amplitude and phase distribution images measured by scanning the homogeneous water phantom containing the same abnormal shape with 2.5g of NaCl, and Fig. 26c and Fig. 26d are respectively the corresponding receiving magnetic field sensitive field amplitude and phase images; in the figure, the gradual distribution of the magnetic field phase reflects the change of the electromagnetic characteristics of the content of the water phantom.

[0350] Figures 27a and 27b show the amplitude and phase distribution images of the sensitive field of the emitted magnetic field obtained by scanning a homogeneous water phantom with the same abnormal shape, after adding 2.5g NaCl and 0.25mL of gadotyl diglucate at a concentration of 377mg / mL. Figures 27c and 27d show the amplitude and phase distribution images of the sensitive field of the received magnetic field, respectively. In the emitted magnetic field and received magnetic field distribution images reconstructed from actual scanning measurements as shown in the figures and above, in order to effectively use the scanning data of the spherical homogeneous water phantom to assist in solving the emitted magnetic field and received magnetic field distribution data of the homogeneous water phantom with abnormal shape and the aforementioned cranial tissue imaging, the reconstructed image results are all the intersection of the signal regions of the spherical homogeneous water phantom and the target imaging body.

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

[0352] Based on the fundamental Helmholtz equation method that ignores electromagnetic property gradients, the conductivity and permittivity of the homogeneous water model with an anomalous shape, obtained from the original magnetic resonance equivalent proton density measurement and substituted with the emitted and received magnetic field distribution data, and inverted and reconstructed, are shown in Figures a, b, c, and d of Figure 29, respectively. When calculating the magnetic field gradient and second derivative, the Savitzky-Golay filter design principle was used to solve for the least-squares quadratic parabolic fitting coefficients, avoiding the accuracy degradation caused by direct correlation division. In the figures, the voxel electrical property parameters shown are the original values ​​obtained from the inversion, as no regularization or normalization processing was performed. In the figures, the changes in brightness reflect the changes in local CuSO4 concentration.

[0353] Based on the fundamental Helmholtz equation method that ignores electromagnetic gradients, the emitted and received magnetic field distribution data of the cranial tissue scan, calculated from the original magnetic resonance equivalent proton density measurement, were substituted into the inverted and reconstructed conductivity and permittivity, as shown in Figures a, b, c, and d of Figure 30. In Figure c of Figure 30, the anatomical structures of tissues such as the corpus callosum are relatively clearly visible; while... Figure 32 The obvious salt-and-pepper noise present in Figures (a) and (b) can be further processed by median filtering.

[0354] In another embodiment of the imaging scan of the present application for measuring the magnetic resonance radio frequency magnetic field, the fast spin echo pulse sequence set parameters used to maximize PD weighting while minimizing Tl and T2 weighting are I (120°, 24 ms, 9000 ms), II (120°, 24 ms, 9000 ms), and III (120°, 32 ms, 9000 ms). For sequences I and II, with a receive bandwidth of 62.5 kHz and a number of phase encoding steps of 128, the single shot sampling time is 8 μs and the single echo acquisition time is 1.024 ms. With an echo train length of 8, the inter-echo spacing is 6 ms, taking into account the transmit coil time constant (dead time) and the pre-arming control time of the radio frequency power amplifier. The frequency encoding read gradient strength is calculated to be 11.46 mT / m based on the echo acquisition time, and the Rel. SNR is calculated to be 28% based on the definition of relative signal to noise ratio from the bandwidth.

[0355] For sequence III, the echo time is 32 ms, and with the same echo train length as I and II, the inter-echo spacing is 8 ms. Since the echo train is relatively shifted later in the transverse relaxation decay curve, implying a decrease in signal to noise ratio, the bandwidth is reduced to maintain uniform amplitude (although the phase calculation only involves the argument data from sequences II and III). Assuming the receive bandwidth is halved to 31.25 kHz, the sampling time and acquisition time are 16 μs and 2.048 ms, respectively. Overall, the single inter-echo spacing of 8 ms is appropriate, resulting in a gradient strength of 5.73 mT / m. The gradient slew rates for sequences I, II, and III are all small, which is advantageous in avoiding the effects of eddy currents, particularly in the argument scan.

[0356] Although the shortest possible echo times are used in sequences I, II, and III above to achieve PDW imaging and reasonably accommodate the inherent parameter inter-relationships of magnetic resonance scanning, they do not help to shorten the imaging time. Since the TR is long, the acquisition time of the fast spin echo train in one excitation is much less than the TR, and for a limited slice scan only one cycle of acquisition is needed. With a phase encoding step of 128 for the sequence set of this embodiment, the times for sequences II and III are 9 s * 128 / 8 = 2:24 s and 7:12 s, respectively.

[0357] The above merely illustrates the embodiments of the present application and is not intended to limit the present application. The present application can have various modifications and changes for those skilled in the art. Any modification, equivalent replacement, improvement, etc. made within the spirit and principle of the present application shall be included in the scope of the claims of the present application.

Claims

1. A magnetic resonance imaging radio frequency magnetic field distribution measurement and human tissue data inversion method, characterized by, The method comprises the following steps: Step (1) scanning the radio frequency magnetic field measurement data of the imaging body: Step (1.1) performing spin echo scanning on the homogeneous water phantom, and the sequence excitation deflection angle, echo time and repetition time parameters are (90°, TE, TR) respectively to obtain the amplitude image and the argument image data; Step (1.1) performing spin echo scanning on the homogeneous water phantom, and the sequence excitation deflection angle, echo time and repetition time parameters are (90°, TE, TR) respectively to obtain the amplitude image and the argument image data; Step (1.1) performing spin echo scanning on the homogeneous water phantom, and the sequence excitation deflection angle, echo time and repetition time parameters are (90°, TE, TR) respectively to obtain the amplitude image and the argument image data; Step (1.1) performing spin echo scanning on the homogeneous water phantom, and the sequence excitation deflection angle, echo time and repetition time parameters are (90°, TE, TR) respectively to obtain the amplitude image and the argument image data; Step (1.1) performing spin echo scanning on the homogeneous water phantom, and the sequence excitation deflection angle, echo time and repetition time parameters are (90°, TE, TR) respectively to obtain the amplitude image and the argument image data; Step (1.1) performing spin echo scanning on the homogeneous water phantom, and the sequence excitation deflection angle, echo time and repetition time parameters are (90°, TE, TR) respectively to obtain the amplitude image and the argument image data; In each spin echo scanning, the field of view, slice thickness, slice interval, frequency encoding number, phase encoding step and layer number geometry parameters are the same and consistent; Step (1.2) processing the scanning data of the homogeneous water phantom and the imaging human body part to obtain the radio frequency receiving magnetic field sensitive field distribution data and the transmitting magnetic field sensitive field distribution data of the human body tissue in sequence; Step (1.3) processing to obtain the radio frequency receiving magnetic field phase distribution data and the transmitting magnetic field phase distribution data of the human body tissue; Step (2) inverting the conductivity and permittivity from the magnetic field measurement data: Assuming that the magnetic permeability of the human body tissue is constant and equal to the magnetic permeability of vacuum, the distribution of the electrical conductivity and the relative permittivity in the tissue is obtained by the magnetic resonance electrical property imaging method from the transmitting magnetic field or the receiving magnetic field data of the human body; Step (3) inverting the magnetic permeability or the magnetic susceptibility from the magnetic field measurement data and the existing electrical property: It is assumed that the magnetic permeability of the human body tissue varies with space, and the distribution of the magnetic permeability in the tissue is obtained from the receiving magnetic field of the human body or the transmitting magnetic field of the human body combined with the electrical property data obtained, based on the magnetic field partial differential equation; the electrical property data includes the electrical conductivity and the relative permittivity.

2. A magnetic resonance imaging radio frequency magnetic field distribution measurement and human tissue data inversion method according to claim 1, characterized in that, The spin echo sequence adopts a 2-dimensional imaging scanning mode, i.e. using the frequency encoding gradient as the layer selection gradient, or a 3-dimensional imaging scanning mode, i.e. using the phase encoding gradient as the layer selection gradient; When the 2-dimensional imaging scanning mode is adopted, it cannot be ensured that the sensitive field and the phase of the measured magnetic field between different layers are continuous in the longitudinal direction, and a 2-dimensional approximation algorithm is selected when inverting the electrical property and the magnetic property based on the measured magnetic field, i.e. considering that the longitudinal spatial partial derivative of the electromagnetic property is equal to 0; When the 3-dimensional imaging scanning mode is adopted, it can be ensured that the sensitive field and the phase of the measured magnetic field between different layers are continuous in the longitudinal direction, and a 2-dimensional approximation algorithm or a 3-dimensional volume algorithm is selected when inverting the electrical property and the magnetic property based on the measured magnetic field.

3. A method of magnetic resonance imaging radio frequency magnetic field distribution measurement and human tissue data inversion as claimed in claim 1, wherein, The sequences are based on the signal formation mechanism of fast spin echo, not gradient echo, which is advantageous in eliminating the inconsistency distortion of the proton precession distribution caused by the weak inhomogeneity of the main static magnetic field, so that the measured radio frequency magnetic field is only related to the inhomogeneity of the transmitting and receiving coil magnetic field and the tissue non-single substance composition, and is affected by the eddy current effect caused by the gradient coil.

4. A method of magnetic resonance imaging radio frequency magnetic field distribution measurement and human tissue data inversion as claimed in claim 1 or 3, characterized in that, The in vivo eddy current effect caused by the gradient coil is eliminated by the method of implementing two scans with opposite gradient pulse polarity and other sequence parameters unchanged, and the spatial domain second harmonic interference in the phase distribution of the receiving magnetic field and the transmitting magnetic field is significantly reduced by averaging the phase image data of the two scans.

5. A method of magnetic resonance imaging radio frequency magnetic field distribution measurement and human tissue data inversion as claimed in claim 1, wherein, In order to eliminate the inhomogeneity of the magnetic field sensitive field distribution inherent in 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 for the sequence used for scanning the human body to measure the magnetic field distribution, and the related coil field inhomogeneity is eliminated in the data processing for calculating the receiving and transmitting magnetic field amplitudes.

6. A magnetic resonance imaging radio frequency magnetic field distribution measurement and human tissue data inversion method according to claim 1, characterized in that, The transverse relaxation time constant and the longitudinal relaxation time constant of the imaged object respectively control the rate of decay of the transverse component and the rate of recovery of the longitudinal component of the precession magnetization vector, and by keeping the echo time TE and the repetition time TR of the two sequences the same and the flip angle FA different, the amplitude of the magnetization vector of the same voxel in the two scans is different but points to the same direction during the echo signal acquisition.

7. A method of magnetic resonance imaging radio frequency magnetic field distribution measurement and human tissue data inversion as claimed in claim 1, wherein, By keeping the echo time TE of the two scans in a fixed ratio during the sequence execution and keeping the flip angle FA the same, the purpose of keeping the phase difference of the transverse component of the precession magnetization vector constant during the readout of the two scans is achieved, so as to measure the initial phase of the receiving magnetic field; the equal repetition time of the two scans makes the amplitude of the longitudinal magnetic vector equal at the beginning of each sequence period, and the amplitudes of the transverse magnetic vectors remain relevant consistency during the imaging signal acquisition.

8. A magnetic resonance imaging radio frequency magnetic field distribution measurement and human tissue data inversion method according to claim 1 or 6, characterized in that, When the imaged object is scanned with sequence parameters (90°, TE, TR) and (45°, TE, TR) respectively, the equal TR in the front and back scans ensures that the amplitude of the transverse component of the initial magnetization vector and the ending longitudinal component of each k-line acquisition period are equal, and the equal TE ensures that the direction of the transverse component of the vector is always consistent during the precession of the echo acquisition, so that the amplitude part of the precession vector of the two scans can be divided by the scalar to eliminate the proton density coupling.

9. A magnetic resonance imaging radio frequency magnetic field distribution measurement and human tissue data inversion method according to claim 1 or 7, characterized in that, When the imaged object is scanned with sequence parameters (45°, TE, TR) and (45°, TE / 2, TR) respectively, the equal excitation flip angle in the front and back scans determines the specific decay mode of the transverse magnetic vector in a single sequence period, that is, the amplitudes of the two echoes are related, and the relaxation echo time ratio is controlled to be 1 / 2, so that the phase difference of the transverse component of the precession magnetic vector remains constant during the refocusing period of the echo acquisition, thereby eliminating the same excitation radio frequency pulse initial phase distribution contained in the receiving magnetic resonance signal phase, and the remaining phase amount is equal to 1 / 2 of the receiving magnetic field phase of the target scanning sequence; the receiving magnetic field phase is subtracted from the original image signal phase, and then a theoretical pi / 2 constant bias phase amount is subtracted, to obtain the transmitting magnetic field phase.

10. The method of claim 1, wherein, To adapt to the partial saturation phenomenon caused by the fact that the repetition time TR parameter of the measurement sequence set does not satisfy the condition of being much larger than the longitudinal relaxation time of the tissue, the TR is set to be greater than 5 times the transverse relaxation time T2 of the target imaging tissue, so as to ensure that there is no residual transverse magnetic vector coherence at the end of a pulse sequence TR period; And for the imaging scan based on the fast spin echo sequence, in one TR period, the signals before and after the adjacent echo formation are not avoided to be connected end to end, that is, the echo time TE is required to be set to be greater than 5 times the magnetic sensitive transverse relaxation time T2* to make the transverse magnetic vector sufficiently incoherent.

11. A method of magnetic resonance imaging radio frequency magnetic field distribution measurement and human tissue data inversion according to claim 1 or 6 or 8, characterized by, The intermediate process data m is obtained by dividing the amplitude data m_9_p acquired with the sequence (90°, TE, TR) by the amplitude data m_4_p acquired with the sequence (45°, TE, TR) and is free of the tissue proton density weighting but contains the ratio of the emitted magnetic field amplitude sine and the normalized received magnetic field amplitude p where m denotes the amplitude and p stands for the water phantom; The intermediate process data m which eliminates the tissue proton density weight but contains the ratio of the sine of the amplitude of the emitted magnetic field and the normalized received magnetic field is obtained by dividing the amplitude data m_9_h acquired by the sequence (90°, TE, TR) by the amplitude data m_4_h acquired by the sequence (45°, TE, TR) h where h represents the human body; Because of the water phantom homogeneity, the process data m h The magnetic field sensitive field inhomogeneity of the receiving coil in m p is corrected, i.e. the human receiving magnetic field sensitive field where the assumption or approximation that the human magnetic permeability is homogeneous and equal to the water phantom magnetic permeability is included.

12. A magnetic resonance imaging radio frequency magnetic field distribution measurement and human tissue data inversion method according to claim 8, wherein, The human body emitted magnetic field sensitive field is calculated by the following steps, Step (1) dividing the original magnitude data m_9_h acquired in a scan of a human body (90°, TE, TR) by the human body receive magnetic field sensitivity field The intermediate process quantity magnitude data t1 is calculated; m represents the amplitude, and h represents the human body; Step (2) dividing the original magnitude data m_4_h acquired in a scan of a human body sequence (45°, TE, TR) by the human body receive magnetic field sensitivity field An intermediate process quantity magnitude data t2 is calculated; The data t1 and t2 both contain the contributions of the macroscopic magnetization, that is, the proton density and the sine value of the deflection angle related to the emitted magnetic field amplitude; Step (3) dividing the intermediate quantities t1 and t2, obtaining Taking the inverse cosine of this ratio gives the distribution of the target excitation transmit magnetic field sensed inside the imaged human tissue where γ is the gyromagnetic ratio of the target nucleus of the magnetic resonance imaging, τ is the time domain width of the excitation transmit frequency pulse, is the uniform transmit magnetic field amplitude, and the superscript 45 denotes the angle of the excitation flip angle.

13. A magnetic resonance imaging radio frequency magnetic field distribution measurement and human tissue data inversion method according to claim 11 or 12, characterized in that, The received magnetic field sensitive field is a magnetic field distribution generated when a unit current flows through the receiving coil, the amplitude of which ranges in the range of (0.5, 1.5) Tesla; the calculated amplitude of the emitted magnetic field sensitive field is of the order of 10 - 6 Tesla.

14. A magnetic resonance imaging radio frequency magnetic field distribution measurement and human tissue data inversion method according to claim 1 or 12, characterized by, The implemented parameters are the excitation deflection angle selected in the (45°, TE, TR) sequence between 0° and 90°, which ensures that the quantity ratio t1 / t2 is positive when calculating the inverse cosine function value, and the emitted magnetic field sensitive field value falls within the range (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 magnetic resonance imaging radio frequency magnetic field distribution measurement and human tissue data inversion method according to claim 11 or 12, characterized in that, When the sequence set parameters (90°, TE, TR) and (45°, TE, TR) are used to scan the human body, the center value of the histogram distribution of the received magnetic field sensitive field voxel amplitude may deviate seriously from 1 and the center value of the histogram distribution of the cosine value of the emitted magnetic field sensitive field amplitude may deviate seriously from 0.707 and approach or even exceed 1, so the sequence set parameters (120°, TE, TR) and (60°, TE, TR) are used to improve and adjust the calculation of the measured magnetic field amplitude distribution to a reasonable statistical range, without affecting the applicability and reasonableness of the magnetic field amplitude distribution measurement method and without affecting the measurement of the magnetic field phase distribution.

16. A method of magnetic resonance imaging radio frequency magnetic field distribution measurement and human tissue data inversion according to claim 1 or 7 or 9, characterized by, The received coil human body radio frequency magnetic field phase is calculated by the following steps, Step (1) Subtract the radian data a_t_p acquired by the sequence (45°, TE / 2, TR) from the radian data a_4_p acquired by the scan uniform water phantom sequence (45°, TE, TR) to obtain the phase difference Δp p ; a represents the radian, and p represents the water phantom. Step (2) subtract the acquired radian data a_t_h of the sequence (45°, TE / 2, TR) from the acquired radian data a_4_h of the sequence (45°, TE, TR) to obtain the phase difference Δp h ; Step (3) calculates the induced received magnetic field phase in the human tissue for the target sequence (45°, TE, TR) as In the formula, the argument data a represents the angle; t represents the sequence of the short echo time TE / 2, which is different from the sequences (90°, TE, TR) and (45°, TE, TR); and the phase data of the scanned human body is subtracted from the phase data of the scanned homogeneous water model to correct the non-uniformity of the received coil magnetic field phase distribution.

17. A method of magnetic resonance imaging radio frequency magnetic field distribution measurement and human tissue data inversion according to claim 1 or 7 or 9, characterized by, The human body induced radio frequency emitted magnetic field phase is calculated by the following steps: Step (1) The phase of the induced received magnetic field in the human tissue corresponding to the target sequence (45°, TE, TR) Subtracting the original phase data a_4_h acquired by scanning the target sequence to obtain the intermediate process quantity phase data p M , which is the initial transverse component M 0,⊥ of the magnetization vector; Step (2) subtracts π / 2 from the phase p M Subtracting π / 2 from the phase of the human-induced transmitted magnetic field 18. A magnetic resonance imaging radio frequency magnetic field distribution measuring and human tissue data inverting method according to claim 6 or 7 or 8 or 9 or 11 or 12 or 16 or 17, characterized in that, The flip angles of the refocusing radio frequency pulses of the magnetic field measurement sequence are all 180°, which ensures that the precession magnetic vector component is symmetrical around the reference average rotation speed magnetic vector position in the transverse plane before and after the flip.

19. The magnetic resonance imaging radio frequency magnetic field distribution measurement and human tissue data inversion method according to claim 11 or 12 or 16 or 17, characterized in that, When the sequence with the echo time parameter TE is used for scanning, the sampling data interval of each echo is set to be td, N data are collected for each echo, the echo reception time tacq is N*td, and it is assumed that the received signal signal-to-noise ratio at this time is rSNR. When scanning with a sequence with echo time parameter TE / 2, if the sampling data interval of each echo is set as td / 2 accordingly, and N data are collected for each echo, the echo receiving time tacq is N*td / 2, at this time, the received signal-to-noise ratio will change significantly, i.e., decrease; In order to keep the signal-to-noise ratio of the two scans consistent, i.e., the transverse component of the magnetic vector under the action of the same excitation deflection angle pulse starts from the same amplitude and decays according to the same curve, the bandwidth BW of the sequence (45°, TE / 2, TR) is adjusted to adjust td indirectly; since the single echo receiving 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 number of echo sampling data points N and tacq.

20. A magnetic resonance imaging radio frequency magnetic field distribution measurement and human tissue data inversion method according to claim 1, wherein, Based on the electromagnetic field and Maxwell equations, the magnetic property parameter distribution is inverted by two different methods, which are: The first type of magnetic property inversion method: first, assume that the human body permeability is homogeneous and equal to the vacuum permeability μ0, use the existing magnetic resonance electrical property imaging inversion conductivity and permittivity algorithm, and calculate the complex permittivity distribution of the tissue using the measured transmitted or received magnetic field; Then consider that the permeability of the same imaging human body is non-homogeneous, use the remaining measured magnetic field data, or measure new radio frequency magnetic field quantity data again with different pulse sequence scanning parameters, and invert the permeability under the condition that the electrical property parameter value distribution of the tissue is known; The second type of magnetic property inversion method: under the condition that the human body permeability is also non-homogeneous, construct a partial differential equation with multiple sets of magnetic field measurement quantities as known quantities, and the permeability, conductivity and relative permittivity as unknown quantities, and simultaneously invert the related magnetic and electrical properties.

21. A magnetic resonance imaging radio frequency magnetic field distribution measurement and human tissue data inversion method according to claim 20, wherein, The first type of magnetic property inversion method is to obtain the electrical conductivity and relative permittivity distribution parameters of the human body, use the remaining magnetic field measurement data, or measure new available radio frequency magnetic field quantities by changing the sequence parameters on the same imaging body, combine the inverted electrical property parameters, and invert the human tissue permeability μ according to the following equation: where ω is the resonant Larmor angular frequency, where ω is the resonant Larmor angular frequency, is the Laplace operator, is the gradient operator, and 'x' is the curl operator; the assumption that the human body magnetic permeability is equal to the vacuum magnetic permeability is contained twice to simplify the scalar and vector derivations to apply the Gauss magnetic law, obtaining the core equation to invert the magnetic permeability from the measured magnetic field.

22. A magnetic resonance imaging radio frequency magnetic field distribution measurement and human tissue data inversion method according to claim 20, wherein, The second type of magnetic property inversion method is to simultaneously invert the electrical and magnetic properties based on the Maxwell equations, i.e., to invert the three unknowns of permeability, conductivity and relative permittivity using known magnetic field measurement quantities, and to derive the following equation group: wherein is the Laplacian operator, 'v' is the inverse magnetic permeability, is the complex permittivity, containing the electric conductivity and the relative permittivity; is the measured transmit magnetic field excitation amplitude, is the measured receive magnetic field sensitive field conjugate; Further investigations Item, B is or The fluid dynamics model can be expanded into convection terms and diffusion terms to obtain the core calculation equation set for the magnetic field components: Two independent sets of known field quantity data are obtained by changing the pulse sequence parameters twice, a high well-posed equation group containing four equations is constructed based on the core equation form, and the unknown electromagnetic property parameter distribution data is solved.

23. A magnetic resonance imaging radio frequency magnetic field distribution measurement and human tissue data inversion method as claimed in claim 20 or 21, characterized in that, In deriving the equation for calculating the magnetic permeability of class I under the condition of magnetic resonance 2-dimensional tomography scanning, the z-component B z of the magnetic field emitted by the excitation coil and the magnetic field received by the detection coil is neglected, which is parallel to the direction of the main static magnetic field. z And when solving the derivative of the electrical property in a specific layer, it is assumed that the z-gradient is 0, i.e. z B 24. A magnetic resonance imaging radio frequency magnetic field distribution measurement and human tissue data inversion method as claimed in claim 21 or 22, characterized in that, The first-order partial derivative and the second-order partial derivative of the radio frequency magnetic field measurement quantity in each direction are obtained by fitting based on the Savitzky-Golay filter construction principle method; for a 3x3 operation kernel data point array, the complex value of the measurement magnetic field B at the central position, i.e., the coordinate point (x, y), is the fitting target quantity, and the new value obtained by fitting is: where c1to c6are fitting coefficients, (x i ,y i ) are the coordinates of the neighboring field points, i = 1, 2, 3, 4, 5, 6, 7, 8. After implementing this quadratic parabolic numerical fit, the magnetic field gradient and second order partial derivatives at the target point (x5, y5) are calculated by: In the formulae, denotes the partial derivative in space.

25. A magnetic resonance imaging radio frequency magnetic field distribution measurement and human tissue data inversion method as claimed in 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 longitudinal relaxation time longest tissue component T1 to satisfy the signal intensity SI calculated based on the spin echo formation mechanism SE The condition that the tissue and sequence parameter related contrast can be factored; the expression of SI written from the Bloch equation is: SE SI = M0 * (1 - e-aTR) In the formula, 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: Thus, the contrast is achieved in the calculation of the measured magnetic field amplitude and the received magnetic field is only related to sin α(r) and S(r), respectively.

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