Magnetic resonance imaging electromagnetic parameter distribution scanning and calculating method

Abstract

The invention relates to a magnetic resonance imaging electromagnetic parameter distribution scanning and calculation method, which comprises the following steps that a fast spin echo sequence I is adopted to scan a target body, parameters are (alpha1, TE0, TR0 and Bw1), alpha is a deflection angle, TE is echo time, TR repetition time and Bw receiving bandwidth; scanning a target body by adopting a fast spin echo sequence II, wherein parameters are (alpha2, TE0, TR0 and Bw1); scanning a target body by adopting a fast spin echo sequence III, wherein parameters are (alpha2, TE0, TR0 and Bw2); and scanning the homogeneous water model by adopting the fast spin echo sequences I, II and III. And under the condition that the magnetic vector precession meets the conditions that the amplitudes are comparable when the phases are the same and the phases are differentiable when the amplitudes are the same, each magnetic field and electric field parameter is calculated based on the homogeneous water model data. According to the invention, decoupling of the proton density and the sensitive field of the signal is realized, so that magnetic resonance weighted quantization imaging becomes possible.

Description

Technical Field

[0001] This invention belongs to the field of magnetic resonance imaging technology, specifically relating to a method for scanning and calculating electromagnetic parameter distribution in magnetic resonance imaging. Background Technology

[0002] Currently, magnetic resonance imaging (MRI) is one of the most important medical imaging technologies, characterized by non-ionizing radiation, non-invasiveness, and high contrast of human tissues. Advances in superconducting magnet technology have driven the development of MRI systems towards increasingly higher field strengths. However, due to the dielectric properties of tissues, ultra-high field strengths lead to problems such as shorter wavelengths and poorer uniformity of the radiofrequency magnetic field distribution, significantly affecting the quality of reconstructed images. How to effectively measure the distribution of the excitation and reception magnetic fields, especially the specific holographic amplitude and phase distributions, based on the characteristics of tissue electromagnetic parameters, remains an unsolved problem in nuclear magnetic resonance (spectroscopy and imaging) since its discovery in the 1940s. Summary of the Invention

[0003] To address the difficulty in measuring the distribution of radio frequency magnetic fields and related derived parameters in existing magnetic resonance imaging (MRI) methods, this invention provides a method for scanning and calculating the electromagnetic parameter distribution in MRI, and a signal detection method for measuring radio frequency magnetic fields and related derived parameters. This method includes not only the measurement of the amplitude and phase distribution of the excitation transmitted magnetic field, but also the measurement of the amplitude and phase distribution of the detected received magnetic field.

[0004] To achieve the above objectives, the technical solution adopted by the present invention is as follows:

[0005] A method for scanning and calculating electromagnetic parameter distribution in magnetic resonance imaging, comprising:

[0006] Step 1: Perform a fast spin echo first sequence I scan of human tissue. The parameters of the first sequence I are (α1, TE0, TR0, Bw1), where α1, TE0, TR0, and Bw1 are the deflection angle, echo time, repetition time, and receiving bandwidth of the first sequence I, respectively.

[0007] Step 2: Perform a fast spin echo second sequence II scan of human tissue. The parameters of the second sequence II are (α2, TE0, TR0, Bw1), where α2, TE0, TR0, and Bw1 are the deflection angle, echo time, repetition time, and receiving bandwidth for performing the second sequence II, respectively.

[0008] Step 3: Perform a fast spin echo third sequence III scan on human tissue. The parameters of the third sequence III are (α2, TE0, TR0, Bw2), where α2, TE0, TR0, and Bw2 are the deflection angle, echo time, repetition time, and receiving bandwidth for performing the third sequence III, respectively.

[0009] Among them, the TE and TR remain consistent between sequences, ensuring the consistency of magnetic vector precession amplitude and phase between different scans for the same target, thus creating the principle conditions for the measurement and calculation of radio frequency magnetic field. In order to make the dynamic range large when calculating the amplitude of the transmitted magnetic field in the later stage, and considering the rationality of the system gradient pulse intensity, the parameters can be set to include α1=120°, α2=60°, Bw1=62.5kHz, and Bw2=31.25kHz.

[0010] To eliminate sequence-dependent contrast weights in the signal data and the effects of coil magnetic field inhomogeneity when calculating the magnetic field distribution in human tissue, the homogeneous water model was scanned again with the same sequence set:

[0011] Step 4: Perform the first sequence I scan of the homogeneous water model using the fast spin echo;

[0012] Step 5: Perform the fast spin echo second sequence II scan on the homogeneous water model;

[0013] Step 6: Perform the fast spin echo third sequence III scan of the homogeneous water model;

[0014] Step 7: Using the original amplitude and argument image data of the equivalent spin density of the human tissue and homogeneous water phantom obtained by scanning, calculate the distribution data of the radio frequency magnetic field and related derived parameters, including the magnetic field strength of the radio frequency transmitting coil, the proton density of the target body, the complex capacitance of the target body, the magnetic permeability of the excited target body, the magnetic permeability of the detected target body, the transmitted electric field strength, the specific absorptivity of the target body, the magnetic field strength of the radio frequency receiving coil, the magnetic field strength of the receiving coil, the transmission-reception phase of the receiving coil, the magnetization intensity of the detected target body, and the continuous phase of the transmission-reception of the magnetic flux density.

[0015] In the aforementioned sequence set, the echo time parameter TE is sufficiently short, for example, less than the shortest transverse relaxation time T2 in the effectively imaged human tissue component, and the repetition time parameter TR is sufficiently long, for example, greater than three times the longest longitudinal relaxation time T1 of the imaged tissue component, to ignore the unnecessary influence of the potential exponential nonlinear contrast weighting factor related to TE and TR. This makes the calculation of the magnetic field distribution based on the image data reconstructed by the signal more reasonable, which is beneficial for separating the contributions of deflection angle weight and other parameter weights in the calculation of magnetic resonance signals, forming independent factors for magnetization, deflection angle sine, (TE, TR) unit received magnetic field, and sensitive field, thereby improving the accuracy of the measurement method; Sequence I and Sequence... In sequence II, the equal receiving bandwidth ensures that the echo acquisition time is equal, that is, the precession phase between sequences remains consistent. Therefore, under different excitation deflection angles, magnetic vector amplitude division between sequences can be performed to decouple the proton density and coil sensitive field amplitude in the signal. In terms of electromagnetic parameter phase calculation, the equal deflection angles of sequence II and sequence III and the halved receiving bandwidth ensure that the target body undergoes transverse relaxation attenuation according to the same eigenvalue curve in one echo chain cycle, that is, the echo peak amplitude is the same. Therefore, under the same deflection angle, magnetic vector phase subtraction between sequences can be performed. Based on the bandwidth ratio, the single phase accumulation can be calculated, as well as the phase of the transmitting magnetic field and receiving magnetic field, and the phase of the human tissue and receiving coil field can be calculated discretely.

[0016] Below, deflection angle and Relationship satisfaction Given that the receiving bandwidths Bw1 and Bw2 are related by Bw1=2Bw2, write out the overall expression of the magnetic resonance signal containing parameters for different imaging targets.

[0017] Targeting human tissues:

[0018] First sequence I parameters ( Signals acquired during scanning (TE0, TR0, Bw1):

[0019] ;

[0020] The subscript 's' indicates human tissue. Represents the coordinates of human tissue. Indicates the coordinates of the transmitting coil. The coordinates of the receiving coil are indicated to distinguish the target object from the coil; ER represents the echo time TE0 and repetition time TR0; P represents the proton density; and the longitudinal relaxation time T1 and transverse relaxation time T2 are collectively related to the contrast of the reconstructed image; the initial phase of the excitation vector is also indicated. Also related to the initial phase of the radio frequency pulse of the transmitting coil Related, but in magnetic resonance imaging It typically does not change with the sequence, therefore it is unlabeled; It is the gyromagnetic ratio; M0 represents the radio frequency excitation / emission magnetic flux density; M0 represents the magnetization vector. The phase of the received magnetic field is measured in units. Indicates the resonant angular frequency. Indicates the pulse width of the stimulus. This indicates the receiving magnetic field related to TE and TR.

[0021] Second sequence II parameters ( Signals acquired under TE0, TR0, Bw1):

[0022] ;

[0023] ;

[0024] Third sequence III parameters ( Signals acquired under (TE0, TR0, Bw2):

[0025] ;

[0026] For homogeneous water models:

[0027] The signal acquired by the first sequence I parameter:

[0028] ;

[0029] In the formula, the subscript p represents the water model. Because the water model is homogeneous, the signal contribution is dominated by the distribution of the sensitive field of the coil magnetic field.

[0030] Signal acquired by the second sequence II:

[0031] ;

[0032] And the signals acquired in Sequence III:

[0033] .

[0034] In the above formula, for a homogeneous water model, all parameters are also uniformly distributed, so there is a relevant... That is, the parameter values ​​at each position remain unchanged.

[0035] To facilitate the calculation of the magnetic field and its derived parameters, the focus is on examining the amplitude in each set of signal data. and phase Two components. The signal components of the human tissue and the homogeneous water model are re-recorded as shown in Table 1.

[0036] Table 1

[0037]

[0038] Since the amplitude and phase of the emitted magnetic field determine the deflection angle and initial phase of the transverse magnetization vector, respectively, and the amplitude and phase of the radio frequency excitation pulses implemented on each slice in the two-dimensional imaging sequence scan are equal and consistent, the emission-related parameter distribution can be reconstructed in three dimensions based on the data measured and calculated in two-dimensional slices, including the use of three-dimensional inversion algorithms when calculating electrical characteristics.

[0039] Since the amplitude and phase of the received magnetic field depend on the magnetization and sequence parameters, and the magnetic field gradient pulses related to the sheet in the two-dimensional scan are either time-divisionally acquired with the echo or have a mean intensity of zero, and the echo time and repetition time are consistent, the distribution of the received calculation parameters can be reconstructed in three dimensions, including the use of a three-dimensional inversion algorithm when calculating electrical characteristics.

[0040] Based on the original equivalent spin (proton) density image data reconstructed from magnetic resonance imaging, and according to the basic principles of vector amplitude division and phase difference, the amplitude and phase distribution data of each magnetic field and derived parameter are calculated sequentially, and the deflection angle is set... , and Bw1 Bw2 The solution process is as follows.

[0041] (1) Amplitude of magnetic permeability received by human tissue:

[0042] For imaging human tissue, the amplitude data of the scan is acquired by the first sequence I. Divide by the amplitude of the second sequence II acquisition scan ,have:

[0043] ;

[0044] Based on the definition of the receiving magnetic field, and under the conditions of the same target body and the same sequence time parameters, we have:

[0045] ;

[0046] That is, only includes the excitation deflection angle The relevant functional weights and contrast related to TE0, TR0, and the sensitive field of the receiving coil are also considered. Furthermore, to distinguish the signal contributions of the target body and the coil pair in the interaction, the magnetic flux density is written as... , That is, the distribution of the sensitive field, we have:

[0047] ;

[0048] In the formula, This represents the receiving magnetic field related to TE and TR, including the sensitive field distribution of the receiving coil. '~' indicates that it is calculated based on the receiving magnetic field. Similarly, the amplitude data of the first sequence I of the homogeneous water model is acquired and scanned. Divide by the amplitude data acquired by the second sequence II scan ,have:

[0049] ;

[0050] In the formula, The permeability of the homogeneous water model is calculated based on the received magnetic field.

[0051] From the perspective of the emitted magnetic field, assuming that the magnetic permeability of human tissue is uniform and equal to that of a homogeneous water phantom, then we can further reduce the magnetic permeability by... The relative permeability of human tissues is obtained by weighting relevant functionals.

[0052] ;

[0053] remember:

[0054] .

[0055] (2) Amplitude of the magnetic field (magnetic flux density) emitted by human tissue:

[0056] Given the received (relative) permeability amplitude mentioned above, further calculations can be performed based on the magnetic resonance acquisition signal formula:

[0057] ;

[0058] ;

[0059] In the formula, The T in the text indicates that it is related to TE and TR. This indicates the magnetic field strength under specific TE and TR conditions.

[0060] To obtain the amplitude of the emitted magnetic field, perform:

[0061] ;

[0062] Among them, the excitation deflection angle must meet the following requirements. .

[0063] In actual calculations, to reduce accuracy loss, the following can be executed directly:

[0064] ;

[0065] have:

[0066] ;

[0067] ;

[0068] The transmitting magnetic field is the result of the combined effect of human tissue and the sensitive field of the transmitting coil.

[0069] (3) Amplitude of the sensitive field of the receiving coil:

[0070] To avoid introducing new assumptions, and drawing on the principle of calculating the amplitude of the magnetic field emitted by human tissue, a deflection angle is used when scanning the homogeneous water model. The amplitude of the emitted magnetic field is calculated as follows:

[0071] ;

[0072] Taking advantage of the homogeneity of the water model, the amplitude data of the homogeneous water model were further scanned using the second sequence II. Divide by :

[0073] ;

[0074] ;

[0075] In the formula, The superscript '0' in the equation indicates that the magnetization amplitude of the homogeneous water model is uniform. Further derivation and substitution of both sides of the above equation yields:

[0076] ;

[0077] This involves calculating the sensitive field amplitude of the receiving coil, weighted by the magnetization intensity of a specific water model, and relating it to parameters TE0 and TR0. The magnetization weighting also implies a weak correlation with the uniformity of B0.

[0078] (4) Amplitude of magnetic field strength of transmitting coil:

[0079] Directly determined by the launch deflection angle of the water model ,have:

[0080] ;

[0081] ;

[0082] ;

[0083] In the formula, Indicates deflection angle The magnetic field strength of the transmitting coil under the given conditions The permeability of the water model is calculated based on the emission magnetic field.

[0084] (5) Amplitude of magnetic field (magnetic flux density) received by human tissue:

[0085] The amplitude of the magnetic permeability of the human tissue and the amplitude of the sensitive field of the receiving coil, calculated from (1) and (3) respectively, can be used to indirectly calculate the amplitude of the magnetic field received by the tissue:

[0086] ;

[0087] ;

[0088] It can be seen that the amplitude of the magnetic field received by the tissue, calculated from the scan data, contains the magnetization intensity weight of the homogeneous water model.

[0089] (6) Amplitude of magnetic permeability emitted by human tissue:

[0090] Since the aforementioned calculations (2) and (4) yielded the emitted magnetic field of human tissue and the sensitive field of the emitted coil, the permeability related to the emitted magnetic field can be directly calculated according to the definition:

[0091] ;

[0092] (7) Magnetization intensity amplitude of human tissue:

[0093] Based on the holographic expression of magnetic resonance signals containing amplitude and phase, several principles need to be considered when calculating the magnetization amplitude of human tissue:

[0094] 1) It should be calculated based on data generated from sequences of scanned human tissues;

[0095] 2) The calculated amplitude is only related to the data from the first sequence I scan of the homogeneous water model and the second sequence II scan of the human tissue. The magnetization amplitude should be calculated from the II data based on the amplitude of the received magnetic field.

[0096] 3) The permeability of human tissue and the amplitude of the sensitive field of the coil have been obtained, so the synthesized receiving magnetic field can be used in the calculation.

[0097] The expression for amplitude data obtained by scanning human tissue using sequence II is known to be:

[0098] ;

[0099] The relative permeability of human tissue receiving magnetic fields, the sensitive field of the receiving coil, and the sinusoidal quantities related to the transmitting magnetic field are all known.

[0100] ;

[0101] ;

[0102] Substitute the obtained quantities, i.e., the measurement data items, into the table.

[0103] ;

[0104] ;

[0105] The normalized magnetization amplitude is obtained:

[0106] ;

[0107] In the formula, The magnetization intensity amplitude of the homogeneous water model.

[0108] Generally, the magnetization intensity of human tissue is less than or equal to the magnetization intensity of a homogeneous water model. Therefore, under the condition that the gain of the system scanning receiver is the same, the above-mentioned normalized magnetization intensity distribution takes a value less than or equal to 1.

[0109] (8) Proton density of human tissue:

[0110] In quantum mechanics, based on magnetization and proton density Definition of relationship

[0111] ;

[0112] In the formula, To reduce Planck's constant, For spin quantum number, Boltzmann's constant, The temperature is the temperature of human tissue. As described in the background section, the magnetization of pure water in 3T is approximately 9.4 mA / m, from which (7) can be calculated. If the approximate distribution is obtained, then:

[0113] ;

[0114] The above is the derivation of the calculation formulas for various amplitude-related parameters in the electromagnetic measurement problem of magnetic resonance imaging. Since the signal-to-noise ratio of measured data is not ideal, it poses certain challenges to the calculation of radio frequency magnetic fields and derived parameters, and even the subsequent inversion of electromagnetic properties. In engineering, a reasonable range of sequence parameters can be set in conjunction with the amplitude data processing procedure. For example, when the deflection angle is 60°, the dynamic range of the measurement data can be [30°, 90°], which corresponds to the median pixel value in the image. The value is multiplied by 1, and the pixels outside the interval are re-interpolated.

[0115] Apart from the receiving bandwidth, the excitation deflection angle, echo time, repetition time, geometric definition, and other echo and timing parameters are identical between the second sequence II and the third sequence III, including the initial phase of the RF pulse, which remains the same (manufacturer's underlying settings). Therefore, for the same imaging target, with the relaxation time constant and the target position unchanged, the difference in the echo signal sampling process during the two sets of sequence scans lies in the constant difference between the precession phase distributions, which manifests in the time domain waveform as an increase in echo duration and a decrease in echo peak value when the receiving bandwidth is narrow.

[0116] (9) Phase of the magnetic field received by human tissue:

[0117] To calculate the reconstructed received magnetic field phase distribution for the second sequence II of the imaging scan, the phase data of the second sequence II is subtracted from the phase data of the third sequence III of the human tissue to obtain the phase data under given TE and TR conditions:

[0118] ;

[0119] The derivation of the formula is based on the fact that the radio frequency pulses implemented in the two sets of scans are the same, that is, the deflection angle generated by the transmitted magnetic field acting on the magnetization vector between sequences is the same, and the initial transverse precession phase is also the same, which are respectively equal to BW. and Under these conditions, a single-fold phase difference of the receiving magnetic field is directly obtained:

[0120] ;

[0121] In actual phase data acquisition, it is usually necessary to first set up the system to save the image phase data of a single channel of the receiving coil, and then perform additional vector combining in post-processing to obtain the true Fourier phase data. It is also important to be aware of the potential entanglement characteristics of the phase data; therefore, it is essential to ensure that the raw data is untangled using third-party calculation tools before proceeding with subsequent data processing.

[0122] (10) Phase of the receiving coil's sensitive field:

[0123] Similar to the contribution of the coil-sensitive field to the received magnetic field amplitude, calculations are performed on the water model data to measure the phase component of the receiving coil-sensitive field:

[0124] ;

[0125] In the formula, because the initial phase of the radio frequency pulse in magnetic resonance imaging does not affect the specific precession imaging method as much as its amplitude, the phase of the magnetization is... Can be changed to .

[0126] The property of addition and subtraction based on phase can be remembered as follows: ,in The phase of the complex permeability of human tissue. Let be the frequency domain phase of the coil's sensitive field, i.e., the magnetic field strength, which is related to the coil's geometry. Further, let the coordinates of the homogeneous water model be denoted... ,have:

[0127] ;

[0128] In the formula, The characteristics representing the target entity itself, that is, equivalent to In the overall received magnetic field phase, it is important to reflect the phase differences between different tissues, while the phase of the receiving coil's sensitive field shows only the DC component bias. As before, because the receiving bandwidth increases exponentially, the formula for calculating the phase distribution data of the receiving coil's sensitive field in Sequence II is:

[0129] ;

[0130] (11) Phase of the magnetic field emitted by human tissue:

[0131] The study takes the deflection angle as the basis. Sequence II scanning of human tissue, based on the principle of phase continuity between transmission and reception in the voxel signals reconstructed from the acquired signals by magnetic resonance imaging, yields the following relationship:

[0132] ;

[0133] Then the phase of magnetization:

[0134] ;

[0135] ;

[0136] ;

[0137] The phase of the emitted magnetic field of human tissue is:

[0138] ;

[0139] (12) Phase of the sensitive field of the transmitting coil:

[0140] The phase of the permeability related to the transmission and reception of a homogeneous water model can be considered constant, and theoretically, both... and The distribution is identical, meaning it is strictly consistent. The phase of the receiving coil's sensitive field has been calculated (10), similar to the calculation method for the phase of the emitted magnetic field of human tissue. The phase data of the transmitting coil's sensitive field can be calculated from the phase data of the water model scanned by sequence II, as follows:

[0141] ;

[0142] ;

[0143] ;

[0144] In the formula, L represents the phase change of the emitted magnetic field strength in a non-rotating coordinate system. Related; three sets of sequences The same conditions are met, ensuring that the phase distribution of the sensitive field of the transmitting coil can be measured and calculated according to the method described above. Simultaneously, it is also necessary to ensure that the initial phase of the voltage waveform of the time-domain excitation magnetic field of the RF source in the magnetic resonance system, under pulse modulation, remains constant regardless of the different sequences used, and does not change with variations in waveform amplitude. Rearranging the above equation, we obtain:

[0145] ;

[0146] (13) Transceive phase of human tissue:

[0147] The radio frequency transmit-receive phase of the target signal is a crucial concept in the spin-based scanning mechanism of magnetic resonance imaging (MRI), and arguably the only physical quantity that can be directly measured without computation. It maintains phase continuity during the precession process between excitation and detection. The transmit-receive phase is defined as follows: Starting from the reconstructed image phase data acquired from the second sequence II,

[0148] ;

[0149] This leads to the derivation of the human tissue transmit-receive phase as follows:

[0150] ;

[0151] ;

[0152] (14) Phase of magnetic permeability received by human tissue:

[0153] The phase of the complex permeability of human tissue at the Larmor frequency, calculated based on the received signal, is obtained directly from the phase of the magnetic field received by (9) human tissue and the phase of the sensitive field of (10) receiving coil.

[0154] ;

[0155] ;

[0156] (15) Phase of emitted magnetic permeability of human tissue:

[0157] The phase of the complex permeability of human tissue at the Larmor frequency calculated based on the emission excitation is obtained directly from the phase of the emitted magnetic field of (11) human tissue and the phase of the sensitive field of the transmitting coil of (12):

[0158] ;

[0159] ;

[0160] There are no assumptions made in the calculation of the phase data above. However, it should be noted that there is an intersection effect between the effective area of ​​the water model and the target image in the two phases of the magnetic permeability of human tissue.

[0161] Below, based on the measured and calculated magnetic fields emitted and received by human tissue and the sensitive field of the coil, the distribution of derived parameters such as the electrical properties of the tissue, the intensity of the induced electric field in space, and the specific absorption rate are further calculated according to Maxwell's equations.

[0162] (16) Electrical properties of human tissue (Helmholtz):

[0163] The magnetic properties of human tissue, namely the permeability distribution, have already been calculated. Under high-frequency conditions, the electrical properties of human tissue include conductivity and permittivity, and the classical electromagnetic wave equation provides the relationship between electromagnetic waves and electromagnetic properties. The derivation of conductivity calculation using the circular polarization component of the magnetic field is then presented. and permittivity The formula.

[0164] ;

[0165] Continue to calculate the curl on both sides of the equation.

[0166] ;

[0167] ;

[0168] Making a first assumption of homogeneous electrical properties, we get:

[0169] ;

[0170] Will Substituting into the above formula,

[0171] ;

[0172] In the formula, As a sensitive field source for the coil, Let be an orthogonal unit vector in the coordinate system, and the parameter to be determined is the complex permittivity. In magnetic resonance imaging, the magnetic field used to measure the signal is a transverse magnetic field, which is negligible. And the left side of the above equation:

[0173] ;

[0174] ;

[0175] Under ideal tomographic imaging conditions, assuming the target plane is close to the center of the coil and considering... and The unit vector is omitted, and the right side of the above equation has:

[0176] ;

[0177] In the formula, if the sensitive field of the coil can satisfy the second-order or even higher-order continuity conditions, then the order of differentiation of the magnetic field in the components can be appropriately interchanged. The formula can be written in component form:

[0178] ;

[0179] Note that under magnetic resonance conditions, all the above physical quantities are taken as periodic complex values, and the sensitive field and magnetic flux density in pulse excitation and inductive detection are known to be expressed as follows:

[0180] and ;

[0181] In the formula, '+' and '~' represent the forward transmission and reverse reception propagation of the right-hand component in the rotating coordinate system, respectively. and Equivalent. Because the launch-related parameters in the aforementioned calculations have no weighting coefficients and do not involve issues such as the intersection of human tissue and the homogeneous water model target area, for the sake of accuracy, the formula is further processed based on the launch field quantity, resulting in:

[0182] ;

[0183] ;

[0184] Given measured data or a reference coordinate system, the solution to the alternating problem can be transformed from a rectangular coordinate system to a rotating coordinate system. This is equivalent to keeping the real x-component unchanged and multiplying the y-component by the imaginary unit. Therefore, the field coordinate components in the formula and the following derivation are all real values ​​for ease of derivation. Furthermore, we have;

[0185]

[0186] At this point, the known measurement or calculation quantity is the sensitive field of the transmitting coil. Inductive emission magnetic field from human tissue The parameters to be determined are All are complex numbers. Based on the correspondence between the real and imaginary parts:

[0187] ;

[0188] To reduce the accuracy loss in division operations when performing finite difference operations, the first and second derivative data can be obtained by approximating the parameter distribution with a parabolic surface based on the Savitzky-Golay filter and then using the least squares fitting method.

[0189] Induced electric field strength:

[0190] In magnetic resonance imaging (MRI), the excitation pulse emission magnetic field induces the absorption of radiofrequency energy within the tissue. Therefore, according to Faraday's law, the electric field can be calculated from the coil sensing field sampled from a homogeneous water phantom and the tissue's electrical properties. Since the emission magnetic field is a circularly polarized transverse magnetic field, the induced electric field is mainly... The components are:

[0191] ;

[0192] ;

[0193] Above, the field quantities are defined in a suitable complex plane such that each component of the field can take real values. Under reasonable approximation conditions, directly taking the z-component of the equation yields:

[0194] ;

[0195] Known excitation deflection angle and pulse width amplitude under conditions and phase Based on its relationship with the components in the rectangular coordinate system:

[0196] ;

[0197] Export:

[0198] ;

[0199] Ultimately, it can be calculated that:

[0200] ;

[0201] (18) Specific absorption rate

[0202] Besides electric field strength, calculating the specific absorptivity of a target object from an alternating magnetic field requires knowledge of the electrical conductivity distribution of human tissue. The specific absorptivity is defined as follows:

[0203] ;

[0204] In the formula, For electrical conductivity, The electric field is caused by the magnetic field. For mass density. When proton density ,Right now When the SAR value is known, the equivalent SAR value is:

[0205] ;

[0206] That is, the proton density obtained by the actual measurement calculation according to the present invention can yield a SAR value distribution that is close to the true value. .

[0207] Table 2 below summarizes the calculations of various static and alternating electromagnetic physical quantities and parameters involved in magnetic resonance imaging, in order from source to effect, including amplitude and phase, totaling 18.

[0208] Table 2

[0209]

[0210] The beneficial effects of this invention are as follows:

[0211] The physical principles upon which this invention is based are clear, the derivation process is consistent, and there are few assumptions. The methods for acquiring and processing the radio frequency magnetic field distribution in magnetic resonance imaging are entirely based on real measurement data. The quantitative distribution of electrical and magnetic properties of the imaged human tissue, such as conductivity, capacitance, and permeability, can be further derived from the reconstructed radio frequency magnetic field, without the need for surgical procedures or biopsy pathological analysis. Attached Figure Description

[0212] Figure 1 A schematic diagram of the device structure for implementing the magnetic resonance imaging parameter distribution acquisition and calculation method of the present invention;

[0213] Figure 2 This is a flowchart illustrating the data processing involved in the magnetic resonance imaging parameter distribution acquisition and calculation method of the present invention.

[0214] Figure 3 The timing relationship between the radio frequency excitation pulse, gradient encoding pulse, and detection echo signal during the acquisition of the equivalent proton density magnetic resonance signal for performing magnetic field isoparametric measurements using a fast spin echo sequence;

[0215] Figure 4 This diagram illustrates the superposition of precession magnetization vector amplitudes and phase unwinding between channels for signal acquisition using conventional multi-channel receiving coils for imaging scanning; where (a) is the sensitive field amplitude accumulation diagram, (b) is the sensitive field phase angle winding diagram, and (c) is the unwinding phase.

[0216] Figure 5 The diagram illustrates the evolution of the magnetization vector precession before and after the excitation pulsed circularly polarized emission magnetic field acts on the spin system of the target body. In the diagram, (a) is the thermal equilibrium state diagram, (b) is the stimulated absorption diagram, and (c) is the spontaneous emission diagram.

[0217] Figure 6 The original amplitude and argument images of the equivalent spin density obtained from the factory-specific water phantom A for the second sequence II scan;

[0218] Figure 7 To receive the factory-specific permeability amplitude and phase diagram of the water model A related to the magnetic field;

[0219] Figure 8 Deflection angle The amplitude and phase image of the emission magnetic field of the factory-specific water model A;

[0220] Figure 9 The amplitude and phase images of the sensitive field of the receiving coil with constant magnetization coefficient are obtained based on standard homogeneous water model scanning.

[0221] Figure 10 This is a reconstructed image of the amplitude and phase of the sensitive field of the transmitting coil obtained by measurement and calculation under no assumptions.

[0222] Figure 11 The synthesized factory-specific water model A receives the magnetic field amplitude and phase image;

[0223] Figure 12 A phase image of the transmit-receive magnetic field of a factory-specific water phantom A, suitable for use in magnetic resonance electrical property imaging;

[0224] Figure 13 A diagram showing the amplitude and phase distribution of the permeability of the factory-specific water model A related to the launch magnetic field.

[0225] Figure 14 The image shows the normalized magnetization amplitude of the factory-specific water phantom A during magnetic resonance signal acquisition, and the reconstructed proton density distribution of the factory-specific water phantom A based on reference values.

[0226] Figure 15 The receiver permeability, transmit magnetic field amplitude, and transmit-receive phase image were obtained by magnetic resonance imaging of a high-conductivity sodium chloride-containing specific water phantom N.

[0227] Figure 16 The receiving permeability, transmitting magnetic field amplitude, and transmitting-receiving phase image of a high-permeability gadotyl ether-containing specific water phantom G were obtained by magnetic resonance imaging.

[0228] Figure 17Comparison of the calculated permeability map of specific water model A, the inverted conductivity map and calculated proton density map of specific water model N, and the conductivity statistical histograms of specific water models A, N, and G;

[0229] Figure 18 Inverted conductivity and relative permittivity plots for healthy volunteers and volunteers with cerebral hemorrhage.

[0230] Figure label:

[0231] 1. Main magnet of magnetic resonance imaging system; 2. Imaging target; 3. Radio frequency pulse excitation coil; 4. Induction signal detection coil; 5. Pulse sequence controller. Detailed Implementation

[0232] The technical solutions of the present invention will be clearly and completely described below with reference to the accompanying drawings of the embodiments of the present invention. Obviously, the described embodiments are only some embodiments of the present invention, and not all embodiments. All other embodiments obtained by those skilled in the art based on the embodiments of the present invention without creative effort are within the protection scope of the present invention.

[0233] The equivalent spin density grayscale image reconstructed from a target body obtained by magnetic resonance imaging (MRI) scans physically represents the amplitude distribution of the macroscopic magnetization vector in each voxel, while the corresponding phase image represents the relative phase distribution of the vector rotation angle. This invention proposes a method for measuring the amplitude and phase of the RF transmitted magnetic field and the received magnetic field in MRI, as well as a method for calculating derived parameters such as permeability amplitude and phase. Combined with a magnetic homogeneous water model, by changing the deflection angle of the scanning pulse sequence and the receiving bandwidth parameters, the true electromagnetic wave and related characteristic parameter distributions within human tissue can be reconstructed.

[0234] like Figure 1 The diagram shows a schematic of the magnetic resonance imaging (MRI) apparatus used to implement the method of this invention. Under the static magnetic field generated by the main magnet 1 of the MRI system, the spin system within the imaging target 2 is polarized. The transmitting magnetic field generated by the radio frequency pulse excitation coil 3, controlled by the pulse sequence controller 5, resonates with the spin system of the imaging target 2, performing a fast spin echo sequence scan. The relevant macroscopic magnetization intensity vector generates a k-space voltage signal in the induction signal detection coil 4. This voltage signal is processed by inverse Fourier transform to reconstruct a magnetic resonance proton density amplitude-spin precession phase image related to the amplitude and phase of the transmitting and receiving magnetic fields. Using the special sequence set proposed in this invention, human tissue and a reference homogeneous water model are scanned. The obtained amplitude and argument image data are divided and subtracted to obtain the distribution of the steady transmitting magnetic field, alternating receiving magnetic field, and derived parameters such as permeability.

[0235] The implementation process and related data processing of the parameter distribution acquisition and calculation method for electromagnetic measurements in magnetic resonance imaging of this invention are as follows: Figure 2 As shown, the special sequence set used to scan the target body and its implementation steps are as follows:

[0236] Step 1: Perform a fast spin echo sequence I scan of human tissue. The parameters of the sequence I are (α1, TE0, TR0, Bw1), where α is the deflection angle, TE is the echo time, TR is the repetition time, and Bw is the receiving bandwidth.

[0237] Step 2: Perform a fast spin echo sequence II scan of human tissue, wherein the parameters of sequence II are (α2, TE0, TR0, Bw1);

[0238] Step 3: Perform a fast spin echo sequence III scan of human tissue, wherein the parameters of sequence III are (α2, TE0, TR0, Bw2);

[0239] Among them, the TE and TR remain consistent between sequences, ensuring the consistency of magnetic vector precession amplitude and phase between different scans for the same target, thus creating the principle conditions for the measurement and calculation of radio frequency magnetic field. In order to make the dynamic range large when calculating the amplitude of the transmitted magnetic field in the later stage, and considering the rationality of the system gradient pulse intensity, the parameters can be set to include α1=120°, α2=60°, Bw1=62.5kHz, and Bw2=31.25kHz.

[0240] In order to eliminate the sequence-related contrast weights in the signal data and the influence of the non-uniformity of the coil magnetic field when calculating the magnetic field distribution in human tissue, the homogeneous water model was scanned again with the same sequence set.

[0241] Step 4: Perform the fast spin echo sequence I scan on the homogeneous water model;

[0242] Step 5: Perform the fast spin echo sequence II scan on the homogeneous water model;

[0243] Step 6: Perform the fast spin echo sequence III scan on the homogeneous water model;

[0244] Step 7: Calculate the distribution data of the radio frequency magnetic field and related derived parameters using the scanning data of the imaging body and the homogeneous water model, including (1) the magnetic field strength of the radio frequency transmitting coil, (2) the proton density of the target body, (3) the complex capacitance of the target body, (4) the magnetic permeability of the excited target body, (5) the magnetic permeability of the detected target body, (6) the transmitting electric field strength, (7) the specific absorption rate of the target body, (8) the magnetic field strength of the radio frequency receiving coil, (9) the magnetization of the detected target body and (10) the magnetic flux density transmission-reception continuous phase.

[0245] When scanning a homogeneous water phantom, it is necessary to ensure a one-to-one correspondence between the acquired voxel data positions of the water phantom and the target calibrated human tissue data positions for easy post-processing. Preferably, the deflection angle combination (α1, α2) can also be (90°, 45°), and the receiving bandwidth combination (Bw1, Bw2) can also be (125kHz, 62.5kHz) or other numerical ratios. Figure 2 As shown, the amplitude data of the signal acquired in step 1 Divide by the amplitude data collected in step 2 This yields an approximate received magnetic field amplitude distribution of human tissue after removing proton density weights. The term "approximate" is used because the amplitude data, in addition to retaining the relationship with echo time and repetition time contrast, still includes the weighted contribution of the transmitted magnetic field. The signal amplitude data acquired in step 4... Divide by the amplitude data collected in step 5 The approximate received magnetic field amplitude distribution of the homogeneous water model after removing the proton density weight is obtained.

[0246] Further dividing the approximate received magnetic field amplitude data of the human tissue by the approximate received magnetic field amplitude data of the homogeneous water model yields the relative permeability amplitude distribution of the human tissue after removing the amplitude weight of the sensitive field of the receiving coil. The symbol '~' indicates the distributional alternation parameter related to the detection, and the calculation process assumes that the human tissue is magnetically homogeneous.

[0247] Although the permeability amplitude is obtained, the signal amplitude data acquired in step 1 can actually be used to... Divide directly by the signal amplitude data acquired in step 2 Under the condition of a two-fold relationship between deflection angles, the deflection angle is obtained by simultaneously removing the proton density and the coil sensitive field weight. Given a pulse width and a distribution of twice the cosine value, the amplitude distribution of the emitted magnetic field of human tissue with respect to sequence II is calculated after processing. .

[0248] Regarding phase calculation, the signal argument data acquired in step 3... Subtract the argument data collected in step 2 The phase of the transmitted magnetic field is removed to obtain the phase distribution of the received magnetic field at the voxel level of the imaged human tissue, which is the precession phase distribution of the source magnetization vector that reflects the spatial resolution. This phase distribution contains the superposition contribution of a certain non-uniform distribution of the sensitive field of the receiving coil.

[0249] Signal argument data acquired in step 6 Subtract the argument data collected in step 5 Since the receiving bandwidth satisfies the double value relationship, the phase distribution of the receiving coil sensitive field with respect to sequence II, after removing the phase of the transmitting magnetic field, is obtained.

[0250] By further subtracting the phase data of the receiving coil's sensitive field from the obtained phase data of the received magnetic field in the human tissue, the phase distribution of the permeability of the human tissue itself can be obtained. .

[0251] Based on the definition of the equivalent spin density signal phase in magnetic resonance, the phase relationship between magnetization and the emitted magnetic field, and using the phase data from Sequence II measurements of human tissue and the obtained phase distribution of the received magnetic field, the Sequence III phase is directly subtracted by twice the Sequence II phase, and then subtracted again. Determine the quasi-uniform phase distribution of the excitation radio frequency pulse emission magnetic field within the tissue. .

[0252] In conventional fast spin echo sequence design, to improve image quality, manufacturers have fully considered the continuity and consistency of the acquisition start and end of each round of magnetic field vector precession process within a single sequence period and between sequence periods. This implicitly guarantees the calculation conditions that the amplitude ratio and phase difference involved in the vector measurement of this invention must meet. It also means that, under the condition that the repetition time TR is much greater than the target volume T1, for example, TR >> 3T1, the sequence-related parameter deflection angle... The settings of echo time (TE), repetition time (TR), and receiving bandwidth (Bw) do not affect the correctness of the electromagnetic parameter measurement and calculation method of this invention, but they do affect the specific magnitudes of the amplitude and phase of the transmitted and received magnetic fields, impacting the contrast of the image and the graph, such as the received magnetic field under specific TE and TR. In the sequence group, if the repetition time parameter is sufficiently long, for example, TR greater than or equal to three times the longest longitudinal relaxation time T1 of human tissue components, the excitation deflection angle and the factor terms of the sequence time parameters in the signal can be separated to facilitate data processing, achieving the purpose of reducing the magnitude of the exponential terms related to TR, and improving the accuracy of the measured magnetic field and other derived parameter calculations.

[0253] This invention can employ a proton density-weighted imaging method to scan human tissue and a reference homogeneous water model. Figure 3The diagram illustrates the sequence timing and k-space data filling method used. In the diagram, the excitation deflection angle and refocusing flip angle are set to 90° and 180°, respectively. The nominal echo time is selected to maximize the amplitude of the first echo, i.e., minimize the gradient of the corresponding phase-encoded gradient pulse. As the intensity of subsequent gradients increases, the echo amplitude gradually decreases. According to the Cartesian sampling scanning method, the echo sampling data corresponding to the minimum gradient (i.e., 0) is filled into the central strip of the k-space, while the data with the minimum echo amplitude at the maximum gradient is filled into the edge strips. Then, an inverse Fourier transform is performed on the complete original wavenumber domain k-space data to obtain the magnetic vector precession amplitude and relative phase distribution in the image space. Here, the moderately long echo intervals in the pulse sequence, combined with necessary anti-polarity compensation gradient pulse segments, ensure that the overall magnetic vector astigmatism related to each k-line is consistent throughout the acquisition process. This is controlled by the receiving bandwidth Bw; different Bw values ​​create different proportions of astigmatism, facilitating the measurement and calculation of the magnetic vector phase. In the figure, the echo train length is 8, indicating that the k-space is divided into 8 strips; and the number of k-lines filled in each strip is n, indicating that the pulse sequence is repeated n times, that is, keeping the TE constant, changing the intensity of the gradient pulse n times within a certain range, and collecting the required total The k-line data eventually fills the k-space corresponding to all voxels of the entire slice, i.e., an image. Furthermore, the decrease in echo peak amplitude is a result of the combined effects of endogenous spin-spin interactions and exogenous phase-coding gradient enhancement.

[0254] The inverse Fourier transform of k-space data in the time domain can only yield image domain argument data in the principal value interval [-π, π], which means there is potential entanglement. Therefore, relevant untangling preprocessing is also essential. Figure 4 This is a schematic diagram of data combining from an 8-channel receiving coil. When the received signal is used for image acquisition, based on the distribution characteristics of the sensitive field detected by the sub-coil, the amplitude images of the single channels overlap before vector combining, as shown below. Figure 4 As shown in (a), the argument image exhibits angular entanglement, as... Figure 4 As shown in (b) and (c).

[0255] In order to measure and calculate the magnetic field and related derived parameters, in a specific embodiment of the preprocessing of row vector merging and phase unwrapping of a multi-channel acquired image data, the process for obtaining the precession magnetic vector amplitude and argument data at the voxel level on the fault is as follows:

[0256] Step 1: Read the medical image file to obtain the reconstructed amplitude and argument data of a single channel for later use;

[0257] Step 2: Unwrap the argument image data of a single channel;

[0258] Step 3: Based on the single-channel amplitude and unwinding argument data, obtain the total amplitude and phase data on the corresponding fault.

[0259] ;

[0260] In the formula, For the complex numerical representation of the combined circularly polarized radio frequency wave. The amplitude of a single channel. For the argument of a single channel, The channel ordinal number is used. The vector merging process involves first converting the amplitude and argument of the voxels into complex values, then superimposing the data between channels according to the real and imaginary parts, and finally restoring them to the form of amplitude and argument.

[0261] Step 4: Unwrap the argument data of the fault to obtain the phase data of the true relative argument.

[0262] Based on the obtained magnetic vector merging amplitude and unwinding phase data, the transmitting and receiving magnetic fields and derived parameters can be further calculated.

[0263] The measurement and calculation of the emitted magnetic field phase of the magnetic resonance target in this invention is based on the concept of continuous transmit-receive phase related to the magnetic vector precession during the imaging process. This includes the evolution of magnetic vector precession before and after the excitation radio frequency pulse acts on the spin system, such as the fast spin echo sequence. Figure 5 As shown in (a), for the voxel array, at t=0, a principal static magnetic field B0 exists, resulting in a lumped macroscopic magnetization intensity vector M0 distribution within the magnetized target. In the rotating coordinate system, at t=0, a right-handed circular deflection magnetic field pulse is emitted along the positive x' axis. M0 shifts towards the positive y' axis after transient response (consistent with the left-hand rule), such as Figure 5 As described in (b). After the pulse, starting from t=0+, due to the spin-spin effect, the transverse component Mxy gradually dephases and becomes incoherent, while the amplitude of the longitudinal component Mz, affected by the spin-lattice effect, gradually relaxes and recovers from 0, as shown in section (b). Figure 5 As described in (c). The precession process mathematically obeys the Bloch equation. Under pulsed nuclear magnetic resonance conditions, the pulse width is ignored, and it is assumed that the initial phase of the excited voxels lags behind. Phase pi / 2. After acquiring image data and calculating the received magnetic field phase using a scanning method that maintains the amplitude ratio and controls the phase difference, the transmitted magnetic field phase is further calculated. At that time, considering the phase of the target image To ensure the phase of the detection signal following the excitation pulse, and considering that the receiving and transmitting magnetic fields propagate in opposite directions, therefore... The DC magnetic fields, including the main magnetic field and magnetic field gradient, as well as the necessary phase callback segments in the design of frequency and phase-encoded gradient pulses, are independent of the measurement of the RF magnetic field phase. That is, in the inverse Fourier transform form, the wavenumber domain equivalent Larmor frequency precession phase of the voxel magnetic vector is independent of the initial phase related to the RF magnetic field. Furthermore, the several 180° refocusing pulses (rather than other values) following the initial excitation pulse during the echo chain of a single sequence simplify the analysis of the precession process, thus facilitating the calculation of the receive and transmit magnetic field phases from the spin system's transmit-receive phase using this method.

[0264] Assuming the main magnetic field is absolutely uniform and the imaging target is magnetically homogeneous and constant equal to the vacuum permeability, this invention evaluates the error impact of assuming the tissue magnetic properties are equal to the water magnetic susceptibility when calculating the amplitude and phase of the received magnetic field. Specifically, it addresses the amplitude data processing of two sets of sequences scanning tissue and a water phantom with different excitation deflection angles. , The relative magnetic susceptibility of several typical human tissues involved in magnetic resonance imaging differs significantly. Table 3 lists some reference magnetic susceptibility and the calculated ratios.

[0265] Table 3

[0266]

[0267] In the table above, it is assumed that the sensitive field H1 of the transmitting coil is uniformly distributed, with a uniform pulse width of 3000 μs. It can be calculated that although the differences in magnetic susceptibility of calcifications, fat, and methemoglobin relative to water are 44.4%, 7.8%, and 101.3%, respectively, the differences in the ratios of the characteristic contrast weights in functional form between tissues are only 7.3 ppm, -1.3 ppm, and -17 ppm, respectively. Furthermore, the emitted magnetic flux density is on the order of μT, and the magnetic flux density generated by exciting water at a deflection angle of 60° is approximately 1.31 μT.

[0268] Under specific structure and size conditions of the transmitting coil, the rated excitation power for local organ magnetic resonance imaging is 6kW, and the corresponding effective current is approximately 10.954A, such as when implementing a 180° refocusing flip angle.

[0269] In the following embodiments, the magnetic resonance system field strength is 3.0T, all based on a fast spin echo sequence.

[0270] In one embodiment, the scanning objects are a factory-specific water phantom A containing internal structures, a factory-specific water phantom N containing internal structures with added sodium chloride, and a factory-specific water phantom G containing internal structures with added sodium chloride and the magnetic contrast agent gadotate dimeglumine (Gd-DOTA). The reference longitudinal and transverse relaxation times of the factory-specific water phantom A are 1115.7 ms and 287.5 ms, respectively. The geometric sequence parameters of the imaging settings are shown in Table 4.

[0271] Table 4

[0272]

[0273] In one embodiment, the scanning subject is a control group of healthy subjects, and the parameters of the short echo time series used for head imaging are shown in Table 5.

[0274] Table 5

[0275]

[0276] In one embodiment, the subjects being scanned were clinical patients with cerebral hemorrhage, and the parameters of the long repetitive time series used for head imaging are shown in Table 6.

[0277] Table 6

[0278]

[0279] Using a conductivity meter and consulting relevant data, the parameters such as the magnetic susceptibility and inverted target electrical properties of the specific water module and the test subject, as well as their reference values, are summarized in Table 7.

[0280] Table 7

[0281]

[0282] In the table, the Dotarem dose standard for the NaCl+Gd water model is 0.1 mmol / kg; the SAR value range is provided by the magnetic resonance scanning system during image data acquisition.

[0283] Figure 6 The original amplitude image shows the original amplitude distribution obtained from the factory-specific magnetic resonance scan of the water model A. The single-channel original argument image is the argument data distribution collected and scanned by the corresponding receiving coil channel 4. The inner diameter of the water model is 140mm and the height is 65mm.

[0284] Figure 7 The received permeability amplitude diagram shows the relative permeability amplitude diagram calculated from the water model A. The received permeability phase diagram is the corresponding phase diagram. Due to the data processing involving data related to homogeneous water models, the area where non-specific water model A and homogeneous water models intersect in the diagram is marked as "Not a Number" (NaN). The central block and edge columns can be distinguished on the phase diagram, but the quality of the calculated reconstruction is poor.

[0285] Figure 8 The emitted magnetic field amplitude image shows about The main measurement sequence includes the amplitude image of the emitted magnetic field of water model A and the phase image of the emitted magnetic field, which are the corresponding phase images.

[0286] Figure 9The amplitude image of the received sensitive field shows the amplitude image of the magnetic field intensity sensitive field distribution of the receiving coil itself, based on the measurement of the homogeneous water model. The phase image of the received sensitive field is the corresponding phase image. In the figure, the noise at the edge of the target area is due to the pixel being close to the background area (NaN) during pixel calculation processing.

[0287] Figure 10 The amplitude image of the transmitting sensitive field shows the amplitude image of the transmitting coil's sensitive field, while the phase image of the transmitting sensitive field is the corresponding phase image. In the figure, the magnetic field strength amplitude is in H / m, and the phase is in rad.

[0288] Figure 11 The received magnetic field amplitude image shows the received magnetic field amplitude image of the scanned specific water model A, and the received magnetic field phase image is the corresponding phase image. Due to limitations of the receiving parameter measurement and calculation methods themselves, the amplitude includes the magnetization constant weight related to the homogeneous water model accessories. Note that the amplitude image shows the intersection region image constrained by the homogeneous water model, while the phase image does not have this limitation.

[0289] Figure 12 The transmit-receive phase image shows the transmit-receive magnetic field phase image of a specific water model A. This quantity can be used to invert the numerical distribution of the electrical properties of the target.

[0290] Figure 13 The emission permeability amplitude diagram shows the permeability amplitude diagram of the specific water model A calculated using the emission magnetic field and the sensitive field, while the emission permeability phase diagram is the corresponding phase diagram.

[0291] Figure 14 The magnetization image shows the normalized magnetization image of the specific water phantom A, and the proton density map is the corresponding proton density distribution map. In the figure, the proton density data has been center-shifted to preserve the distribution shape.

[0292] Figure 15 The permeability map shows the receiving relative permeability distribution calculated from a specific water model N containing 6 g / L NaCl, i.e., a numerical map of the magnetic properties of the scanned solution; the emission magnetic field image is the emission magnetic flux density image, and the emission-receive phase image is the emission-receive phase of the continuous magnetic flux density, which is used to further invert and calculate the electrical property distribution of the solution.

[0293] Figure 16 The permeability map shows the relative permeability distribution of a specific water phantom G containing 6 g / L NaCl and 0.1 mmol / kg Gd-DOTA. The emission magnetic field image is the emission magnetic flux density image, and the emission-receive phase image is the emission-receive phase of the continuous magnetic flux density. The emission magnetic field image and the emission-receive phase image implicitly contain the structural feature contrast information of the edge column, which can be recovered and revealed after electrical property inversion.

[0294] Figure 17 The calculated permeability plot of the specific water model A is shown, with values ​​ranging from [value missing]. Inversion conductivity plot of specific water model N, with a numerical range of (0.5, 2.5) S / m; calculated proton density plot, with a numerical range of... Comparison of conductivity statistical histograms for specific water models A, N, and G.

[0295] Figure 18 The inversion conductivity plots for healthy volunteers are shown, with values ​​ranging from (-0.5, 2) S / m; the inversion relative permittivity plots for healthy volunteers are shown, with values ​​ranging from (-50, 200); the inversion conductivity plots for volunteers with cerebral hemorrhage are shown, with values ​​ranging from (-1, 2) S / m; and the inversion relative permittivity plots for volunteers with cerebral hemorrhage are shown, with values ​​ranging from (-200, 300).

[0296] The specific embodiments described above further illustrate the purpose, technical solution, and beneficial effects of the present invention. It should be understood that the above descriptions are merely specific embodiments of the present invention and are not intended to limit the present invention. Any modifications, equivalent substitutions, improvements, etc., made within the spirit and principles of the present invention should be included within the protection scope of the present invention.

Claims

1. A method for scanning and calculating electromagnetic parameter distribution in magnetic resonance imaging, characterized in that, The method includes the following steps: Step 1: Scan human tissue using fast spin echo sequence I. The parameters of sequence I are (α1, TE0, TR0, Bw1), where α1, TE0, TR0, and Bw1 are the deflection angle, echo time, repetition time, and receiving bandwidth for implementing sequence I, respectively. Step 2: Perform a fast spin echo second sequence II scan of human tissue. The parameters of the second sequence II are (α2, TE0, TR0, Bw1), where α2, TE0, TR0, and Bw1 are the deflection angle, echo time, repetition time, and receiving bandwidth for performing the second sequence II, respectively. Step 3: Perform a fast spin echo third sequence III scan on human tissue. The parameters of the third sequence III are (α2, TE0, TR0, Bw2), where α2, TE0, TR0, and Bw2 are the deflection angle, echo time, repetition time, and receiving bandwidth for performing the third sequence III, respectively. Step 4: Scan the homogeneous water model using the first sequence I of the fast spin echo; Step 5: Scan the homogeneous water model using the aforementioned fast spin echo second sequence II; Step 6: Scan the homogeneous water model using the aforementioned fast spin echo third sequence III; Step 7: Based on the original amplitude and argument image data of the equivalent spin density of human tissue and homogeneous water phantom obtained by scanning, calculate the distribution data of the magnetic field strength, proton density and magnetization, permeability, conductivity and permittivity, transmission magnetic flux density, transmission electric field strength, receiving coil sensitive field, receiving magnetic flux density, signal transmission-reception phase and specific absorption rate in the target region of the second sequence II.

2. The method for scanning and calculating electromagnetic parameter distribution in magnetic resonance imaging according to claim 1, characterized in that, The deflection angles of the first sequence I to the third sequence III satisfy The receiving bandwidth meets The refocusing flip angle is a fixed 180°; the repetition time is taken as... , where T1i is the longitudinal relaxation time constant of the i-th component in the target body.

3. The method for scanning and calculating electromagnetic parameter distribution in magnetic resonance imaging according to claim 1, characterized in that, The calculation of the amplitude and phase distribution of the emitted and received magnetic fields within the human tissue involves the following image data processing: Step 3.1: Collect the amplitude data from Step 1. Divide by the amplitude data scanned in step 2 Then scan the amplitude data from step 4. Divide by the amplitude data scanned in step 5 Then execute By removing the common sensitive field weights related to echo time and repetition time, the magnetic permeability data of human tissue relative to a homogeneous water model are obtained. , '~' indicates based on the received magnetic field; where, when considering the sinusoidal weight of the deflection angle, the magnetic homogeneity of human tissue is assumed; Step 3.2: Obtain Data Then, calculate , It is the gyromagnetic ratio. The pulse width is used to obtain the measurement data of the emitted magnetic field of human tissue. ; Step 3.3: Collect the argument data from Step 3. Subtract the scan argument data from step 2 The phase of the received magnetic field of human tissue related to the received bandwidth Bw1 was obtained. ; Step 3.4: Receive the magnetic field phase of the obtained human tissue. minus Subtract again The phase of the emitted magnetic field of human tissue is obtained. .

4. A method for scanning and calculating electromagnetic parameter distribution in magnetic resonance imaging according to claims 1, 2, and 5, characterized in that, The calculation of the electromagnetic sensitive field of the coil and the distribution of transmitted waves and characteristic parameters within human tissue involves the following image data processing: Step 4.1: Replace the amplitude data scanned in Step 4. And the amplitude data scanned in step 5 Input operation , The resonant angular frequency is... Given the permeability of a homogeneous water model, the magnetization in thermal equilibrium is obtained. Weighted receiving coil sensitive field amplitude data , The data represents the proton density of the homogeneous water model; ER represents the sequence echo time and repetition time; the argument data from step 6. Subtract the argument data from step 5. The phase of the receiving coil sensitive field related to bandwidth Bw1 is obtained. distributed; Step 4.2: Replace the amplitude data scanned in step 4. And the amplitude data scanned in step 5 Input operation The amplitude distribution data of the magnetic field strength generated by the transmitting coil were obtained. The phase of the transmitting coil's magnetic field strength is related to the pulse width; the argument data scanned in step 5 is expanded as follows: In the formula The phase of the magnetic permeability of human tissue. The transmit phase is obtained using the receive coil sensitive field phase from step 4.

1. ; Step 4.3: Obtain the magnetization-weighted amplitude of the magnetic field received by human tissue from steps 3.1 and 4.

1. ; The phase of the human tissue permeability in the frequency domain is obtained by subtracting the phase of the receiving coil's sensitive field from the phase of the received magnetic field in step 3.

3. ; Step 4.4: Obtain the magnetic permeability amplitude of the emission-related human tissue from steps 3.2 and 4.

2. ; Step 3.4 Subtracts the phase of the magnetic field strength of the transmitting coil from the phase of the emitted magnetic field in step 4.2 to obtain the phase of the magnetic permeability of the emitted human tissue. ; Step 4.5: Divide the amplitude data from Step 2 by the relative permeability of human tissue, the deflection angle of Sequence II, and the sensitive field of the weighted receiving coil in sequence to obtain the magnetization data relative to the homogeneous water model. ; The magnetization data relative to the homogeneous water model Obtain nonnormalized proton density distribution data of human tissues ; Step 4.6, add the argument data from Step 2 Taking the negative value yields the magnetic flux density emission-receiver phase of human tissue. .

5. A method for scanning and calculating electromagnetic parameter distribution in magnetic resonance imaging according to claim 1, 2, 3, or 4, characterized in that, The conductivity and capacitance of the human tissue are determined by the emitted magnetic field. and the magnetic field strength of the transmitting coil The inversion yields the following assumptions: (1) the human tissue is electrically homogeneous; (2) the longitudinal receiving magnetic field is zero; (3) the first and second longitudinal partial derivatives of the longitudinal magnetic field strength of the transmitting coil are zero; and (4) the order of the second-order transverse partial derivatives of the transverse component of the coil magnetic field strength can be arbitrarily interchanged. In the rotating coordinate system, the real and imaginary parts of each complex magnetic field quantity correspond to the x' and y' axes, respectively. The conductivity... and permittivity The calculation formula is: ； In the formula, Im[] and Re[] represent taking the imaginary and real parts of the complex number, respectively, and i is the imaginary unit. Take conjugate; The transmitting electric field strength is obtained by substituting the coil magnetic field strength, conductivity, and permittivity into Ampere's law. The formula for calculating the longitudinal component of the electric field strength is: ； In the formula, '_' represents a complex number. L1 represents the complex permittivity, and L1 represents the phase of the magnetic field strength. Ignoring the transverse component of the electric field intensity in magnetic resonance, the specific absorptivity associated with the emitted electric field is obtained by substituting the proton density from step 4.5 and the longitudinal component of the emitted electric field intensity into the following equation: ； In the formula, The proton density is equivalent to the mass density.

6. A method for scanning and calculating electromagnetic parameter distribution in magnetic resonance imaging according to claim 1, 2, or 5, characterized in that, α2 = 60° maximizes the dynamic range of the magnetic field emitted by human tissue.

7. A method for scanning and calculating electromagnetic parameter distribution in magnetic resonance imaging according to claim 1, 2, 5, or 6, characterized in that, The relevant parameter distribution of the emission magnetic field is used to perform three-dimensional reconstruction based on data from two-dimensional sheet measurements and calculations.

Images