WO2023246175A1

WO2023246175A1 - Skipped echo coding-based magnetic resonance magnetic field measurement method and apparatus

Info

Publication number: WO2023246175A1
Application number: PCT/CN2023/079934
Authority: WO
Inventors: 关浩男; 王敏敏; 冯涛; 张韶岷; 张祎
Original assignee: 浙江大学
Priority date: 2022-06-22
Filing date: 2023-03-06
Publication date: 2023-12-28
Also published as: CN115267627A; CN115267627B

Abstract

A skipped echo coding-based magnetic resonance magnetic field measurement method and apparatus. First, different positive and negative modulated currents synchronized with refocusing pulses are applied to imaging tissue in two scans until sampling of a first echo signal for each repetition time is completed, enabling phase changes induced by the current-generated magnetic field to accumulate continuously before the end of the echo sampling. In skipped echo coding, sampling of the first n echo signals is skipped, ensuring that the sampled first echo signal experiences a sufficiently long time for external current, thereby improving the signal-to-noise ratio of the result. After the current ends, the signal is collected using a turbo spin echo sequence having a turbo factor of m. Finally, the phase results of the two scans having different initial currents are subtracted, and a linear relationship between the phase difference and the magnetic field generated by the current (B_z) is used for calculation, to obtain a result B_z. In the present method, a tenfold increase in the measurement speed of the B_z magnetic field is achieved without losing signal-to-noise ratio.

Description

A magnetic resonance magnetic field measurement method and device based on jump echo coding

Technical field

This application relates to the field of magnetic resonance technology, and in particular to the field of magnetic resonance MRCDI and MREIT imaging magnetic field measurement.

Background technique

Magnetic Resonance Current Density Imaging (MRCDI) and Magnetic Resonance Electrical Impedance Tomography (MREIT) are two emerging imaging methods for measuring internal electrical property parameters of tissues. Among them, MREIT technology By combining the injection of external current, the conductivity distribution with high spatial resolution inside the tissue can be obtained non-invasively; MRCDI technology can obtain the current density distribution inside the tissue stimulated by external current, which can be used for nerve stimulation such as transcranial electrical stimulation (tDCS). Provide guidance on the specific implementation of control technology.

This technology combines the application of current through two electrodes on the outside of the imaging object, and uses the principle that magnetic resonance phase information is sensitive to the size of the magnetic field in the direction of the main magnetic field (B ₀ ). It measures the spatial distribution of the magnetic field generated by the current inside the imaging object and the main magnetic field ( B ₀ ) parallel direction component (B _z ), and then use B _z to solve and calculate the information of current density or conductivity distribution inside the imaging body.

At present, the measurement of the B _z magnetic field mainly uses the spin echo (SE) sequence. By applying a direct current that changes in direction synchronized with the refocusing pulse, the continuous accumulation of the signal phase is achieved, using two different currents (such as 1mA and 2mA, 2mA and -2mA, etc.), the phase data of the scanning results are compared, and the distribution of B _z is calculated through the linear relationship between the phase difference and B _z . However, due to the inefficiency of the SE sequence in filling k-space when collecting signals, the current measurement of B _z is still not efficient enough. It is difficult to obtain the spatial distribution information of the whole brain in a short time and is difficult to apply in actual needs. The fast magnetic field measurement sequence based on the jump echo coding of the present invention can not only effectively measure the distribution of B _z , but also greatly improves the measurement efficiency and reduces the required time, thus providing the possibility for practical applications of MRCDI and MREIT for clinical needs. .

Contents of the invention

In view of the shortcomings of low efficiency of B _z measurement in the existing technology, the present invention intends to improve the existing technology and propose a more efficient method and device for B _z measurement. The present invention intends to use fast spin echo (turbo spin echo, TSE) method to collect data, and turbo factor is the acceleration factor (m). Compared with the original spin echo (SE) sequence, the acquisition speed of this sequence can be increased nearly ten times. However, since the spin echo sequence echo signal exhibits an exponential attenuation of T ₂ , when the echo spacing is large, a large acceleration factor will make the signal of the echo behind each TR too small, thus reducing the resolution of the final imaging. rate, so the selection of smaller echo spacing is very necessary. At the same time, in the context of this problem, the duration of the external current in the spin echo sequence method is consistent with the echo spacing. Therefore, when the echo spacing is small, the shorter current duration makes the phase change caused by B _z smaller and reduces the The signal-to-noise ratio of the B _z measurement is obtained.

Based on this, the present invention provides a magnetic resonance magnetic field measurement method and device based on jump echo coding, which can realize rapid measurement of the magnetic field generated by external current, while ensuring the measurement signal-to-noise ratio and measurement effect, and improving the magnetic resonance MRCDI and MREIT imaging performance.

The present invention is specifically implemented by adopting the following technical solutions:

In a first aspect, the present invention provides a magnetic resonance magnetic field measurement method based on jump echo coding, specifically as follows:

The same imaging sequence is used to scan the target tissue twice, and during the first scanning process, a first stimulation current with positive and negative conversion synchronized with the refocusing pulse is applied to the imaging tissue and the first scanning result is obtained, and during the second scanning process, the target tissue is scanned twice. During the scanning process, a second stimulation current with positive and negative conversion synchronized with the refocusing pulse is applied to the imaging tissue and the second scanning result is obtained; the phase difference between the first scanning result and the second scanning result is calculated, and the phase difference is combined with the B _z magnetic field The linear relationship between the B _z magnetic field;

The imaging sequence divides all 180° refocusing pulses in each repetition period in the fast spin echo sequence into a jump echo encoding module and a fast spin echo sampling module, in which the first n 180° refocusing pulses belong to the jump echo encoding module. Wave encoding, the remaining 180° refocusing pulse belongs to the fast spin echo sampling part, n is a positive integer not less than 1; in the jump echo encoding module, the 180° refocusing pulse in the module and its corresponding slice selection gradient The encoding is consistent with the fast spin echo sequence, but the phase gradient encoding and frequency gradient encoding after each 180° refocusing pulse in the module are canceled; the fast spin echo sampling module samples the signal according to the fast spin echo method and Filled into k space, the first frequency gradient code in each repetition period is the first frequency gradient code in the fast spin echo sampling module;

In each repetition period, the first stimulation current lasts from the first 90° radio frequency pulse to the end of the first frequency gradient encoding, and is only set to zero during the application of each 90° and 180° radio frequency pulse. In a current duration segment between two adjacent radio frequency pulses, the current magnitude and positive and negative directions are always the same. The currents in adjacent current duration segments have the same magnitude but opposite positive and negative directions; the first stimulation current and the second stimulation current have the same number of current duration segments in each repetition period, and the respective current duration segments are the same. In one correspondence, the start and end times of the corresponding current duration segments of each group are exactly the same but the currents are different.

As a preferred embodiment of the above first aspect, the imaging sequence is specifically implemented in the following manner:

In each repetition period of the imaging sequence, a 90° radio frequency pulse is first applied and a layer selection encoding gradient is applied at the same time, and then a pre-refocusing frequency gradient is applied to restore the spins in phase at the center of the echo, and then jump echo encoding is performed. module and a fast spin echo sampling module; in the jump echo encoding module, n 180° refocusing pulses are applied at intervals, and layer selection gradient encoding is performed while applying each 180° refocusing pulse; the fast auto-rotation In the gyro wave sampling module, m 180° refocusing pulses are applied at intervals, and layer selection gradient encoding is performed while applying each 180° refocusing pulse. After layer selection gradient encoding, phase gradient encoding is performed, and after phase gradient encoding, Frequency gradient encoding is performed, and k-space data is obtained through echo signal sampling while frequency gradient encoding is performed; and the n+m levels applied in the jump echo encoding module and the fast spin echo sampling module select both sides of the gradient encoding A spoiler gradient is set, in which the spoiler gradient area on both sides of the first layer selection gradient encoding is set to a value that can cause 4π phase dispersion, and any two adjacent layers in the n+m layer selection gradient encoding select gradient encoding. Apply spoiler gradients with different areas. The spoiler gradient area on both sides of any one of the n+m level selection gradient codes is not less than the spoiler gradient area on both sides of the first level selection gradient code; finally, for The sampled k-space number is calculated through Fourier transform to reconstruct the amplitude image and phase image of each scan; n and m are positive integers not less than 1 respectively;

As a preferred aspect of the above first aspect, the relationship between spoiler gradient areas on both sides of any two adjacent level selection gradient codes in the n+m layer selection gradient codes is preferably 2 or 1/2 times.

As a preferred option of the first aspect above, the calculation formula for obtaining the B _z magnetic field using the linear relationship between the phase difference and the B _z magnetic field is:
B _z (x,y)=ΔΦ(x,y)/(2γ·Tc·(I ₁ -I ₂ ))

Where: B _z (x, y) represents the B _z magnetic field generated by unit external current at point (x, y), unit: nT/mA, ΔΦ (x, y) represents the first scan result and the second The phase difference of the scanning result at point (x, y), γ is the gyromagnetic ratio of hydrogen atoms, Tc is the duration of the stimulation current in a repeated cycle, I ₁ and I ₂ are the first stimulation current and the second stimulation current that distinguish between positive and negative respectively. 2. Stimulation current.

As a preferred option of the above first aspect, the calculation formula of the duration Tc of the stimulation current in one repetition cycle is:
T _c =(ESP-τ _π )·(n+1)-0.5τ _π/2

Among them: ESP is the time interval between adjacent 180° refocusing pulses, τ _π and τ _π/2 are the action times of a single 180° refocusing pulse and a single 90° radio frequency pulse respectively.

As a preference in the above first aspect, the n is preferably 1 to 3.

As a preference in the above first aspect, the n is further preferably 1.

As a preference in the above first aspect, the m is preferably no more than 10.

As a preference in the above first aspect, the m is further preferably 5.

As a preferred aspect of the above first aspect, in each repetition period, the starting and ending times and sizes of the corresponding current duration segments of each group of the first stimulation current and the second stimulation current are exactly the same, but the positive and negative directions of the current on the contrary.

As a preferred aspect of the above first aspect, among the first stimulation current and the second stimulation current, the size of the first current duration is preferably the maximum value within the safe range of the external current allowed to be applied to the target tissue.

In a second aspect, the present invention provides a magnetic resonance magnetic field measurement device based on jump echo coding for implementing the magnetic resonance magnetic field measurement method described in any of the above-mentioned solutions of the first aspect, which includes magnetic resonance equipment and external electrical stimulation equipment. and computing modules;

The external electrical stimulation device is used to apply the first stimulation current and the second stimulation current to the target tissue during scanning of the magnetic resonance device;

The magnetic resonance equipment is used to perform the imaging sequence and obtain first scan results and second scan results;

The calculation module is used to calculate the phase difference between the first scan result and the second scan result, and use the linear relationship between the phase difference and the _Bz magnetic field to obtain the _Bz magnetic field.

As a preferred aspect of the above second aspect, during the process of executing the imaging sequence to scan the target tissue, the magnetic resonance equipment sends a first synchronization signal to the external electrical stimulation equipment before applying a 90° radio frequency pulse. A second synchronization signal is sent to the external electrical stimulation device before each 180° refocusing pulse of the jump echo encoding module, and a second synchronization signal is sent to the external electrical stimulation device before the first 180° refocusing pulse of the echo signal sampling module is applied. The device sends a third synchronization signal;

After receiving the first synchronization signal, the external electrical stimulation device begins to apply the first current duration to the target tissue according to the set time delay;

Each time the external electrical stimulation device receives the second synchronization signal, it applies the next current duration segment in the opposite positive and negative direction to the previous current duration segment to the target tissue according to the set time delay;

After receiving the third synchronization signal, the external electrical stimulation device applies the last current duration period in the opposite positive and negative direction to the previous current duration period to the target tissue according to the set time delay.

As a preferred aspect of the above second aspect, during the two scanning processes of the magnetic resonance equipment, the external electrical stimulation equipment applies the first current duration section after receiving the first synchronization signal. The currents are the same but positive. The negative direction is the opposite.

As a preference of the above second aspect, it is further preferred that the size of the first current duration segment applied by the external electrical stimulation device after receiving the first synchronization signal during the two scanning processes of the magnetic resonance device is the target. The maximum value within the safe range of external current that is allowed to be applied by an organization.

Compared with the prior art, the present invention has the following beneficial effects:

By reasonably skipping the sampling of this part of the echo in the jump echo encoding, the invention makes the measured echo signal experience a long enough external current action time, improves the phase accumulation induced by the external current, and improves the overall Measuring the quality of the magnetic field solves the problem of low signal-to-noise ratio caused by short echo intervals in traditional measurement methods based on spin echo sequences. At the same time, the fast spin echo sampling method greatly reduces the imaging time, improves the overall magnetic resonance MRCDI and MREIT imaging performance, and the imaging time meets the limitations of clinical application, which also makes the present invention have very important clinical application value.

Description of the drawings

Figure 1 is a block diagram of the magnetic resonance magnetic field measurement sequence based on jump echo coding.

Figure 2 is a comparison and quantitative analysis of the B _z image of the water model scanning experiment based on the magnetic resonance magnetic field measurement sequence based on jump echo encoding and the conventional multi-echo spin echo (MESE) sequence weighted method.

Figure 3 is the large-scale coverage B _z image effect of the magnetic resonance magnetic field measurement sequence scanning water model experiment based on jump echo encoding.

Figure 4 is the large-scale coverage B _z image effect of the magnetic resonance magnetic field measurement sequence scanning pork experiment based on jump echo encoding.

Detailed ways

In order to make the above objects, features and advantages of the present invention more obvious and easy to understand, the specific embodiments of the present invention will be described in detail below with reference to the accompanying drawings. In the following description, numerous specific details are set forth in order to provide a thorough understanding of the invention. However, the invention can be practiced in many other ways than described herein. Those skilled in the art can make similar improvements without departing from the connotation of the present invention, and therefore the present invention is not limited to the specific embodiments disclosed below. The technical features in various embodiments of the present invention can be combined accordingly as long as they do not conflict with each other.

In the description of the present invention, it should be understood that the terms "first" and "second" are only used for distinction and description purposes, and cannot be understood as indicating or implying relative importance or implicitly indicating the number of indicated technical features. . Therefore, features defined as "first" and "second" may explicitly or implicitly include at least one of these features.

Please refer to Figure 1. The present invention provides a magnetic resonance magnetic field measurement method based on jump echo encoding. This method can actually be expressed as a magnetic resonance magnetic field measurement sequence based on jump echo encoding. The magnetic resonance magnetic field measurement sequence It is achieved by coupling the external stimulation current applied by the external electrical stimulation equipment based on the imaging sequence of the magnetic resonance equipment. The specific improvement points and design principles of the magnetic resonance magnetic field measurement sequence are described in detail below.

When measuring the B _z magnetic field, the present invention uses magnetic resonance equipment to scan the target tissue twice based on an imaging sequence that is improved based on the traditional fast spin echo (TSE) sequence. In the two scans, the external electrical stimulation device applies positive and negative switching currents of the same size and opposite starting direction to the imaging tissue in synchronization with the refocusing pulse. This current lasts until the first time in each repetition time TR. The echo signal sampling ends to ensure that the phase changes induced by the magnetic field generated by the current can continue to accumulate before the echo sampling ends. The phases of the two scan results are Φ ⁺ (x, y) and Φ ^- (x ,y).

Each repetition period of the fast spin echo sequence contains a 90° radio frequency pulse and several refocusing pulses (i.e., 180° radio frequency pulses) applied after the 90° radio frequency pulse. In the present invention, all 180° refocusing pulses in each repetition period of the fast spin echo sequence are divided into a jump echo encoding module and a fast spin echo sampling module, in which the first n 180° refocusing pulses belong to the jump echo Encoding, the remaining 180° refocusing pulses (the number is recorded as m) belong to the fast spin echo sampling part, and n is a positive integer not less than 1. In the jump echo encoding module, the 180° refocusing pulses in the module and the corresponding layer selection gradient encoding are consistent with the TSE sequence, but the difference lies in the phase gradient encoding after each 180° refocusing pulse in the jump echo encoding module. and frequency gradient encoding are cancelled; and the fast spin echo sampling module samples the signal according to the TSE sequence and fills it into k-space. Since the frequency gradient encoding after the 180° refocusing pulse in the jump echo encoding module in each repetition period is canceled, the first frequency gradient encoding in the entire repetition period is the first frequency gradient encoding in the fast spin echo sampling module. A frequency gradient encoding.

In jump-echo coding of imaging sequences, n represents the 180° refocusing applied in the jump-echo coding module The number of focal pulses. This parameter is a skipped echo parameter that needs to be set appropriately. It is used to skip the sampling of the first n echo signals so that the first echo signal collected has experienced a long enough external current action time. In order for the current applied by the external electrical stimulation device to continuously accumulate linearly in each echo, it is necessary to eliminate the influence of the stimulated echo caused by imperfect refocusing pulses in each echo. The present invention adopts the method of transforming adjacent spoiler gradients. area way.

In addition, the number m of 180° refocusing pulses in the fast spin echo sampling module is the acceleration factor of the imaging sequence. After the current ends, a conventional fast spin echo (TSE) sequence is used to collect signals. Each TR collects m echo signals to fill m rows of k space, which increases the imaging speed by m times.

Based on the above imaging sequence, after two scans under external stimulation current, the phase results of the two scans are differenced to obtain the phase difference ΔΦ(x,y) of each point position on the image, and ΔΦ(x,y) and B The linear relationship between _z solves for B _z , which is used in the calculation of MREIT or MRCDI imaging.

It can be seen that the magnetic resonance magnetic field measurement method based on jump echo encoding of the present invention extends the external current action time through jump echo encoding and improves the effectiveness of B _z . At the same time, the use of fast spin echo method to collect signals can increase the imaging speed by up to 10 times, and realize the measurement of B _z magnetic field efficiently, stably and quickly, thus improving the magnetic resonance MRCDI and MRCDI imaging performance.

In the present invention, the magnetic resonance imaging equipment and the external electrical stimulation equipment are two different equipment systems, because the two need to rely on synchronization signals to achieve coordination between the applied external current and the imaging sequence. In the present invention, the external stimulation current applied by the external electrical stimulation device can be implemented according to the following method:

Before all radio frequency pulses are applied before the first sampling echo of each TR, the magnetic resonance machine is programmed in the sequence to give synchronization pulse signals with specific widths according to three different current command (start, flip, end) requirements. The external electrical stimulation device receives this synchronization signal and programs the applied stimulation current by identifying the pulse width. The three command requirements are: a "start" synchronization signal is applied before the 90° radio frequency pulse, an "end" synchronization signal is applied before the 180° radio frequency pulse before echo signal sampling, and an "end" synchronization signal is applied before all 180° radio frequency pulses between the two. Flip” sync signal. The application of external current requires programming that takes into account the specific timing relationship between synchronization signals and RF pulses. It is important to note that in order to reduce the complex impact of external current on the results during the application of the RF pulse, the application of the external current should be zeroed during the RF pulse. The specific current application mode can be seen in Figure 1.

Traditional fast spin echo sequences usually start collecting data from the first echo signal and fill all echo signals into k-space. But in the context of the problem of B _z measurement, it is required to apply an external stimulation current to the At the end of an echo sampling, for traditional fast spin echo sequences, a long echo spacing (ESP) will cause a large loss of echo signals at the back, which greatly affects the imaging resolution in the case of a large acceleration factor (m). , while a small acceleration factor (m) weakens the acceleration effect; at the same time, a short echo spacing (ESP) will reduce the current application time, resulting in non-complementary accumulation of phases, making it difficult to measure signals with a higher signal-to-noise ratio. Therefore, in order to extend the duration of the current under the conditions of selecting a small echo spacing (ESP) and a large acceleration factor (m), the present invention designs a sequence that skips the first n echo samples, that is, the first n echoes The time is only used to apply the current accumulation phase, and the echo signal collection starts after the current ends. In this case, within a repeated cycle, the action time of the external stimulation current can be expressed by the following formula:
T _c =(ESP-τ _π )·(n+1)-0.5τ _π/2 formula (1)

Among them, ESP is the time interval between adjacent 180° refocusing pulses, and τ _π and τ _π/2 are the action times of the 180° refocusing pulse and a single 90° radio frequency pulse respectively. The specific relationship between echo signal sampling and external current application can be seen in Figure 1.

In addition, in the magnetic resonance magnetic field measurement sequence of the present invention, the way of ensuring continuous linear accumulation of current between consecutive echoes in the jump echo encoding module is improved as follows:

The stimulated echo component in the echo is eliminated by changing the adjacent spoiler gradient area. The specific method is: the first group of spoiler gradient areas is set to a value that can cause the corresponding 4π phase dispersion, and the remaining spoiler gradient areas are The gradient area is 2 or 1/2 times that of the corresponding previous group, but it must be guaranteed not to be smaller than the area of the first group of spoiler gradients. The sequence designed according to this method can realize that only the main echo signal is retained in the echo, and the excitation echo component is eliminated, so that the phase caused by the magnetic field generated by the current can be linearly accumulated in the continuous echo signal.

However, it should be noted that the spoiler gradient area of adjacent layer selection gradient encoding does not necessarily satisfy the relationship of 2 times or 1/2 times. This is just an optimal way. Theoretically, any two adjacent layer selection gradients Encoding can apply spoiler gradients with different areas.

In addition, in the magnetic resonance magnetic field measurement sequence of the present invention, echo data sampling and k-space filling are improved as follows:

Different from the traditional B _z measurement sequence based on single-echo spin echo (SE) sequence, in which each TR only collects one echo signal and fills it into k space, the sequence designed in this invention adopts the fast spin echo (TSE) sequence. For signal sampling, m is the acceleration factor. Each TR collects m echo signals to fill m rows in k space, and the scanning time is shortened by m times.

In the magnetic resonance magnetic field measurement sequence of the present invention, the B _z magnetic field can be calculated according to the following method:

In the two scanning results with opposite initial current directions, the signal at each point (x, y) on the image can be expressed as:

Among them, ρ(x,y) is the signal density, TE is the echo time, _T2 is the relaxation time of the imaging body, δ(x,y) is the intrinsic phase of the system, γ is the gyromagnetic ratio of hydrogen atoms, γ=26.75×10 ⁷ rad/(T·s), current action time T _c =(ESP-τ _π )·(n+1)-0.5τ _π/2 . Therefore, the phase difference between the two scan results can be expressed as:

Therefore, the required B _z (x, y) can be calculated by the following formula:
B _z (x,y)＝ΔΦ(x,y)/(2γ·Tc·(I ₁ -I ₂ )) Formula (4)

In the formula: I ₁ and I ₂ are the stimulation currents applied respectively during the two scanning processes, recorded as the first stimulation current and the second stimulation current. I ₁ and I ₂ need to be distinguished between positive and negative, that is, the current in the opposite direction is positive. Negative values are used to differentiate.

Based on the above description of the specific improvements and design principles of the magnetic resonance magnetic field measurement sequence shown in Figure 1, in a preferred embodiment of the present invention, a magnetic resonance magnetic field measurement method based on jump echo encoding is further provided. The specific method and process are as follows:

The same imaging sequence is used to scan the target tissue twice, and during the first scanning process, a first stimulation current with positive and negative conversion synchronized with the refocusing pulse is applied to the imaging tissue and the first scanning result is obtained, and during the second scanning process, the target tissue is scanned twice. During the scanning process, a second stimulation current with positive and negative conversion synchronized with the refocusing pulse is applied to the imaging tissue and the second scanning result is obtained; the phase difference between the first scanning result and the second scanning result is calculated, and the phase difference is combined with the B _z magnetic field The linear relationship between gives the B _z magnetic field.

Among them, the sequence formed by the temporal coupling of the synchronously applied imaging sequence and the stimulation current (the first stimulation current or the second stimulation current) is shown in Figure 1, and the details are as follows:

In each repetitive cycle of the imaging sequence, a 90° radio frequency pulse is first applied and a layer selection encoding gradient is applied at the same time, and then a pre-refocusing frequency gradient is applied to restore the spins to the same phase at the center of the echo, and then jump echo encoding and fast Spin echo sampling. In the jump echo encoding module, n 180° refocusing pulses are applied at intervals, and slice selection gradient encoding is performed while applying each 180° refocusing pulse; in the fast spin echo sampling module, m 180° refocusing pulses are applied at intervals. 180° refocusing pulse, and slice selection gradient encoding is performed while applying each 180° refocusing pulse, slice selection gradient encoding is followed by phase gradient encoding, phase gradient encoding is followed by frequency gradient encoding, and frequency gradient encoding is performed At the same time, k-space data is obtained through echo signal sampling. And in order that the phase induced by the external current can be retained in the reconstructed phase In bit images, the component of the excitation echo caused by the imperfect refocusing pulse needs to be eliminated in the sampled echo. This problem can be solved by reasonably setting the spoiler gradient in the imaging sequence, specifically: the jump echo encoding module and The n+m layer-selective gradient codes applied in the fast spin echo sampling module all set spoiler gradients on both sides. The spoiler gradient area on both sides of the first layer-selective gradient code is set to cause 4π phase dispersion. Numerical value, among n+m layer selection gradient codes, any two adjacent layer selection gradient codes apply spoiler gradients with different areas (preferably, the area of the spoiler gradient applied by two adjacent layer selection gradient codes is 2 times or 1/2 times), the spoiler gradient area on both sides of any one of the n+m level selection gradient codes is not less than the spoiler gradient area on both sides of the first level selection gradient code; n and m are positive integers not less than 1 respectively. Finally, the k-space numbers sampled from the above imaging sequence are transformed through Fourier transformation to calculate and reconstruct the amplitude image and phase image of each scan.

In each repetition cycle, the first stimulation current lasts from the first 90° radio frequency pulse to the end of the first frequency gradient encoding, and only repeats the pulse current after each radio frequency pulse (including the 90° radio frequency pulse and the 180° radio frequency pulse). (Polymer pulse) is set to zero during application. The current magnitude and positive and negative directions in a current duration segment between two adjacent radio frequency pulses are always the same. The current magnitude and positive and negative directions in any two adjacent current duration segments are the same but the positive and negative directions are opposite. ; The first stimulation current and the second stimulation current have the same number of current durations in each repetition period and the respective current durations correspond one to one, and the start and end times of each group of corresponding current durations are exactly the same. But the current is different.

As a specific implementation manner of the embodiment of the present invention, as mentioned above, the linear relationship between the phase difference and the B _z magnetic field is used to obtain a specific calculation formula for the B _z magnetic field, such as formula (4). In the formula, B _z (x, y) represents the B _z magnetic field at point (x, y), and ΔΦ (x, y) represents the difference between the first scan result and the second scan result at point (x, y). Phase difference, γ is the gyromagnetic ratio of hydrogen atoms, Tc is the duration of stimulation current in a repeated cycle, I ₁ and I ₂ are the first stimulation current and the second stimulation current that distinguish between positive and negative respectively. The calculation formula of Tc can adopt the aforementioned formula (1).

In addition, the parameters n and m in the present invention need to be optimized and adjusted according to actual conditions. Among them, n is preferably 1 to 3, and more preferably 1. m is preferably no more than 10, and more preferably 5.

It should be noted that in each repetition period, in principle, the starting and ending times of the corresponding current duration segments of each group of the first stimulation current and the second stimulation current are exactly the same but the currents are different. The difference can be The magnitude of the current is different (for example, currents of 1 mA and 2 mA are applied respectively), and the direction of the current can also be different (for example, currents of 2 mA and -2 mA are applied respectively). However, as a specific implementation of the embodiment of the present invention In the same way, the starting and ending times and sizes of the current duration segments corresponding to each group of the first stimulation current and the second stimulation current are exactly the same, but the positive and negative directions of the currents are opposite.

In another embodiment of the present invention, based on the same inventive concept, in order to implement the above magnetic resonance magnetic field measurement method based on jump echo coding, the present invention also provides a magnetic resonance magnetic field measurement device based on jump echo coding, The device includes magnetic resonance equipment, external electrical stimulation equipment and computing modules;

It should be noted that when the magnetic resonance equipment performs the imaging sequence to scan the target tissue, it sends a first synchronization signal representing "start" to the external electrical stimulation equipment before applying the 90° radio frequency pulse. Before each 180° refocusing pulse of the jump echo encoding module, a second synchronization signal representing a "flip" is sent to the external electrical stimulation device, and before the first 180° refocusing pulse of the echo signal sampling module is applied to The external electrical stimulation device sends a third synchronization signal representing "end";

After receiving the first synchronization signal, the external electrical stimulation device begins to apply a first current duration period to the target tissue according to the set time delay,

As mentioned above, in principle, the starting and ending times of the corresponding current duration segments of each group of the first stimulation current and the second stimulation current in each repetition period are exactly the same but the currents are different, but it is preferably set to each group. The starting and ending times and sizes of the corresponding current duration segments are exactly the same, but the positive and negative directions of the current are opposite.

It should be noted that after the external electrical stimulation device receives each synchronization signal, the specific electrical stimulation delay that needs to be executed needs to be based on the actual start time from the synchronization signal to the corresponding current duration segment in the sequence shown in Figure 1. Determination is based on the ability to accurately apply the corresponding external stimulation current to the target tissue according to the sequence shown in Figure 1. Moreover, if the target tissue can withstand it, the larger the external current applied, the better. Among the first stimulation current and the second stimulation current, the size of the first current duration is preferably the maximum value within the safe range of the external current allowed to be applied to the target tissue.

It should be noted that the opposite positive and negative directions of the first stimulation current and the second stimulation current can be controlled by setting the positive and negative directions of the first current duration period, that is, the external electrical stimulation equipment performs two scans of the magnetic resonance equipment. During the process, the positive and negative directions of the first continuous period of current applied after receiving the first synchronization signal are opposite. For example, the first stimulation current can be applied according to the external current sequence shown in Figure 1, that is, the current direction of the first current duration is forward; and the second stimulation current can be applied according to the reverse direction of the external current sequence shown in Figure 1 applied, that is, the current direction of the first current duration period is negative. Since other subsequent current duration segments alternate between positive and negative, as long as the current direction of the first current duration segment is opposite, the current direction of any subsequent set of corresponding current duration segments in the first stimulation current and the second stimulation current It will all be the opposite.

Those skilled in the art should know that the computing module involved in the present invention can be implemented by circuits, other hardware or executable program codes, as long as the corresponding functions can be achieved. If code is used, the code may be stored in a storage device and executed by corresponding components in the computing device. Implementations of the invention are not limited to any specific combination of hardware and software. The computing module in the above device can also be integrated into the internal computing unit of the magnetic resonance imaging equipment, or exist in the form of a separate data processing device, without limitation.

In addition, the magnetic resonance imaging equipment and external electrical stimulation equipment in the present invention can be commercially available products or self-made equipment, and can be selected according to actual user needs.

The specific technical effects will be demonstrated below based on the above method in conjunction with the embodiments, so that those skilled in the art can better understand the essence of the present invention.

Example:

The above-mentioned magnetic resonance magnetic field measurement method and device based on jump echo encoding was tested in magnetic resonance magnetic field measurement experiments on a spherical water model and pork tissue, and compared with the magnetic field measurement results of conventional multi-echo spin echo sequence weighted methods. A comparison was made. The specific scheme of the magnetic resonance magnetic field measurement method and device based on jump echo coding of the present invention is as mentioned above. The imaging sequence applied during the first scan and the first stimulation current applied simultaneously are shown in Figure 1. The two There is a 2-fold or 1/2-fold relationship between the spoiler gradient areas applied by selective gradient encoding at adjacent levels; while the imaging sequence applied during the second scan is the same as the first scan, but the second stimulation current is different from the first stimulation The starting and ending times and sizes of the current duration segments corresponding to each group in the current are exactly the same, but the positive and negative directions of the current are opposite. The specific imaging sequence and stimulation current will not be described in detail here. The following will only introduce The specific parameters here are introduced. In this embodiment, the skipped echo parameter n=1 and the acceleration factor m=5. The experimental results in this embodiment are shown in Figure 2, Figure 3 and Figure 4:

As can be seen from Figure 2, in the magnetic resonance magnetic field measurement experiment on the water model, the measurement results of the magnetic resonance magnetic field measurement method based on jump echo encoding are highly consistent with the magnetic field measurement results of the conventional multi-echo spin echo sequence weighted method. At the same time, the measurement time is shortened by 5 times. The quantitative analysis further illustrates the effect of the proposed sequence on the magnetic field measurement problem (regression result r=0.98, s=1.02, p<0.001; Bland-Altman Plot result shows 97 % points are within the confidence interval), indicating the effectiveness of the present invention.

It can be seen from Figure 3 and Figure 4 that the magnetic resonance magnetic field measurement method based on jump echo coding can provide a wide range of imaging coverage within the clinically allowed measurement time (z is different) whether it is used on water models or biological tissues. level), and no obvious artifacts were found in either the magnetic field measurement results or the amplitude results, which further proves the effectiveness of the present invention.

The above-described embodiment is only a preferred solution of the present invention, but it is not intended to limit the present invention. Those of ordinary skill in the relevant technical fields can also make various changes and modifications without departing from the spirit and scope of the present invention. Therefore, any technical solution obtained by adopting equivalent substitution or equivalent transformation shall fall within the protection scope of the present invention.

Claims

A magnetic resonance magnetic field measurement method based on jump echo coding, which is characterized by:

The same imaging sequence is used to scan the target tissue twice, and during the first scanning process, a first stimulation current with positive and negative conversion synchronized with the refocusing pulse is applied to the imaging tissue and the first scanning result is obtained, and during the second scanning process, the target tissue is scanned twice. During the scanning process, a second stimulation current with positive and negative conversion synchronized with the refocusing pulse is applied to the imaging tissue and the second scanning result is obtained; the phase difference between the first scanning result and the second scanning result is calculated, and the phase difference is combined with the B z magnetic field The linear relationship between the B z magnetic field;

The imaging sequence divides all 180° refocusing pulses in each repetition period in the fast spin echo sequence into a jump echo encoding module and a fast spin echo sampling module, in which the first n 180° refocusing pulses belong to the jump echo encoding module. Wave encoding, the remaining 180° refocusing pulse belongs to the fast spin echo sampling part, n is a positive integer not less than 1; in the jump echo encoding module, the 180° refocusing pulse in the module and its corresponding slice selection gradient The encoding is consistent with the fast spin echo sequence, but the phase gradient encoding and frequency gradient encoding after each 180° refocusing pulse in the module are canceled; the fast spin echo sampling module samples the signal according to the fast spin echo method and Filled into k space, the first frequency gradient code in each repetition period is the first frequency gradient code in the fast spin echo sampling module;

In each repetition period, the first stimulation current lasts from the first 90° radio frequency pulse to the end of the first frequency gradient encoding, and is only set to zero during the application of each 90° and 180° radio frequency pulse. The current magnitude and positive and negative directions in a current duration segment between two adjacent radio frequency pulses are always the same, and the current magnitudes and positive and negative directions in any two adjacent current duration segments are the same but the positive and negative directions are opposite; the first stimulation current and the said The second stimulation current has the same number of current duration segments in each repeated cycle, and the respective current duration segments correspond one to one. The start and end times of each group of corresponding current duration segments are exactly the same but the currents are different.
The magnetic resonance magnetic field measurement method based on jump echo coding as claimed in claim 1, characterized in that: the imaging sequence is specifically implemented in the following manner:

In each repetition period of the imaging sequence, a 90° radio frequency pulse is first applied and a layer selection encoding gradient is applied at the same time, and then a pre-refocusing frequency gradient is applied to restore the spins in phase at the center of the echo, and then jump echo encoding is performed. module and a fast spin echo sampling module; in the jump echo encoding module, n 180° refocusing pulses are applied at intervals, and layer selection gradient encoding is performed while applying each 180° refocusing pulse; the fast auto-rotation In the echo wave sampling module, m 180° refocusing pulses are applied at intervals, and layer selection gradient encoding is performed while applying each 180° refocusing pulse. Layer selection gradient encoding Then perform phase gradient encoding, and then perform frequency gradient encoding after phase gradient encoding, and obtain k-space data through echo signal sampling while frequency gradient encoding; and in the jump echo encoding module and fast spin echo sampling module Spoiler gradients are set on both sides of the applied n+m level selection gradient codes, where the spoiler gradient area on both sides of the first level selection gradient code is set to a value that can cause 4π phase dispersion. n+m level selections In gradient coding, any two adjacent levels select gradient coding to apply spoiler gradients with different areas. Among n+m level selection gradient coding, the spoiler gradient area on both sides of the gradient coding in any one level selection gradient coding is not less than that of the first level selection. The spoiler gradient area on both sides of the gradient encoding; finally, the sampled k-space numbers are calculated through Fourier transform to reconstruct the amplitude image and phase image of each scan; n and m are positive integers not less than 1 respectively;
The magnetic resonance magnetic field measurement method based on jump echo coding according to claim 2, characterized in that: the spoiler gradient areas on both sides of any two adjacent level selection gradient codes in the n+m layer selection gradient codes The relationship is preferably 2 or 1/2 times.
The magnetic resonance magnetic field measurement method based on jump echo encoding as claimed in claim 1, characterized in that: the calculation formula of using the linear relationship between the phase difference and the Bz magnetic field to obtain the Bz magnetic field is:
B z (x,y)=ΔΦ(x,y)/(2γ·Tc·(I 1 -I 2 ))

Where: B z (x, y) represents the B z magnetic field generated by unit external current at point (x, y), unit: nT/mA, ΔΦ (x, y) represents the first scan result and the second The phase difference of the scanning result at point (x, y), γ is the gyromagnetic ratio of hydrogen atoms, Tc is the duration of the stimulation current in a repeated cycle, I 1 and I 2 are the first stimulation current and the second stimulation current that distinguish between positive and negative respectively. 2. Stimulation current.
The magnetic resonance magnetic field measurement method based on jump echo coding according to claim 4, characterized in that: the calculation formula of the duration Tc of the stimulation current in one repetition cycle is:
T c =(ESP-τ π )·(n+1)-0.5τ π/2

Among them: ESP is the time interval between adjacent 180° refocusing pulses, τ π and τ π/2 are the action times of a single 180° refocusing pulse and a single 90° radio frequency pulse respectively.
The magnetic resonance magnetic field measurement method based on jump echo coding according to claim 1, characterized in that: the n is preferably 1 to 3, more preferably 1; the m is preferably no more than 10, further preferably 5 .
The magnetic resonance magnetic field measurement method based on jump echo coding according to claim 1, characterized in that: in each repetition period, each group of the first stimulation current and the second stimulation current corresponds to The start and end times and sizes of the current duration segments are exactly the same, but the positive and negative directions of the current are opposite.
The magnetic resonance magnetic field measurement method based on jump echo coding according to claim 1, characterized in that: among the first stimulation current and the second stimulation current, the size of the first current duration segment is preferably the target tissue. The maximum value within the safe range of external current that is allowed to be applied.
A magnetic resonance magnetic field measurement device based on jump echo coding for implementing the magnetic resonance magnetic field measurement method according to any one of claims 1 to 8, characterized in that it includes magnetic resonance equipment, external electrical stimulation equipment and a computing module;

The external electrical stimulation device is used to apply the first stimulation current and the second stimulation current to the target tissue during scanning of the magnetic resonance device;

The magnetic resonance equipment is used to perform the imaging sequence and obtain first scan results and second scan results;

The calculation module is used to calculate the phase difference between the first scan result and the second scan result, and use the linear relationship between the phase difference and the Bz magnetic field to obtain the Bz magnetic field.
The magnetic resonance magnetic field measurement device based on jump echo encoding according to claim 9, wherein the magnetic resonance equipment performs the imaging sequence to scan the target tissue before applying a 90° radio frequency pulse. Send a first synchronization signal to the external electrical stimulation device, send a second synchronization signal to the external electrical stimulation device before applying each 180° refocusing pulse of the jump echo encoding module, and before applying the echo signal sampling module Send a third synchronization signal to the external electrical stimulation device before the first 180° refocusing pulse;

After receiving the first synchronization signal, the external electrical stimulation device begins to apply the first current duration to the target tissue according to the set time delay;

Each time the external electrical stimulation device receives the second synchronization signal, it applies the next current duration segment in the opposite positive and negative direction to the previous current duration segment to the target tissue according to the set time delay;

After receiving the third synchronization signal, the external electrical stimulation device applies the last current duration period in the opposite positive and negative direction to the previous current duration period to the target tissue according to the set time delay.