CN106821500B

CN106821500B - Navigation system for minimally invasive surgery

Info

Publication number: CN106821500B
Application number: CN201710167799.7A
Authority: CN
Inventors: 罗会俊
Original assignee: Nanjing Tuobao Medical Technology Co ltd
Current assignee: Nanjing Tuobao Medical Technology Co ltd
Priority date: 2017-03-21
Filing date: 2017-03-21
Publication date: 2023-11-17
Anticipated expiration: 2037-03-21
Also published as: CN106821500A

Abstract

The invention discloses a navigation system and a navigation method for minimally invasive surgery. The system features to be modified aiming at minimally invasive surgery and intra-operative navigation mainly relate to an open U-shaped magnet structure, a magnet gap of more than 500mm, a radio frequency receiving and transmitting coil designed as a biplane circular polarization type, a gradient coil designed as a biplane main coil and a biplane axial shielding coil added on the outer side; and relates to a clinically applicable navigation control procedure and an intra-operative navigation imaging method. The structure and the method ensure the system opening degree, the safety and the convenience required by the minimally invasive surgery; in particular, in order to improve the quality and the real-time performance of the magnetic resonance imaging and ensure the accurate positioning of surgical instruments and the accurate control of surgical routes.

Description

Navigation system for minimally invasive surgery

Technical Field

The invention relates to the technical field of magnetic resonance, in particular to a navigation system for minimally invasive surgery.

Background

The accurate minimally invasive treatment technology is beneficial to improving the curative effect and relieving the pain of patients, and is increasingly applied to clinical medicine. Radiation therapy techniques such as laser, X-ray and gamma ray, high-intensity focused ultrasound techniques, interventional therapy techniques such as thermal ablation and cryoablation, and other various minimally invasive operations all require accurate positioning of lesions by means of image guidance techniques and real-time accurate monitoring of the range of action and the efficacy of electromagnetic waves, ultrasound waves or surgical instruments on targets during treatment. Compared with the imaging technologies such as ultrasound and CT, the Magnetic Resonance Imaging (MRI) technology not only has the advantages of high resolution and multi-azimuth and multi-parameter imaging, but also can clearly display the boundary, nerves and blood vessels of an anatomical structure, even monitor the physiological activity of a human body in real time and detect the temperature of metabolites and target areas, and has no damage and radiation injury to the human body, so the MRI technology is particularly suitable for image navigation.

Modern MRI-guided stereotactic brain surgery and other interventional therapy systems are typically composed of minimally invasive surgical instruments (or surgical robots), an optical tracking positioning subsystem, and an MRI image guidance and navigation subsystem. For navigation in MRI, magnet opening degree, scanning speed and image artifact are the most important technical indexes, the image signal to noise ratio and resolution are required to accurately display the target position and the medical instrument position, and temperature distribution monitoring is required to be accurate and reliable for interventional therapy such as thermal/cryoablation. Because of the above technical requirements, there are few MRI image guidance products that have been in clinical use so far, mainly open medium-low field MRI systems. For example, sign SP (0.5T) of general electric company in the united states is a midfield superconducting open system, which generates a horizontal magnetic field by vertically placing two superconducting coils made of niobium-tin alloy material, increasing the width of the accessible region; magnetom Open 0.2T from siemens, germany and proview0.23T Open from philips, netherlands are low field normally Open systems, with vertical magnetic fields generated by upper and lower resistive magnets, allowing access to the patient from one side and performing interventional procedures in the near 280 range. Such image navigation systems have low requirements for magnetically compatible surgical instruments and imaging quality and speed that can meet general clinical interventional therapy requirements, such as biopsy or minimally invasive surgical treatment of prostate cancer, uterine fibroids, and joint diseases. While high field or ultra high field MRI systems provide the necessary high signal-to-noise ratio or high resolution conditions for many important imaging methods and techniques applications, the cylindrical enclosure structure and high cost of superconducting magnets severely limit MRI technology applications in interventional procedures and minimally invasive procedures. For a long time, superconducting MRI interventional therapy systems are mainly used for preoperative operation route planning and postoperative efficacy evaluation, and in intraoperative navigation, a time-sharing scanning scheme is often used, for example, the GE IntraOp sign HD 3T system needs to be carried back and forth between a shielding room and an operating room through a sliding rail during scanning and surgical operation, instrument mark tracking needs to be realized by means of a stereotactic technology (such as an optical tracking technology) and through conversion of a physical space and a logical space (i.e., an image domain), and even an intraoperative displacement error of a tissue organ needs to be corrected by developing a correction algorithm which is complicated and has to be operated with high efficiency, so that the application of such high-field MRI systems in interventional therapy is considerably limited. Radiation therapy is relatively less subject to the closed structure of the superconducting magnet, which may allow the superconducting MRI system to be used in combination with a linac to effect intra-operative navigator interventions, but requires safe and effective control of the dose and targeting accuracy of tumor radiotherapy. In any event, the development of modern magnetic resonance imaging devices into open architecture is a major trend, which not only improves the comfort of examination for patients (particularly claustrophobic patients), but also opens up a wide space for intraoperative MRI image guidance applications.

Since the 90 s of the last century in China, a small number of permanent magnet MRI systems have been used for interventional therapy clinical trials of medical research institutions, but basically the simple combination of the existing diagnostic permanent magnet imaging systems and a general optical tracking and positioning instrument has not yet developed a special navigation system with unique MRI technical advantages and an imaging method suitable for navigation, which are widely applied to minimally invasive surgery and interventional therapy.

Disclosure of Invention

Aiming at overcoming the opening degree limitation of the existing imager magnet and coil structure and aiming at the biopsy of craniocerebral and joint diseases and the treatment needs of minimally invasive surgery (such as drug injection, puncture drainage and the like), the invention provides an imaging magnet structure of a special magnetic resonance image navigation system with high opening degree and accurate positioning.

The invention provides a navigation system for minimally invasive surgery, comprising: the magnetic field sensor comprises a magnetic yoke, magnetic steel, a polar plate, a shielding coil, a gradient coil, a radio frequency coil, a scanning bed, a heating rod, a heating plate, a temperature sensor, a shell, a T/R receiving switch, a gradient power amplifier, a radio frequency power amplifier, a preamplifier, a temperature control unit, an optical tracking and positioning system, a coil tuning control unit, a scanning bed control unit, a navigation imaging control unit, a system control interface, a host and a display, and is characterized in that a magnet is of an open U-shaped structure, a magnet gap is larger than 500mm, and the radius of the polar plate is 350-400 mm; the radius of the polar plate is 380mm; the field intensity of the magnet is between 0.2T and 0.7T; the radio frequency coil of the magnet is of a biplane circular polarization structure, and a plurality of nonmagnetic capacitors are arranged between the outer conductor ring and the inner conductor ring; the radio frequency receiving and transmitting coil realizes the conversion of transmitting and receiving through a T/R switch; the number of the non-magnetic capacitors is more than 100; two groups of X, Y and Z axis gradient coils are respectively arranged, one group is arranged on the inner side of one magnetic pole, the other group is arranged on the inner side of the other magnetic pole, and the coil planes are parallel to the surface of the polar plate and are separated from the polar plate through an anti-vortex plate; the Z-axis gradient coils are composed of a main gradient coil and an axial shielding coil, belong to MAXWELL coils, are in a concentric circle type, have 13 turns of the main gradient coil and 14 turns of the axial shielding coil, and are connected to a gradient power amplifier through a gradient power filter by 6 gradient cables; the main gradient coil and the axial shielding coil are manufactured by adopting a printed circuit board technology; the axial shielding coil is arranged outside the main gradient coil and is close to the polar plate, gradient current directions of the axial shielding coil and the main gradient coil are opposite, and magnetic field gradients are limited in the circumference of the axial shielding coil, so that thrust generated by interaction of current pulses and static magnetic fields in the coils is counteracted; an insulating layer is arranged between the main gradient coil and the axial shielding coil; the primary gradient coils, the axial shield coils and the insulating layer preferably have a radius of 350mm, each coil having a thickness of 4mm to 5mm, and the primary gradient coils and the axial shield coils have a spacing of 2mm to 3mm.

The invention also provides a navigation method for minimally invasive surgery, which adopts the navigation system for minimally invasive surgery according to claim 1, and comprises the following steps:

step 100: the magnetic susceptibility artifact or metal artifact of the non-magnetic medical instrument is detected and the navigation sequence parameter is optimized in advance according to the operation flow, a three-dimensional high-resolution T1 weighted image is collected on a navigation imager or a high-field imager before operation, scanning enhancement, angiography or functional imaging and the like are added when necessary, and then operation route planning is carried out, including target setting, needle insertion path planning and operation scheme planning.

Step 200: then, the scanning part is placed in an isocenter area through the calibration of the horizontal plane height of a scanning bed and the guidance of a laser marking line of a laser positioning instrument on a navigation imaging system, a dynamic local shimming technology is selected for shimming, three-dimensional space selective radio frequency pulses consisting of three selective layer gradients and three sine waveform pulses excite protons or other magnetic atomic nuclei in a small volume, then free induction attenuation signals FID are acquired, the sequence repeatedly operates, and the magnetic field uniformity is continuously adjusted through linear gradients or shimming gradients until the frequency spectrum integration area is maximum, so that the magnetic field uniformity of an imaging area is optimized.

Step 300: then the navigation flow establishes an internal reference coordinate system through a navigation module and a navigation interface and controls a local rapid navigation sequence to run, fault signals in three orthogonal directions are repeatedly collected according to a surgical route, a single layer can be collected in each direction, multiple layers can be synchronously collected, or a real-time path tracking mode is adopted, a large-view positioning image is firstly obtained through scanning of navigation software, the center position of a scanning layer is firstly positioned at a surgical entrance position on the positioning image, a plurality of groups of scanning layers are planned towards a target position, each group of layers can be arranged at different positions to avoid nerves or blood vessels, and the interval between the layers is set to be half of the thickness of the layers; the positioning parameters and waveform parameters are continuously updated in the operation process and scanned in real time so as to continuously track and position the surgical instrument.

The local rapid navigation sequence design mode is characterized in that radio frequency pulses have three-dimensional space selectivity, layer selection gradients corresponding to a first radio frequency excitation pulse, a first radio frequency refocusing pulse and a second radio frequency refocusing pulse are arranged in different directions, an SLR pulse waveform with an excitation profile highly optimized is adopted, a plurality of adjacent frequency bands can be selectively and uniformly excited at the same time, the range and the interval of each frequency band can be accurately adjusted on a positioning image through layer thickness and layer spacing, the amplitude of the pulse excitation profile can be corrected and consistent according to the integral area of each frequency band, a part of k-space lines can be repeatedly acquired and pulse phase circulation can be carried out, the frequency encoding gradient and the layer selection gradient can be simultaneously applied during data acquisition, the bandwidth of a receiver is set to be 100kHz or higher, and k-space data accumulation and partial Fourier image reconstruction can be carried out. The navigation imaging method not only can improve the uniformity, resolution and signal-to-noise ratio of the image, but also can inhibit susceptibility artifacts and metal artifacts and shorten the scanning time;

Step 400: after the navigator treatment is finished, local high resolution scanning and enhanced scanning are performed on a navigator imager or a high field imager to confirm the curative effect. The acquisition matrix of the frequency encoding and phase encoding directions is doubled here to further improve the image resolution compared to intra-operative navigator scans.

Preferably, the navigation process includes the step of dynamic local shimming: the scanning position is arranged in an isocenter area through the calibration of the plane height of a scanning bed and the guidance of a laser marking line of a laser positioning instrument, a dynamic local shimming technology is selected for shimming, three layer-selecting gradients and sine or SLR radio frequency pulses are adopted for exciting protons or other magnetic atomic nuclei in a small volume in sequence, then free induction attenuation signals are collected, the sequence is operated repeatedly, and meanwhile, the magnetic field uniformity is adjusted through linear gradients or shimming gradients until the frequency spectrum integration area is maximum, so that the magnetic field uniformity of an imaging area is optimized.

Preferably, tracking the position of the surgical instrument is accomplished by metal artifact testing and calibration steps: firstly, testing magnetic field distribution graphs based on a standard water model and a gradient echo sequence of a copper sulfate water solution, then adding a strip-shaped or cross-shaped nonmagnetic metal material into the central area of the standard water model, testing the magnetic field distribution graphs through the gradient echo sequence, calculating the difference between the two magnetic field distribution graphs, selecting the nonmagnetic metal material or the composite material with the smallest difference to manufacture a surgical instrument, and finally measuring the size of a non-signal area on a gradient echo image and carrying out consistency verification with the actual size of a metal strip to be used as an accurate mark of the position of the surgical instrument.

Preferably, a proton magnetic resonance signal in a small volume is excited by using a radio frequency pulse with a multi-azimuth layer selection gradient and an excitation profile highly optimized, wherein a first 90-degree radio frequency excitation pulse and a first 180-degree refocusing pulse adopt an SLR pulse waveform which is uniformly excited, and the layer selection gradient amplitude corresponding to the first radio frequency pulse is set as G in a navigation sequence _s1 ＝2πΔf/γ/FOV _x Setting the gradient amplitude of the layer selection corresponding to the second radio frequency pulse as G _s2 ＝2πΔf/γ/FOV _y Setting the gradient amplitude of the layer selection corresponding to the third radio frequency pulse as G _s3 =2pi Δf/γ/THK, where FOV _x 、FOV _y And THK can be directly regulated on the sequence parameter table.

Preferably, the rf pulse is preferably an SLR pulse, and the waveform characteristic parameters thereof are set as follows: the time-band product TBP is 4 or 8, the out-band ripple coefficient and the in-band ripple coefficient are not more than 0.5%, the flip angle of the excitation pulse is pi/2, and the selected layer gradient strength is set according to the required layer thickness, for example, 0.7G/cm; or the time-zone product TBP is 16, the out-band ripple coefficient and the in-band ripple coefficient are not more than 0.5%, the refocusing pulse flip angle is pi, and the selected layer gradient strength is 0.5G/cm; and optimizing excitation profile based on synchronous scanning calibration sequence under synchronous multilayer excitation condition, wherein the sequence is characterized in that radio frequency pulse adopts synchronous excitation pulse waveform, layer gradient and frequency coding gradient are arranged in the same direction, k-space data are acquired at echo time, and excitation profile and amplitude image are calibrated according to the following mode:

A. The SLR pulse waveform of each phase code is selected and used with the synchronous scanning calibration sequence shown in FIG. 22

The signal is acquired and fourier transformed into the frequency domain to obtain excitation profiles, denoted C respectively ₁ 、C ₂ And C ₃ ；

B. Calculation C ₁ 、C ₂ And C ₃ Is denoted by k respectively ₁ ＝∫∫C ₁ dxdy、k ₂ ＝∫∫C ₂ dxdy and k ₃ ＝∫∫C ₃ dxdy, where ≡ ≡denotes two-dimensional integral, x and y denote two-dimensional coordinates;

C. calibration profile C ₁ 、And->

D. Also, the two-dimensional image obtained by three scans is calculated according to the integral area ratio k ₁ 、k ₂ And k ₃ Performing calibration and decomposing images of each layer according to a formula (4);

although the above approach can be directly applied to the three-layer simultaneous excitation case, it is apparent that the method can be generalized to the calibration of other multi-layer simultaneous excitation images in a similar manner.

Preferably, in the case of metal artifact interference, the receiver bandwidth is preferably above 100kHz, the bandwidth Δf of the radio frequency pulse needs to be optimized in advance, Δf can be set to different values according to the operation flow shown in fig. 7, Δf when the magnetic susceptibility artifact or the metal artifact is minimum is searched from the bandwidth Δf, and when the frequency encoding gradient and the layer selecting gradient are applied simultaneously during data acquisition, the optimal value of Δf is close to the single echo acquisition time length, otherwise Δf is preferably 2kHz or higher.

Preferably, the navigation sequence Local-SE-NV is selected The navigation imaging sequence adopts three-azimuth slice selection gradients and SLR pulses to realize Local excitation on the basis of a spin echo sequence, or adopts the navigation sequence Local-ME-NV, the navigation imaging sequence adopts the three-azimuth slice selection gradients and SLR pulses to excite magnetic resonance signals in a small visual field, then a plurality of gradient echoes are acquired during the application of a frequency coding gradient with alternating positive and negative polarities, and T1 weighted imaging can be realized rapidly by minimizing the echo time and the optimal phase coding steps to be 32 or 64 or other smaller integers. Here, a layer gradient G _s1 、G _S2 And G _S3 Is set to be orthogonal to the orientation of the orientation image G _s1 ，G _s2 And G _s3 Representing different azimuthal slice gradients, the first two of which are perpendicular or parallel to the surgical path and the other of which is parallel or perpendicular to the surgical path, the other slice gradients being of azimuthal and amplitude and G _s3 The same, signal acquisition starts from the third radio frequency pulse later, k space filling and image reconstruction adopt a partial Fourier mode; the Local imaging sequence Local-ME-NV is characterized by applying a G at the moment of switching the positive and negative polarities of the read gradient when the scan speed is prioritized _blip The gradient, so that the data acquisition part is equivalent to echo plane acquisition, the number of echoes in the dashed line frame can be set to be 32 or other smaller integers, and the acquired data is processed according to the common phase correction and image reconstruction modes of echo plane imaging.

Preferably, a navigation sequence Local-HASTE-NV is selected, the navigation imaging sequence adopts three-azimuth slice selection gradients and SLR pulses to realize Local excitation on the basis of single excitation of a fast spin echo sequence, and a half Fourier acquisition mode is adopted to rapidly realize T2 weighted imaging; here, the echo time is set in the range of 80ms to 140ms, and the slice gradient G is selected _s1 、G _S2 And G _S3 Is arranged in an orthogonal direction on the positioning image, wherein the first two are perpendicular (or parallel) to the surgical path, the other are parallel or perpendicular to the surgical path, and the other gradient of the selected layer is oriented with G _s3 The same; and, here, additionally applying a compensation gradient G in the layer-selecting direction _sc1 、G _sc2 、G _sc3 …G _scn The gradient amplitude is equal to G _S3 Identical toThe gradient width is equal to G _r1 The same; first pulse waveform characteristic parameter: minimum phase SLR, time-banded TBP of 8, in-banded ripple coefficient of 0.5%, out-banded ripple coefficient of 0.1%, pulse flip angle of pi/2, selected layer gradient strength set according to desired layer thickness, e.g., 0.7G/cm; second pulse waveform characteristic parameter: minimum phase SLR, time-banded TBP of 16, in-banded ripple coefficient of 0.5%, out-banded ripple coefficient of 0.1%, pulse flip angle of pi, selected layer gradient strength set according to the desired layer thickness, e.g., 0.5G/cm; third pulse and subsequent pulse waveform characteristic parameters: minimum phase SLR, time-band product TBP of 16, in-band ripple factor of 0.5%, out-of-band ripple factor of 0.1%, pulse flip angle of pi, selected layer gradient strength of 0.5G/cm, or (pulse waveform parameters of FIG. 26 need to be described herein), linear phase SLR, time-band product TBP of 16, in-band and out-of-band ripple factors of 0.1%, pulse flip angle of pi, selected layer gradient strength of 0.5G/cm.

The signal acquisition starts after the third radio frequency pulse and the k-space filling and image reconstruction takes place in a partial fourier manner.

Preferably, the radio frequency pulse has spatial selectivity, and can uniformly excite a plurality of adjacent frequency bands at the same time, and the radio frequency pulse waveform of multi-layer synchronous imaging at any position accords with the following formula:

in the above formula, SLR (t) represents a linear phase SLR pulse, gamma represents a gyromagnetic ratio,and->Respectively represent pulse code phase, G _s Represents the gradient of the selected layer, r _i (i=0-n) represents a set of space vectors of equally spaced synchronous excitation levels; setting corresponding positioning line on positioning image through user interface> Transmitting azimuth parameters of each layer to a pulse waveform calculation program of a navigation module, outputting updated pulse waveforms, loading the updated pulse waveforms to a navigation imaging control unit, then operating a navigation sequence Local-MSME-NV, realizing Local uniform excitation by adopting three layer-selecting gradients and the pulse waveforms defined by the sequence, acquiring multi-echo signals under the action of frequency coding gradients with alternating positive and negative polarities, and then reconstructing a real-time image; for the three-layer synchronous excitation condition, the radio frequency excitation pulse waveform is calculated according to the following formula:

r in the above ₀ Space vectors representing synchronous excitation levels, ΔS representing synchronous excitation level spacing, the excitation profile can be optimized to achieve uniform excitation of magnetic resonance signals at three levels in three immediately adjacent frequency bands by adjusting waveform parameters in the above equation, e.g., G here _s The bandwidth of the SLR pulse is set to 1kHz, the time-band product TBP is set to 4, and the in-band and out-of-band ripple coefficients are not more than 0.5% based on the required layer thickness set to 1.5G/cm, three scansAnd->Are respectively arranged as And->Refocusing pulses in spin echo acquisition modePreferably 180 DEG minimum phase SLR pulse or linear phase SLR pulse, G _s The value is not more than the maximum gradient amplitude, and is set to be 2.0G/cm, and the bandwidth requirement of SLR pulse is not less than (3. THK+2. Delta. S). Gamma.G _s /(2pi), here set to 4kHz, where THK represents the thickness of the layer corresponding to each band, the time-band product TBP of the SLR pulse is set to 8 or 16, the in-band ripple coefficient is set to 0.1%, and the out-of-band ripple coefficient is set to 0.1%; each time scanning acquires a part of k-space lines, and three times of scanning obtain an image S ₁ Image S ₂ And image S ₃ Finally, carrying out image decomposition according to the following formula to obtain images of each layer:

s 'in formula (4)' ₁ ,S' ₂ And S' ₃ Representing images corresponding to the three levels of simultaneous excitation.

When the scanning speed is prioritized, a G is applied at the moment of switching the positive and negative polarities of the read gradient _blip The gradient, so the data acquisition part is equivalent to echo plane acquisition, the number of echoes in the dotted line frame can be set to be 32 or other smaller integers, the acquired data is processed according to the common phase correction and image reconstruction modes of echo plane imaging, and then the image decomposition is carried out according to the formula (4) to obtain images of each layer.

The beneficial effects are that: the structure of the invention ensures the system opening degree, safety and convenience required by the minimally invasive surgery; in particular, in order to improve the quality and the real-time performance of the magnetic resonance imaging and ensure the accurate positioning of surgical instruments and the accurate control of surgical routes.

Drawings

Fig. 1 is a block diagram of a navigation system for minimally invasive surgery according to an embodiment of the present invention.

Fig. 2 is a sectional view taken along the direction A-A in fig. 1.

Fig. 3 is a schematic diagram of a planar rf coil according to an embodiment of the present invention.

Fig. 4 is a diagram of a gradient coil structure according to an embodiment of the present invention.

Fig. 5 is a schematic diagram of a winding manner of the main coil and the axial shielding coil according to the embodiment of the present invention.

Fig. 6 is a schematic diagram of an MRI image guided informationized operating room in accordance with an embodiment of the present invention.

Fig. 7 is an instrument detection flow in accordance with an embodiment of the present invention.

Fig. 8 is a local dynamic shimming sequence according to an embodiment of the present invention: g _s1 ，G _s2 ，G _s3 Logical gradients representing different orientations whose orientation can be set by positioning the image.

FIG. 9 is a navigation workflow of an embodiment of the present invention.

FIG. 10 is a functional block diagram of navigation software according to an embodiment of the present invention.

FIG. 11 is a navigation cross section and parameters of an embodiment of the present invention: the left hand diagram is a navigation interface including "surgical assessment", "path planning" and "navigation scan" sub-interfaces corresponding to 1031, 1032 and 1033 functional modules, respectively, in fig. 10. The right hand graph is a navigation parameter table, in the form of a floating window, for modifying the sequence parameters and the reconstruction parameters.

Fig. 12 is a schematic diagram of a real-time path tracking mode according to an embodiment of the present invention. The rectangle represents the scan slice, and the circle represents the scan slice center position.

FIG. 13 is a Local spin echo navigator sequence (Local-SE-NV) according to one embodiment of the present invention: g _s1 ，G _s2 And G _s3 Represents the gradient of the selected layer in different directions, and the gradient G of the selected layer is shown by a dotted line _s4 As an option Δt represents the echo time.

Fig. 14 shows a (left) minimum phase SLR excitation pulse waveform and a (right) minimum phase SLR excitation pulse excitation profile according to an embodiment of the present invention. Waveform characteristic parameters: the time-band product TBP is 8, the in-band ripple coefficient is 0.5%, the out-of-band ripple coefficient is 0.1%, the pulse flip angle is pi/2, and the gradient strength of the selected layer is 0.7G/cm.

FIG. 15 is a (left) minimum phase SLR refocusing pulse waveform according to an embodiment of the present invention; the (right) minimum phase SLR refocuses the pulse excitation profile. Waveform characteristic parameters: the time-band product TBP is 16, the in-band ripple coefficient is 0.5%, the out-of-band ripple coefficient is 0.1%, the pulse flip angle is pi, and the gradient strength of the selected layer is 0.5G/cm.

FIG. 16 is an intra-operative navigator sequence (Local multi-echo sequence, local-ME-NV) according to one embodiment of the present invention: g _s1 ，G _s2 ，G _s3 Representing the gradient of layers in different directions G _s4 ，G _s5 ，G _s6 ......G _sn As an alternative to represent the applied layer gradient during sampling, G _p Representing phase encoding gradient, G _pre Representing a preliminary read gradient, G _r1 ，G _r2 ，G _r3 ......G _rn Representing the first to n read gradients, adjacent read gradients are equal in area but opposite in polarity, the 90 RF pulse is preferably the minimum phase SLR pulse shown in FIG. 14, and the 180 RF pulse is preferably the minimum phase SLR pulse shown in FIG. 15.

FIG. 17 is an intraoperative navigator sequence (three-layer synchronous Local multi-echo sequence, local-MEEPI-NV) according to an embodiment of the present invention: g _blip The phase encoding gradient is applied to the positive and negative polarity switching time of the frequency encoding gradient, and other parameters are the same as those of fig. 16.

FIG. 18 is an intraoperative navigation sequence (three-layer synchronous Local multi-echo sequence, local-MSME-NV) of an embodiment of the present invention: g _s1 ，G _s2 ，G _s3 Representing the gradient of layers in different directions G _s4 ，G _s5 ，G _s6 ......G _sn As an alternative to represent the applied layer gradient during sampling, G _p Representing phase encoding gradient, G _pre Representing a preliminary read gradient, G _r1 ，G _r2 ，G _r3 ......G _rn Representing first to n read gradients, adjacent read gradients are equal in area but opposite in polarity, the first rf pulse is preferably a three-layer simultaneous excitation SLR pulse as shown in fig. 15, the second and third rf pulses are selected from 180 deg. refocusing SLR pulses as shown in fig. 15 or 26, and the selected layer gradient amplitude is appropriately adjusted within the imaging sequence according to the desired layer thickness.

FIG. 19 is an intraoperative navigation sequence (three-layer synchronous Local multi-echo sequence, local-MSPI-NV) of an embodiment of the present invention: g _blip The phase encoding gradient is applied to the positive and negative polarity switching time of the frequency encoding gradient, and other parameters are the same as those of fig. 18.

Fig. 20 shows a three-layer simultaneous excitation rf pulse waveform (left) and excitation profile (right) according to an embodiment of the present invention. The dotted line in the right figure represents C ₁ ' Star line indicates C ₂ ' solid line represents C ₃ ’。

FIG. 21 is a plot of the excitation profile of the pre-calibration tri-layer simultaneous excitation pulse according to an embodiment of the present invention. In the figure, the dotted line represents C ₁ ' Long and short dash line indicates C ₂ ' solid line represents C ₃ ’。

FIG. 22 is a multi-layer synchronous scanning calibration sequence according to an embodiment of the present invention: the radio frequency pulse is multilayer synchronous excitation pulse, G _s1 、G _s2 And G _s3 Representing the layer selection gradient.

FIG. 23 is a plot of a three-layer synchronized excitation pulse excitation profile after calibration in accordance with an embodiment of the present invention. In the figure, the dotted line represents C ₁ ' Long and short dash line indicates C ₂ ' solid line represents C ₃ ’。

Fig. 24 is a (left) five-layer synchronous excitation pulse waveform of an embodiment of the present invention, the solid line represents the real part of the waveform, and the broken line represents the imaginary part of the waveform; and (right) five layers of synchronous excitation profiles, wherein different types of curves in the figure correspond to different layers of synchronous excitation respectively.

Fig. 25 is a (left) nine-layer synchronous excitation pulse waveform according to an embodiment of the present invention, the solid line represents the real part of the waveform, and the broken line represents the imaginary part of the waveform; nine layers of synchronous excitation profiles are (right), and different types of curves in the figure correspond to different layers of synchronous excitation respectively.

Fig. 26 shows a linear phase SLR refocusing rf pulse waveform (left) and excitation profile (right) according to an embodiment of the invention. Waveform characteristic parameters: the time-band product TBP is 16, the in-band ripple coefficient is 0.1%, the out-of-band ripple coefficient is 0.1%, the pulse flip angle is pi, and the gradient strength of the selected layer is 0.5G/cm.

FIG. 27 is an intraoperative navigator sequence (Local fast spin echo sequence, local-HASTE-NV) of an embodiment of the present invention: the first 90 ° rf excitation pulse and the first 180 ° refocusing pulse are respectively SLR pulses shown in fig. 14 (left) and fig. 15 (left), and the second and subsequent 180 ° refocusing pulses are respectively SLR pulses shown in fig. 15 (left) or fig. 26 (left). 90 ° rf excitation pulse and 180 ° refocusing pulseThe time interval between the punches is TE/2 and the time interval between 180 deg. refocusing pulses is TE. G _s1 、G _s2 、G _s3 …G _sn Representing the gradient of layers selected in different directions, G _pre Representing a preliminary read gradient, G _r1 、G _r2 、G _r3 …G _rn Representing read gradient, G _sc1 、G _sc2 、G _sc3 …G _scn Represents the compensation gradient applied in the selected layer direction, ±g _P1 、±G _P2 、±G _p3 …±G _pn Representing the phase encoding gradient of positive and negative polarity. Partial repeat execution N within the dashed box _PE Signal acquisition starts after the third SLR pulse/4 times.

Detailed Description

In order to make the technical problems solved by the invention, the technical scheme adopted and the technical effects achieved clearer, the invention is further described in detail below with reference to the accompanying drawings and the embodiments. It is to be understood that the specific embodiments described herein are merely illustrative of the invention and are not limiting thereof. It should be further noted that, for convenience of description, only some, but not all of the matters related to the present invention are shown in the accompanying drawings.

1. Principle and structure

The basic framework of the navigation system and method for minimally invasive surgery of the invention is shown in fig. 1, and technical features different from the conventional diagnostic MRI system and other MRI navigation systems are described as follows:

(1) The magnet is designed into an open U-shaped structure shown in fig. 1, and the field intensity of the magnet is in the range of 0.2T-0.7T, preferably 0.3T, so as to avoid susceptibility artifacts or metal artifacts of high field intensity aggravated images. An important technical indicator, unlike conventional magnet designs, is that the magnet gap is greater than 500mm, and the plate diameter is reduced by 1/3 to 1/2, preferably 360mm, over conventional diagnostic designs, in order to increase the convenience of the surgical procedure. (2) The rf transceiver coil is designed as a biplane circular polarization type as shown in fig. 2. Unlike the planar RF coil for diagnosis and scanning, the T/R switch realizes the emitting and receiving functions simultaneously, and this can avoid the spatial limitation of independent receiving coil to minimally invasive operation, especially the capacity number of at least 100, and improve the homogeneity of RF field emission and signal reception. (3) The gradient coil is designed as a double-plane main coil, and a double-plane axial shielding coil is additionally arranged on the outer side of the gradient coil, the winding mode is shown in fig. 3, the installation position is shown in fig. 1, the diameters of each coil and the shielding plate are preferably 360mm, an eddy current shielding plate with high magnetic permeability and high resistivity is additionally arranged between the main coil and the axial shielding coil, eddy current interference generated during gradient switching is sufficiently weakened, and the thickness increase of the gradient coil is limited.

The technical characteristics mainly ensure the system opening degree, safety and convenience required by the minimally invasive surgery; in particular, in order to improve the quality and real-time performance of magnetic resonance imaging and ensure the accurate positioning of surgical instruments and the accurate control of surgical routes, the invention provides the following intra-operative navigation technical scheme and imaging method:

firstly, susceptibility artifacts or metal artifacts of the nonmagnetic medical instrument are detected and navigation sequence parameters (such as radio frequency pulse width and echo acquisition time) are optimized in advance according to the operation flow shown in fig. 7, a three-dimensional high-resolution T1 weighted image is acquired on a navigation imager or a high-field imager before an operation, scan enhancement, angiography or functional imaging are added if necessary, and then operation route planning is performed, including target point setting, needle insertion path planning and operation scheme making.

Secondly, the scanning part is placed in an isocenter area through the calibration of the horizontal plane height of a scanning bed and the guidance of a laser marking line of a laser positioning instrument on a navigation imaging system shown in fig. 1, and a dynamic local shimming technology is selected for shimming.

Then, an internal reference coordinate system is established through a navigation module shown in fig. 10 and a navigation interface shown in fig. 11 according to the navigation flow shown in fig. 9, and the local rapid navigation sequence is controlled to run, fault signals in three orthogonal directions are repeatedly collected according to the operation route, and each direction can collect a single layer or multiple layers simultaneously, or the positioning parameters and the waveform parameters are continuously updated and scanned in real time in the operation process by adopting the real-time path tracking mode shown in fig. 12, so that the position of the surgical instrument is continuously tracked and positioned.

The local rapid navigation sequence design mode is shown in fig. 13-27, and is basically characterized in that radio frequency pulses have three-dimensional space selectivity, layer selection gradients corresponding to a first radio frequency excitation pulse, a first radio frequency refocusing pulse and a second radio frequency refocusing pulse are arranged in different directions, an SLR pulse waveform with highly optimized excitation profile is adopted, a plurality of adjacent frequency bands can be selected to be excited uniformly at the same time, the range and the interval of each frequency band can be accurately adjusted on a positioning image through layer thickness and layer interval, the pulse excitation profile amplitude can be corrected and consistent according to the integral area of each frequency band, a part of k-space lines can be repeatedly acquired for a plurality of times, pulse phase circulation can be carried out, a frequency encoding gradient and a layer selection gradient can be applied simultaneously during data acquisition, the bandwidth of a receiver is set to be 100kHz or more, and k-space data accumulation and partial Fourier image reconstruction can be carried out. The navigation imaging method not only can improve the uniformity, resolution and signal-to-noise ratio of the image, but also can inhibit susceptibility artifacts and metal artifacts and shorten the scanning time.

After the navigator treatment is finished, local high resolution scanning and enhanced scanning are performed on a navigator imager or a high field imager to confirm the curative effect. The acquisition matrix of the frequency and phase encoding directions is doubled here to further improve the image resolution compared to intra-operative navigator scans.

As shown in fig. 1, the present magnet structure is used in a minimally invasive surgical navigation system comprising: yoke 1, magnet steel 2, polar plate 3, shielding coil 4, gradient coil 5, radio frequency coil 6, scanning bed 7, heating rod 8, heating plate 9, temperature sensor 10, shell 11, T/R send-receive switch 12, gradient power amplifier 13, radio frequency power amplifier 14, preamplifier 15, temperature control unit 16, optical tracking positioning system 17, coil tuning control unit 18, scanning bed control unit 19, navigation imaging control unit 20, system control interface 21, host 22, display 23.

The magnet 60 is designed as an open U-shaped structure as shown in FIGS. 1 and 2, and the field strength of the magnet is in the range of 0.2T to 0.7T, preferably 0.3T, so as to avoid susceptibility artifacts or metal artifacts of the high field strength emphasized image. An important technical index, unlike conventional magnet designs, is that the magnet gap is greater than 450mm, and the plate 3 radius is reduced by 1/5 to 1/4 as compared to conventional diagnostic type, so as to increase the convenience of the surgical operation. The rf coil 6 is designed as a double-plane circular polarization type, a plurality of nonmagnetic capacitors 51 are arranged between the outer conductor ring 50 and the inner conductor ring 52, as shown in fig. 3, unlike a flat rf coil for diagnostic scanning, the transmission and receiving functions are realized simultaneously by the T/R switch 12, the space limitation of the independent receiving coils to the minimally invasive surgery is avoided, particularly, the number of nonmagnetic capacitors 51 is at least 100, the uniformity of rf field transmission and signal reception is fully improved, and the matched rf power amplifier has 15kW or higher power to meet the requirement of an ultra-fast navigation imaging sequence.

As shown in fig. 4, the biplane gradient coil 5 is designed as three sets of biplane main gradient coils 101, each set of two sets of biplane main gradient coils are respectively located at the upper and lower positions of an imaging area and near to the polar plate 3, linear gradient fields are respectively generated in the directions of an X axis, a Y axis and a Z axis, a biplane axial shielding coil 102 is additionally arranged at the outer side of the axial main gradient coil 101, a winding form of the axial biplane main gradient coil 101 and the axial shielding coil 102 is calculated by a target field method, the coils are in a concentric circle form as shown in fig. 5 and are manufactured by adopting a printed circuit board technology, the wiring mode of the axial shielding coil 102 is similar to that of a gradient cable as shown in fig. 4, but the directions of gradient currents are opposite, and the installation positions of the main gradient coils are shown in fig. 1 and fig. 2; in addition, an insulating layer 103 is added between the biplane axial main gradient coil and the axial shielding coil and is tightly connected, an eddy current shielding plate with high magnetic permeability and high resistivity is added between the main gradient coil 101 and the polar plate 3, so that eddy current and thrust generated during gradient switching are sufficiently weakened, and the thickness increase of the gradient coil is limited.

Other parameters are as follows: the field intensity of the magnet 60 is preferably 0.3T, the air gap is preferably 500mm, the radius of the magnet polar plate 3 is smaller than 380mm, a laser positioning instrument 1 is arranged above the magnet 60, 12V direct current power supply is adopted, the shape of a linear light spot is adopted, the output wavelength is in the range of 635nm to 650nm, and the position coordinate is accurate to 0.1mm. The maximum gradient strength of the gradient coil is 15mT/m, the switching rate is 50mT/m/ms or higher, the radius of the main gradient coil 101 and the axial shielding coil 102 is preferably 350mm, the thickness is 4mm to 5mm, the distance is 2mm to 3mm, the gradient linearity is usually limited to be within 5%, the DSV of the effective imaging area is not less than 250mm multiplied by 300mm, and the size of the gradient linear area is calculated according to the Biot-Savart theorem for design confirmation.

The whole set of device forms an informationized operating room suitable for MRI image navigation, as shown in FIG. 6, comprising: magnet 60, laser positioning instrument 61, scanning bed 7, bayonet 63, very large liquid crystal display 64, keyboard 66, cabinet 67 (with host and magnetic field alarm inside), operating bed 68, auxiliary positioning marking 69. The console adopts a mobile small computer desk, lowers the host computer and is provided with a magnetic field alarm for preventing the computer desk from being interfered by a strong magnetic field when entering a 10-Gaussian range. The imaging data adopts an optical fiber transmission mode and a double-screen display mode, a small liquid crystal display 65 (medical touch screen display) is fixed on the desktop of a computer desk, and an ultra-large liquid crystal display 64 is arranged on the back wall of an imager. The operation table 68 is in butt joint with the scanning table 7 through the bayonet 63, so that a patient can be pushed into an imaging area to perform an operation while scanning, and can also be moved out of the imaging area to perform certain special operation according to operation requirements. During clinical scanning and operation, the main machine loads the navigation sequence to the navigation imaging control unit, the latter controls the radio frequency transmitting coil to generate the required radio frequency pulse according to the navigation sequence and the navigation parameter, and controls the gradient coil to generate the required gradient current pulse, the magnetic resonance signal of three-dimensional space coding is generated at the imaging part of the human body, then the magnetic resonance signal is received by the radio frequency receiving coil, and then uploaded to the main machine through the navigation imaging control unit, and the navigation control software carries out data processing and image display in real time, so that the doctor can obtain the position information of the surgical instrument and the target in time.

The invention provides a navigation system and a navigation method for minimally invasive surgery, which are a diagnosis and interventional therapy dual-purpose magnetic resonance navigation system and a technical scheme with a highly open structure and navigation calibration and real-time high-resolution scanning functions, and particularly provide a navigation-dedicated local imaging technology and a precisely-positioned multilayer synchronous excitation technology.

The highly open system structure comprises a U-shaped magnet with an opening of more than 500mm, and is provided with a transmitting-receiving type double-plane radio frequency coil and a special shielding gradient coil for improving the convenience of operation.

The transmitting and receiving type double-plane radio frequency coil simultaneously realizes transmitting and receiving functions through the T/R switch, avoids the space limitation of an independent receiving coil on minimally invasive surgery, particularly has the capacity number of at least more than 100, and fully improves the uniformity of radio frequency field transmission and signal reception.

The special shielding gradient coil is an X/Y/Z axis double-plane gradient coil, the diameter of the special shielding gradient coil is reduced by 1/3 to 1/2, preferably 360mmm, the gradient linearity is limited to be within 10%, and a double-plane shielding gradient is additionally arranged outside the Z axis main gradient coil, and the winding form is shown in figure 5.

The implementation mode of the navigation calibration function comprises dynamic local shimming, wherein a scanning part is placed in an isocenter area through the calibration of the plane height of a scanning bed and the guidance of a laser marking line of a laser positioning instrument, and a dynamic local shimming technology is selected for shimming.

The implementation mode of the navigation calibration function comprises metal artifact test and calibration, wherein a magnetic field distribution diagram is tested firstly based on a standard water model of a copper sulfate aqueous solution and a gradient echo sequence, then a strip-shaped or cross-shaped nonmagnetic metal material is added in the central area of the standard water model, the magnetic field distribution diagram is tested through the gradient echo sequence, the difference between the two magnetic field distribution diagrams is calculated, the nonmagnetic metal material or a composite material with the smallest difference is selected for manufacturing a surgical instrument, and finally the size of a non-signal area is measured on a gradient echo image and consistency verification is carried out on the size of the non-signal area and the actual size of the metal strip, so that the non-signal area is used as an accurate mark of the position of the surgical instrument.

The implementation mode of the navigation calibration function is that an internal reference coordinate system is established through a navigation module shown in fig. 10 and a navigation interface shown in fig. 11 according to a navigation flow shown in fig. 9, operation route coordinates are preset, local rapid navigation sequence operation is controlled, imaging area signals of three orthogonal planes are gradually acquired along an operation route, or a real-time path tracking mode shown in fig. 12 is adopted, the central position of a scanning layer is sequentially increased by delta L/2 from the position of an operation entrance until the position of a target point along the direction of the operation path shown by an arrow, and positioning parameters and waveform parameters are continuously updated in the operation process and scanned in real time so as to continuously track and position the surgical instrument.

The local imaging technique is characterized by that it adopts multi-azimuth selective layer gradient and radio-frequency pulse whose excitation profile is highly optimized to excite proton signal in a small volume, and the first 90 deg. radio-frequency excitation pulse and first 180 deg. refocusing pulse adopt uniformly excited SLR pulse waveform, and the correspondent selective layer gradient amplitude of first radio-frequency pulse is set in navigation sequence as G _s1 ＝2πΔf/γ/FOV _x Setting the gradient amplitude of the layer selection corresponding to the second radio frequency pulse as G _s2 ＝2πΔ/fγ/FOV _y Setting the gradient amplitude of the layer selection corresponding to the third radio frequency pulse as G _s3 =2pi Δf/γ/THK, where FOV _x 、FOV _y And THK can be directly regulated on the sequence parameter table.

The essential feature of the local imaging technique is that the bandwidth Δf of the radio frequency pulses needs to be optimized in advance in the case of metal artifact interference, the receiver bandwidth is preferably above 100kHz, the excitation profile of the SLR pulses is optimized in the case of TBP + 8, the waveform characteristics of which are shown in fig. 14 and 15, and the excitation profile or calibration image amplitude is optimized in the case of simultaneous multi-layer excitation based on the integrated area of each frequency band obtained by the calibration sequence shown in fig. 22, then gradient pulses are applied simultaneously in the frequency encoding direction and the slice selection direction during k-space data acquisition.

The bandwidth of the radio frequency pulses af is optimized by setting af to different values, e.g. 750hz,1kHz,1.25kHz,1.5kHz, etc., from which af with minimal susceptibility or metal artifacts is searched, according to the operational flow shown in fig. 7, the af optimum value is close to the single echo acquisition time length when both the frequency encoding gradient and the layer selection gradient are applied during data acquisition, otherwise af is preferably 2kHz or higher.

The fundamental feature of the Local imaging technique is that T1 weighted imaging can be quickly achieved by selecting either Local-SE-NV as shown in FIG. 13 or Local-ME-NV as shown in FIG. 13 and minimizing the echo time. Here, a layer gradient G _s1 、G _S2 And G _S3 Is arranged in an orthogonal direction on the positioning image, wherein the first two are perpendicular (or parallel) to the surgical path, the other are parallel (or perpendicular) to the surgical path, and the other gradient of the selected layer has the orientation and amplitude of G _s3 Similarly, the signal acquisition starts after the third radio frequency pulse, and the k-space filling and image reconstruction take part in fourier fashion.

The Local imaging sequence Local-ME-NV is characterized by applying a G at the moment of switching the positive and negative polarities of the read gradient when the scan speed is prioritized _blip The gradient, as shown in fig. 17, thus the data acquisition part is equivalent to echo plane acquisition, the number of echoes in the dashed box can be set to 32 or other smaller integers, and the acquired data is processed according to the phase correction and image reconstruction modes commonly used in echo plane imaging.

The Local imaging technology is characterized in that T2 weighted imaging can be rapidly realized by selecting the Local-HASTE-NV sequence shown in figure 27. Here, the echo time is set in the range of 80ms to 140ms, and the slice gradient G is selected _s1 、G _S2 And G _S3 Is arranged in an orthogonal direction on the positioning image, wherein the first two are perpendicular (or parallel) to the surgical path, the other are parallel (or perpendicular) to the surgical path, and the other gradient is oriented with G _s3 The same; and, here, additionally applying a compensation gradient G in the layer-selecting direction _sc1 、G _sc2 、G _sc3 …G _scn The gradient amplitude is equal to G _S3 The gradient width is the same as G _r1 The same; the first pulse and the second pulse take waveforms similar to those shown in fig. 14 and 15, the third pulse and the subsequent pulse take waveforms similar to those shown in fig. 15 or 26, the waveform parameters are optimized as described above, and the signal is acquired from the third pulseThe radio frequency pulse is followed by a partial fourier mode for k-space filling and image reconstruction.

The basic characteristic of the precisely positionable multilayer synchronous excitation technology is that the radio frequency pulse has spatial selectivity, and can uniformly excite a plurality of adjacent frequency bands simultaneously, for example, for the three-layer synchronous excitation case, the navigation sequence Local-MSME-NV shown in FIG. 18 is selected, wherein the waveform of the radio frequency excitation pulse is calculated according to the following formula:

the waveform characteristics and excitation profile are shown in FIG. 20 or FIG. 23, the refocusing pulse is preferably 180 DEG minimum phase SLR pulse or linear phase SLR pulse in the spin echo acquisition mode, as shown in FIGS. 15 and 26, and the RF bandwidth requirement of the refocusing pulse is not less than (3. THK+2. Delta. S). Gamma. G _s /(2 pi). Likewise, other multi-layer simultaneous excitation scenarios may be achieved in a similar manner, as shown in fig. 24 and 25.

The essential feature of the navigation sequence Local-MSME-NV is that the frequency band range and spacing of the synchronous excitation pulses can be precisely adjusted on the scout image by layer thickness and layer spacing, in three scansAnd->Are respectively arranged asAnd-> Each time scanning acquires a part of k-space lines, and three times of scanning obtain an image S ₁ ，S ₂ And S is ₃ Finally, image decomposition is carried out according to the following formula to obtain eachLayer image:

the basic feature of the navigation sequence Local-MSME-NV is that when the scanning speed is prioritized, a G is applied at the moment of switching the positive and negative polarities of the read gradient _blip The gradient, as shown in fig. 19, is such that the data acquisition portion is equivalent to echo plane acquisition, the number of echoes in the dashed box can be set to 32 or other smaller integers, the acquired data is processed according to the phase correction and image reconstruction modes commonly used in echo plane imaging, and then image decomposition is performed according to the formula (4) to obtain multiple layers of images.

In the case of incomplete image decomposition (e.g., TBP > 8), the multiple closely adjacent bands are uniformly excited in a calibrated manner

(1) The excitation profile is obtained by using the synchronous scanning calibration sequence shown in FIG. 22 and selecting the SLR pulse waveform acquisition signal of each phase code in equation (3) and Fourier transforming to the frequency domain, respectively denoted as C ₁ 、C ₂ And C ₃ ；

(2) Calculation C ₁ 、C ₂ And C ₃ Is denoted by k respectively ₁ ＝∫∫C ₁ dxdy、k ₂ ＝∫∫C ₂ dxdy and k ₃ ＝∫∫C ₃ dxdy, where ≡ζ represents two-dimensional integration, x and y represent two-dimensional coordinates;

(3) Calibration profile C ₁ 、And->As shown in fig. 23.

(4) Also, the two-dimensional image obtained by three scans is calculated according to the integral area ratio k ₁ 、k ₂ And k ₃ Calibration is performed and image decomposition of each layer is performed according to equation (4).

The essential feature of the precisely positioned multilayer synchronous excitation technique is that the multilayer synchronous excitation method can be used for layer imaging at any position, namely

And corresponding positioning lines are set on the positioning image through the user interface shown in FIG. 11 And transmits the azimuth parameters of each layer to the pulse waveform calculation program of the navigation module shown in fig. 10, outputs updated pulse waveforms, loads the updated pulse waveforms to the navigation imaging control unit shown in fig. 1, and then runs a scanning sequence and performs real-time image reconstruction.

2. Examples of the invention

The navigation method and the software of the invention are suitable for various field intensity magnetic resonance imaging systems, in particular for an open magnetic resonance imaging system, such as the magnetic resonance imaging navigation system shown in fig. 1 and 2, and the basic framework and technical implementation details are as follows:

The field intensity of the magnet is preferably 0.3T, the air gap is preferably 550mm, the diameter of a magnet shimming ring (or called a polar ring) is smaller than 360mm, a laser positioning instrument 1 is arranged above the magnet, 12V direct current power supply is adopted, the shape of a linear light spot is adopted, the output wavelength is in the range of 635nm to 650nm, and the position coordinate is accurate to 0.1mm. The radio frequency receiving and transmitting coil adopts a double-plane circular polarization type shown in fig. 3, wherein the number of the capacitors is at least 100 or more, the transmitting and receiving functions are realized simultaneously through a T/R switch, and the matched radio frequency power amplifier has 15kW or higher power. The gradient coil adopts a flat plate type main coil, the maximum gradient strength is 20mT/m, the switching speed is 60mT/m/ms, the coil diameter is preferably 360mm, the gradient linearity is generally limited to be within 10%, and the size of a gradient linear region is calculated according to the Biot-Savart theorem for design confirmation; in addition, a shielding coil is added to the outside of the axial gradient coil, and the winding forms of the axial gradient coil and the shielding coil are calculated by a target field method, as shown in fig. 5. The console adopts a mobile small computer desk, lowers the host computer and is provided with a magnetic field alarm for preventing the computer desk from being interfered by a strong magnetic field when entering a 10-Gaussian range. The imaging data adopts an optical fiber transmission mode and a double-screen display mode, the medical touch screen display is fixed on the desktop of a computer desk, and the oversized liquid crystal display is arranged on the back wall of the imager. The operation bed is in butt joint with the scanning bed through the bayonet, so that a patient can push into the imaging area to perform an operation while scanning, and can also move out of the imaging area to perform certain special operation according to the operation requirement. Thus, the kit constitutes an informationized operating room suitable for MRI image navigation, as shown in FIG. 6. In order to realize that image guidance in operation is not seriously interfered by magnetic susceptibility artifact or metal artifact of a surgical instrument, the invention detects and calibrates the metal artifact of the surgical instrument according to a detection flow shown in fig. 7, wherein a reference water model is a standard water model containing a copper sulfate water solution used for daily quality inspection, a target water model is a water model made of a strip-shaped or derrick-shaped nonmagnetic metal material is added in the central area of the standard water model, firstly, a magnetic field distribution diagram is tested based on the reference water model and a gradient echo sequence, then, the magnetic field distribution diagram is tested based on the target water model and the gradient echo sequence, the difference between the two magnetic field distribution diagrams is calculated, the surgical instrument is manufactured by using nonmagnetic metal materials (such as titanium alloy and ceramic composite materials) with the smallest difference value, and parameters such as optimal pulse width, echo acquisition time and the like are obtained, and finally, the size of a dead zone corresponding to the metal strip is measured on an image and consistency verification is carried out with the actual size of the metal strip.

On the other hand, various navigation methods and imaging methods of the magnetic resonance imaging navigation system are implemented by navigation software including 1031, 1032, and 1033 function modules shown in fig. 10 for controlling the operation of the hardware systems 100, 101, and 102. The control functions of the module 1031 include: performing image measurement and analysis; an assessment report is generated. The control functions of module 1032 include: establishing a magnetic resonance navigation scanning sequence, a protocol and a parameter table; selecting a navigation scheme and mode; and carrying out system calibration, scanning and image reconstruction in real time. The control functions of module 1033 include: establishing an image domain internal reference coordinate system; defining and visualizing a surgical route; displaying target points and instrument tracks in real time by multiple planes; the surgical route deviation is analyzed in real time and warned. The operation interface design mode of the navigation software is shown in fig. 11, in which S11 to S31 represent images of different orientations or layers, and the images are switched between a single Zhang Xianshi mode and a plurality of display modes by double mouse click. The system calibration sequence comprises local shimming and gradient nonlinear calibration, is used for improving magnetic field uniformity and compensating gradient nonlinear errors, the T1 navigation sequence refers to a T1 weighted imaging rapid navigation sequence and comprises sequences of global and local SE-NV, ME-NV, MSME-NV and the like, the T2 navigation sequence refers to a T2 weighted imaging rapid navigation sequence and comprises sequences of global and local HASTE-NV, FISP-NV and the like, and the T2 navigation sequence refers to a T2 weighted imaging rapid navigation sequence and comprises sequences of global and local EPI-NV and the like. The sequence parameter table and the reconstruction parameter table are in the form of floating windows, each sequence name corresponds to a respective parameter table page, and parameter values can be changed and stored, as shown in fig. 11 (below). In the figure, a start key updates a pulse waveform according to the selected layer gradient azimuth parameters in the parameter table after the mouse is clicked, and scanning and real-time image reconstruction are performed.

The MRI navigator system implements clinical navigator operation according to the navigator flow shown in fig. 9, and is specifically described as follows:

first, a three-dimensional isotropic high-resolution T1 weighted image is obtained on a high-field imager, scanned for enhanced scanning for craniocerebral surgical navigation, and scan diffusion tensor imaging and angiography as necessary to display lesions, blood vessels, and nerve fibers. Then, surgical route planning is performed based on the internal reference coordinate system. For craniocerebral operation navigation, the internal reference coordinate system is established by setting the connecting line from the midpoint of the anterior commissure trailing edge to the midpoint of the posterior commissure leading edge as the commissure middle diameter, setting the horizontal plane made by the connecting line as the HO plane, setting the coronal plane of the commissure middle diameter as the FO plane, and adding the mid-sagittal plane SO plane of the brain to form three reference planes of the positioning image. The intersection of these three reference planes is defined as the origin. The axis passing through the front and back direction of the origin is a sagittal axis (coincident with the commissure diameter) and is defined as a Y axis; the vertical axis perpendicular to the Y axis through the up-down direction of the origin is defined as the Z axis; the X-axis is defined as the coronal axis passing through the left-right direction of the origin and intersecting the Y-axis perpendicularly. By using the planes and the axes, three-dimensional space coordinates of each structure in the brain can be drawn. The route specification mode is that multi-plane image reconstruction is carried out on three-dimensional full-brain high-resolution T1 weighted k space data, an optimal operation route is determined on an image according to an internal reference coordinate system and marked as a series of node positions, corresponding gradient azimuth parameters are read, then the gradient azimuth parameters are corrected according to the ratio of the maximum gradient intensity of each of a high-field imager and a navigation imager, and the gradient azimuth parameters are stored in each navigation sequence parameter table of the navigation imager to be used as navigation azimuth parameter default values.

The navigator software shown in fig. 10 and 11 is then run on the MRI navigator imager, the navigator sequences are selected and cycled, with specific design and feature parameter set requirements for each navigator sequence as described in examples 1-6 below. For each navigation sequence, a variable angle scanning mode is adopted to read navigation azimuth parameters and scan images, and simultaneously, images in other two orthogonal directions are additionally scanned, three groups of tomographic images are obtained each time, each group can be a single image or a plurality of images acquired by synchronous excitation, and the anatomical structure and the medical instrument position of an imaging area are displayed on a medical display. Or, under the condition that the nonlinear error of the gradient system is large, the real-time path tracking mode shown in fig. 12 is adopted to realize real-time continuous tracking of the probe position, the distance between adjacent layers is delta L in fig. 12, the center position of the scanning layer is sequentially increased by delta L/2 along the direction of the surgical path shown by an arrow, the real-time scanning starts from the surgical entrance position to the target position, the current scanning can be paused during the real-time scanning, the positioning line of the next node position can be reset on the scanned image by referring to the preset surgical route and the internal reference coordinate system, and then the image is scanned or the pulse waveform is updated according to the positioning parameters and synchronous scanning is started. Since the surgical instrument (such as a non-magnetic metal probe) is displayed as a non-signal characteristic shape on the MRI image, the contrast with surrounding tissues is provided, the position of the surgical instrument can be accurately marked, the sequence parameters and the magnetic susceptibility artifact can be optimized through the operation flow shown in fig. 7, so that the contrast between the surgical instrument and the surrounding tissues is increased, and the positioning accuracy of the position of the surgical instrument is not influenced by external factors due to the navigation mode; in contrast, the optical tracking and positioning system in fig. 1 (including an infrared navigation camera, a positioning tracer, a puncture needle equipped with a navigation light ball, a magnetic compatible power supply, a communication cable, a navigation function module, and the like) is used as a conventional alternative for tracking azimuth information of a surgical instrument in real time, and needs to establish a relative relation between an MRI scanning system coordinate system and an optical positioning system coordinate system, and then convert the positioning system coordinate of the surgical instrument into a scanning system coordinate and display the scanning system coordinate and the MRI image of the patient together on a screen in real time, but positioning accuracy is easily affected by body position variation or tissue organ displacement in the operation of the patient and bending deformation of the tip of the instrument.

Finally, after the navigator treatment is finished, local high-resolution scanning and enhanced scanning are performed on the navigator imager or the high-field imager to confirm the curative effect. The acquisition matrix of the frequency and phase encoding directions is doubled here to further improve the image resolution compared to intra-operative navigator scans.

Example 1 fast local spin echo scanning

Based on the MRI navigator system and the navigator flow, the Local spin echo sequence Local-SE-NV shown in fig. 13 is selected as the fast navigator sequence, wherein the 90 ° rf pulse is preferably the minimum phase SLR pulse shown in fig. 14, and the two 180 ° refocusing rf pulses are preferably the minimum phase SLR pulse shown in fig. 15 or the linear phase SLR pulse shown in fig. 26. When the radio frequency power is insufficient, the radio frequency bandwidth can be reduced, the time-zone product TBP is kept unchanged, the SLR pulse width is correspondingly adjusted, the data point interval time of the waveform file is adjusted according to the required SLR pulse width, and meanwhile, the gradient amplitude in the sequence parameter table is adjusted until the required layer thickness is obtained. In the case of metal artifact interference, the receiver bandwidth is preferably 100kHz or higher, the echo time is optimized to a minimum, and the frequency encoding gradient and amplitude G are applied simultaneously during data acquisition _s3 And the radio frequency pulse bandwidth deltaf is set to different values, e.g., 750hz,1khz,1.25khz,1.5khz, etc., from which the deltaf at which susceptibility or metal artifacts are minimal is searched for according to the operational flow shown in fig. 7. When the slice-selective gradients are applied simultaneously during data acquisition, the optimal value of TBP/af is close to the single echo acquisition time length, otherwise af is typically preferably a value of 2kHz or higher. In real-time path tracking mode, positioning lines of navigation sequence are arranged on a multi-plane (such as tri-plane) positioning image, the thickness, spacing and azimuth of the positioning lines are adjusted, and the gradient strength corresponding to the first radio frequency pulse is set as G in the imaging sequence _s1 ＝2πΔf/γ/FOV _x The gradient strength corresponding to the second radio frequency pulse is G _s2 ＝2πΔf/γ/FOV _y The third RF pulse has a gradient strength of G _s3 =2pi Δf/γ/THK, where Δf is the rf pulse bandwidth, FOV _x And FOV (field of view) _y The fields of view representing the x-axis and y-axis directions, respectively, can be set to 32mm, 64mm or 128mm, the layer thickness THK is set to 8mm, and the sequence repetition time tr=300 ms, the number of phase encoding steps N are set _pe =32, 64, or 128, the number of sequence repetitions nex=1. The k-space data is acquired by partial fourier acquisition, and other adjacent layers are selectively excited in the same way in the remaining time of each TR, and the scanning layers are set to be perpendicular or parallel to the direction of the operation route in real time according to the navigation mode. Then, the navigation sequence is operated, a group of k-space lines are acquired each time of phase coding circulation, interpolation or zero filling is carried out after data acquisition, and partial Fourier reconstruction is carried out to obtain a T1 weighted image. The scan time is calculated as

T _ACQ ＝55％·NEX·N _PE ·TR＝0.3·55％·1·32＝5.3(s)

Since TR is not less than 300ms, the proton magnetization vector of human tissue can be restored to a large extent, thus avoiding interlayer overlapping artifacts. Compared with a spin echo Sequence (SE) for conventional diagnosis, the scanning method can remarkably improve the scanning efficiency, and the image has high resolution. On a high field intensity imaging system, the combination of the rapid navigation sequence and the parallel acquisition technology of the invention can further shorten the scanning time by several times.

Example 2 fast local Multi-echo scanning

Based on the MRI navigator system and navigator flow described above, the Local multi-gradient echo sequence Local-ME-NV shown in fig. 16 is selected for the fast navigator sequence, wherein the 90 ° rf pulse is preferably the minimum phase SLR pulse shown in fig. 14, and the first and second 180 ° refocusing rf pulses are selected from the minimum phase SLR pulse shown in fig. 15 or the linear phase SLR pulse shown in fig. 26. When the radio frequency power is insufficient, the radio frequency bandwidth can be reduced, the time-zone product TBP is kept unchanged, the SLR pulse width is correspondingly adjusted, the data point interval time of the waveform file is adjusted according to the required SLR pulse width, and meanwhile, the gradient amplitude in the sequence parameter table is adjusted to the required layer thickness. In the case of metal artifact interference, the receiver bandwidth is preferably 100kHz or higher, the echo time is optimized to a minimum, and the frequency encoding gradient and amplitude G are applied simultaneously during data acquisition _s3 And optimizes deltaf according to the operational flow shown in figure 7. In real-time path tracking mode, positioning lines of a navigation sequence are arranged on a multi-plane (such as three-plane) positioning image, the thickness, the spacing and the azimuth of the positioning lines are adjusted, and the gradient strength corresponding to the first radio frequency pulse is set as G in the imaging sequence _s1 ＝2πΔf/γ/FOV _x The gradient strength corresponding to the second radio frequency pulse is G _s2 ＝2πΔf/γ/FOV _y The third RF pulse has a gradient strength of G _s3 =2pi Δf/γ/THK, set FOV _x And FOV (field of view) _y 32mm, 64mm or 128mm, THK 8mm, and setting the sequence repetition time TR=300 ms, the number of phase encoding steps N _pe =32, 64, or 128, the number of sequence repetitions nex=1 or 2. The k-space data acquisition adopts a partial Fourier acquisition mode, other adjacent layers are selectively excited in the same mode in the residual idle time of each TR, and the scanning layer is set to be perpendicular or parallel to the direction of the operation route in real time according to the navigation mode. When NEX=2, the polarity of each reading gradient is alternately reversed every time of scanning, the phase encoding gradient is sequentially increased or decreased every time of phase encoding circulation, k-space data are acquired, complex signals are accumulated after the data acquisition is finished to enhance the signal-to-noise ratio and eliminate the phase error, and then a partial Fourier reconstruction mode is adopted to obtain a T1 weighted image. Scan time T _ACQ Can be countedCalculated as

T _ACQ ＝55％·NEX·N _PE ·TR＝0.3·55％·2·32＝10.6(s)

When nex=1, scan time T _ACQ Can be calculated as

T _ACQ ＝55％·NEX·N _PE ·TR＝0.3·55％·1·32＝5.3(s)

The imaging method greatly improves the scanning efficiency based on a local scanning mode under the condition of ensuring the high resolution and the signal-to-noise ratio of the focus area, and can inhibit metal artifacts. The number of echoes in the dashed box of fig. 16 can be further increased at a relatively fast gradient switching rate (e.g. 100 mT/m/ms) to further improve the image signal-to-noise ratio while ensuring the required contrast.

In addition, when the scanning speed is particularly prioritized, G can be selected in the sequence parameter table _blip Gradient option, applying a G at the moment of switching positive and negative polarities of reading gradient _blip Gradient, as shown in fig. 17. Number of echoes N in the dashed line box in the figure _pe Can be set to 32 or 64, and the intensity of the phase encoding gradient satisfies G _pe ＝G _blip ·N _pe Condition/2, scan time is

T _ACQ ＝55％·NEX·N _PE ·TR＝0.3·55％·2＝0.33(s)

In this way, real-time scanning is fully realized, and due to the adoption of a short echo chain and the simultaneous application of frequency coding gradients and layer selection gradients during data acquisition, common nyquist artifacts and susceptibility artifacts can be significantly reduced, and the interference of image artifacts can be sufficiently eliminated on the navigation system shown in fig. 1 by combining the pre-scanning and phase correction schemes commonly used for echo planar imaging.

Example 3 fast local three-layer synchronous scanning

Based on the MRI navigation system and the navigation flow, the rapid navigation sequence selects the multilayer synchronous excitation Local-MSME-NV sequence shown in figure 18. To increase the rf excitation uniformity, the waveform of the first rf pulse of the imaging sequence employs the linear phase SLR pulse waveform shown in fig. 20; in order to excite multiple layers of the human body simultaneously, the radio frequency pulse waveform is phase coded as follows:

here, SLR (t) represents a linear phase SLR pulse, G _s Indicating the direction of the selected layer asGradient intensity of>Is a vector representing the spatial position of the center layer, which can be set directly on the positioning image, here as the isocenter, the position of the adjacent layer is set by the layer thickness THK in the selected layer direction and the layer spacing DeltaS, deltaS being defined by G _S And bandwidth Δf is defined as Δs=2pi·Δf/γ/G _s The setting is carried out such that,and->Is the initial phase, and DeltaS>THK. In order to achieve uniform excitation of each layer, avoid interlayer overlapping artifacts, and enable accurate positioning in clinical navigation applications, where the bandwidth of the first rf pulse is set to 1kHz at single layer excitation, the time-band product TBP is set to 4, the in-band and out-of-band ripple coefficients are set to 0.5% and 0.1%, respectively, the selected layer gradient is 0.7 gauss/cm (G/cm), then pulse waveforms for simultaneous uniform excitation of multiple adjacent layers are obtained according to equation (3), each phase cycle produces a waveform, and the pulse waveforms and excitation profiles for three layer simultaneous excitation corresponding to one of the phase cycles are shown in fig. 20. The second and third RF pulses of the Local-MSME-NV sequence may employ Hamming windowed 180 deg. sinc pulses, preferably here 180 deg. minimum phase SLR pulses (see FIG. 15) or linear phase SLR pulses (see FIG. 26), and are adjusted to have a bandwidth of (3 THK+2. Delta. S). Gamma. G _s /(2 pi). To decode simultaneous excitation of multiple slice signals from phase encoded magnetic resonance signals and reduce radio frequency power peaksValues, for example, for three-layer synchronous excitation case +.>And->The phase cycling is performed in such a way that +.> And

three scans obtain an image S ₁ ，S ₂ And S is ₃ Then carrying out image decomposition according to the following formula to obtain images of each layer:

the middle-low field imaging generally needs to be accumulated three times or four times clinically, and the synchronous excitation acquisition mode improves the image signal-to-noise ratio by about 40% compared with the conventional single-layer excitation acquisition mode under the condition of the same other scanning parameters.

When the time-band product TBP is large or the magnetic field uniformity is poor, the signal amplitude of the synchronous scanning may be inconsistent, resulting in incomplete image decomposition, see tbp=16 shown in fig. 21, and the amplitude error calibration needs to be performed according to the following steps:

(3) Calibration profile C ₁ 、And->As shown in fig. 23.

In addition, in the case of metal artifact interference, the receiver bandwidth is preferably 100kHz or higher, the echo time is optimized to a minimum, and the frequency encoding gradient and amplitude G are applied simultaneously during data acquisition _s3 And optimizes deltaf according to the operational flow shown in figure 4.

In addition, when the scanning speed is particularly prioritized, G can be selected in the sequence parameter table _blip Gradient option, applying a G at the moment of switching positive and negative polarities of reading gradient _blip Gradient, as shown in FIG. 19, echo number N in the dashed box _pe Can be set to 32 or 64, and the intensity of the phase encoding gradient satisfies G _pe ＝G _blip ·N _pe Under the condition of/2, the scanning time is calculated as

T _ACQ ＝55％·NEX·N _PE ·TR·3＝55％·2·0.3·3≈1(s)

MRI images without artifact interference are then obtained on the navigator system shown in fig. 1 using a combination of pre-scan, phase correction and image reconstruction schemes commonly used for echo planar imaging.

Example 4: fast local arbitrary multi-layer synchronous scanning:

the multilayer synchronous excitation method described in example 3 can also be combined with SLR pulse techniques and generalized to the general case, i.e

Here, the SLR (t) pulse shape is obtained by designing a linear phase digital filter based on Parks-mccclellan algorithm and combining Shinnar-Le-Roux transform algorithm. For example, the five-layer simultaneous excitation pulse waveforms and excitation profiles shown in fig. 24 (left) and (right), where SLR (t) is a linear phase SLR pulse with a time-banded product TBP of 4, an in-banded ripple coefficient of 0.5%, an out-of-band ripple coefficient of 0.5%, a pulse flip angle of pi/2, a selective layer gradient strength of 1.5G/cm, and a pulse bandwidth of 1kHz. Nine layers of simultaneous excitation pulse waveforms and excitation profiles shown in fig. 25 (left) and (right), where SLR (t) is a linear phase SLR pulse with a time-banded TBP of 4, an in-banded ripple coefficient of 0.5%, an out-of-band ripple coefficient of 0.5%, a pulse flip angle of pi/2, a selected layer gradient strength of 1.5G/cm, and a pulse bandwidth of 1kHz. When the scan sequence is of spin echo type, the refocusing pulse is a linear phase SLR pulse as shown in fig. 26, and the characteristic parameters are: the time-band product TBP is 16, the pulse bandwidth is 4kHz, the in-band ripple coefficient is 0.1%, the out-of-band ripple coefficient is 0.1%, the pulse flip angle is pi, and the selected layer gradient strength can be adjusted according to the required layer thickness, for example, the layer gradient strength is set to be 2.0G/cm for the five-layer synchronous excitation case. In the case of an insufficiently sharp image, the optimum pulse width of the desired rf pulse can be obtained according to the detection procedure shown in fig. 7 and the data point interval time of the waveform file can be adjusted until the pulse width reaches the optimum value, while the gradient amplitude in the sequence parameter table is adjusted until the desired layer thickness is obtained. When 2n+1 layers of k-space data are acquired simultaneously, the RF pulses defined by equation (5) can meet the requirement of uniform excitation and no overlapping artifacts at adjacent layers. In clinical navigation applications, accurate positioning of the imaging region is achieved according to the navigation procedure shown in fig. 9. Imaging sequences based on five or more layers of synchronously excited radio frequency pulses can significantly improve the signal to noise ratio by multiple summations under medium and low field conditions, with significantly less scan time increase relative to the asynchronous excitation case.

Example 5: fast local arbitrary five-layer synchronous scanning:

the multilayer synchronous excitation method described in example 4 can also be used for slice imaging at arbitrary locations, i.e

The specific positioning mode is that the corresponding positioning line is set on the positioning image through the user interface shown in FIG. 11And transmits the azimuth parameters of each layer to the SLR waveform calculation program of the navigation module shown in fig. 10, outputs an updated SLR pulse waveform, loads the updated SLR pulse waveform to the navigation imaging control unit shown in fig. 1, and then runs a scanning sequence and performs real-time image reconstruction. When the scan sequence is of spin echo type, the refocusing pulse is a linear phase SLR pulse as shown in fig. 26, and the characteristic parameters are: the time-band product TBP is 16, the pulse bandwidth is 4kHz, the in-band ripple coefficient is 0.1%, the out-of-band ripple coefficient is 0.1%, the pulse flip angle is pi, and the gradient strength of the selected layer can be adjusted according to the required layer thickness, and is set to be 2.0G/cm.

Example 6 local Single-shot T2 weighted scanning

Based on the MRI navigation system and navigation procedure described above, the fast navigation sequence selects the Local-HASTE-NV sequence shown in FIG. 27, and the first 90 RF excitation pulse and 180 refocusing pulse of the sequence select the SLR pulse with highly optimized excitation profile in FIGS. 14 and 15, respectively. The waveform characteristic parameter of the minimum phase SLR excitation pulse shown in FIG. 14 is that the time-band product TBP is 8, the in-band ripple coefficient is 0.5%, the out-of-band ripple coefficient is 0.1%, the pulse flip angle is pi/2, and the selective layer gradient strength is 0.7G/cm; the waveform characteristic parameter of the minimum phase SLR refocusing pulse shown in FIG. 15 is that the time-band product TBP is 16, the in-band ripple coefficient is 0.5%, the out-of-band ripple coefficient is 0.1%, the pulse flip angle is pi, and the selective layer gradient strength is 0.5G/cm. Under the condition of insufficient radio frequency power, the second and subsequent 180 DEG refocusing pulses can be SLR pulses shown in FIG. 26, and under the condition of serious magnetic susceptibility artifact or metal artifact, the optimal SLR pulse width can be obtained according to the detection flow shown in FIG. 7, and the data point interval time of the waveform file can be adjusted until The image sharpness is optimal while the gradient amplitude in the sequence parameter table is adjusted until the desired layer thickness is obtained. In fig. 27, the time interval between the 90 ° rf excitation pulse and the first 180 ° refocusing pulse is TE/2, the time interval between the subsequent 180 ° refocusing pulse is TE, TE is typically in the range of 80ms to 140ms, signal acquisition is started after the third SLR pulse, and N is repeatedly performed in part within the dashed box _PE N/4 times _PE A smaller integer, such as 32 or 64, may be taken. G _s1 、G _s2 、G _s3 、G _s4 、G _s5 …G _sn Representing a selected layer gradient, the gradient amplitude being determined by the layer thickness, the gradient pulse lengths being greater than the RF pulse length, wherein G _s1 、G _S2 And G _S3 The orientation of the gradient can be set to be orthogonal on the positioning image, and the orientation of other selected layer gradients and G _s3 The same applies. G _pre Representing a preliminary read gradient, G _r1 、G _r2 、G _r3 …G _rn Are all denoted by G _pre A read gradient with opposite polarity and twice the integral area, G _sc1 、G _sc2 、G _sc3 …G _scn As selectable items indicating application in the selected layer direction and G _S3 The same amplitude as G _r1 Compensating gradients of the same width. G+ -G _P1 、±G _P2 、±G _p3 …±G _pn The phase encoding gradient which represents positive and negative polarities and sequentially increases or decreases in amplitude is set to zero, the phase encoding gradient corresponding to the effective echo is set to be zero, and the rest echoes are set to be corresponding to the amplitude and the polarity of the phase encoding gradient according to a k-space filling mode. Odd echoes in front of the effective echo sequentially fill the upper half part of the k-space central area, even echoes sequentially fill the lower half part of the k-space central area, parity echoes behind the effective echo respectively fill the upper and lower parts of the k-space central area until the lower half part (accounting for 1/10 of the whole k-space) of the central area is filled, then the rest parity echoes sequentially fill the blank part of the upper half k-space (or the lower half k-space) from inside to outside, and real-time partial Fourier reconstruction is carried out after scanning is finished. Due to N here _PE The length of echo chain is several times smaller than that of HASTE sequence for conventional diagnosis, and the influence of noise and clutter can be obviously reduced so as to ensure navigation chartSuch as higher signal-to-noise ratio and resolution.

Finally, it should be noted that: the above embodiments are only for illustrating the technical solution of the present invention, and not for limiting the same; although the invention has been described in detail with reference to the foregoing embodiments, it will be understood by those of ordinary skill in the art that: the technical scheme described in the foregoing embodiments is modified or some or all of the technical features are replaced equivalently, so that the essence of the corresponding technical scheme does not deviate from the scope of the technical scheme of the embodiments of the present invention.

Claims

1. A navigation system for minimally invasive surgery, comprising: the magnetic yoke (1), magnetic steel (2), polar plates (3), shielding coils (4), gradient coils (5), radio frequency coils (6), scanning beds (7), heating rods (8), heating plates (9), temperature sensors (10), a shell (11), T/R receiving and sending switches (12), gradient power amplifiers (13), radio frequency power amplifiers (14), preamplifiers (15), temperature control units (16), optical tracking and positioning systems (17), coil tuning control units (18), scanning bed control units (19), navigation imaging control units (20), system control interfaces (21), a host (22) and a display (23), and is characterized in that a magnet (60) is of an open U-shaped structure, a magnet gap is larger than 500mm, and the radius of the polar plates (3) is 350-400 mm; the radius of the polar plate (3) is 380mm; the magnet (60) has a field strength between 0.2T and 0.7T; the radio frequency coil (6) of the magnet (60) is of a biplane circular polarization structure, and a plurality of nonmagnetic capacitors (51) are arranged between the outer conductor ring (50) and the inner conductor ring (52); the radio frequency receiving and transmitting coil (6) realizes the conversion of transmitting and receiving through a T/R switch; the number of the nonmagnetic capacitors (51) is more than 100; two groups of X, Y and Z axis gradient coils are respectively arranged, one group is arranged on the inner side of one magnetic pole, the other group is arranged on the inner side of the other magnetic pole, and the coil planes are parallel to the surface of the polar plate (3) and are separated from the polar plate (3) through an anti-vortex plate; the Z-axis gradient coils are composed of a main gradient coil (101) and an axial shielding coil (102), belong to MAXWELL coils, are in a concentric circle type, have 13 turns of the main gradient coil (101) and 14 turns of the axial shielding coil (102), and are connected to a gradient power amplifier (13) through a gradient power filter by 6 gradient cables (104); the main gradient coil (101) and the axial shielding coil (102) are manufactured by adopting a printed circuit board technology; the axial shielding coil (102) is arranged outside the main gradient coil (101) and is close to the polar plate (3), gradient current directions of the axial shielding coil (102) and the main gradient coil (101) are opposite, and magnetic field gradients are limited in the circumference of the axial shielding coil (102), so that thrust generated by interaction of current pulses in the coils and a static magnetic field is counteracted; an insulating layer (103) is arranged between the main gradient coil (101) and the axial shielding coil (102); the radius of the main gradient coil (101), the radius of the axial shielding coil (102) and the radius of the insulating layer (103) are 350mm, the thickness of each coil is 4mm to 5mm, and the distance between the main gradient coil (101) and the axial shielding coil (102) is 2mm to 3mm.