CN111067520B - Magnetic nanoparticle imaging system - Google Patents
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Abstract
The application is suitable for the technical field of magnetic nanoparticle imaging, a magnetic nanoparticle imaging system is provided, the magnetic nanoparticle imaging system comprises a control module, a gradient magnetic field generating module, a driving magnetic field generating module and a signal receiving module, the control module is electrically connected with the driving magnetic field generating module and the signal receiving module, the driving magnetic field generating module generates a variable driving magnetic field under the driving of the control module, the gradient magnetic field generating module generates a gradient magnetic field, the gradient magnetic field and the driving magnetic field are superposed to form an imaging region, magnetic nanoparticles are positioned in the imaging region, the imaging region generates a zero magnetic field point and changes the position of the zero magnetic field point in the imaging region, and the signal receiving module is used for receiving response magnetic field signals of the magnetic nanoparticles at different positions and outputting the response magnetic field signals to the control module for imaging. The magnetic nanoparticle imaging system can detect and obtain response magnetic field signals of magnetic nanoparticles at different positions in an imaging area, and realizes reliable detection of magnetic nanoparticle imaging.
Description
Technical Field
The application belongs to the technical field of magnetic nanoparticle imaging, and particularly relates to a magnetic nanoparticle imaging system.
Background
The computer tomography technology has been developed to a great extent in the past decades, and both the hardware structure and the reconstruction algorithm have greatly improved the imaging speed and accuracy. With the rapid development of technologies such as targeted drug therapy, there is an urgent need for accuracy of therapy, which requires tracking and displaying tracer or contrast agent by using medical equipment, and the tracer and contrast agent are developed by using computed tomography technology, and the most used methods are computed tomography technology, positron emission computed tomography technology, magnetic resonance imaging and other technologies, but the above technologies have various defects, such as low spatial resolution, long detection period, radiation hazard of tracer to human body and the like. Imaging systems based on the various imaging techniques described above have poor reliability.
With the development of materials, a nano-sized magnetic nanoparticle is developed, which has a size of nanometer scale, usually less than 10nm, and when the size of the particle is at this size, it shows an aggregate of single-domain magnetic particles, which has a very high spin quantum number, so that the aggregate is freely arranged under an applied magnetic field, and thus the quantum effect hardly works for the whole system, and at this time, a classical limit model of the curie paramagnetic theory is applied to describe the magnetization behavior of the magnetic system, so the magnetic nanoparticle is also called a superparamagnetic nanoparticle.
Disclosure of Invention
In view of this, the present application provides a magnetic nanoparticle imaging system to solve the problem of poor reliability of the existing imaging system.
The present application provides a magnetic nanoparticle imaging system, comprising:
a control module;
a gradient magnetic field generating module;
a driving magnetic field generating module; and
a signal receiving module;
the control module is electrically connected with the driving magnetic field generating module and the signal receiving module, the driving magnetic field generating module generates a variable driving magnetic field under the driving of the control module, the gradient magnetic field generating module is used for generating a gradient magnetic field, the gradient magnetic field and the driving magnetic field are superposed to form an imaging region, the imaging region is used for generating a zero magnetic field point and changing the position of the zero magnetic field point in the imaging region, the imaging region is used for setting the magnetic nanoparticles, and the signal receiving module is used for receiving response magnetic field signals of the magnetic nanoparticles at different positions and outputting the response magnetic field signals to the control module for imaging.
In one embodiment, the driving magnetic field generating module includes a driving coil unit, and the signal receiving module includes a receiving coil.
In one embodiment, the drive coil unit includes a first drive coil having a radius larger than a radius of the receive coil, the receive coil coaxially sleeved within the first drive coil, the imaging region being within the receive coil.
In one embodiment, the drive coil unit further comprises a second drive coil, a third drive coil, a fourth drive coil and a fifth drive coil, wherein the axes of the second drive coil, the third drive coil, the fourth drive coil and the fifth drive coil are in the same plane, and the plane is perpendicular to the axis of the first drive coil; the second driving coil and the third driving coil are coaxially arranged, the axis of the second driving coil is perpendicular to the axis of the first driving coil, the second driving coil and the third driving coil are respectively arranged on two sides of the first driving coil, and the second driving coil and the third driving coil are connected in series; the fourth driving coil and the fifth driving coil are coaxially arranged, the axis of the fourth driving coil is perpendicular to the axis of the first driving coil, the fourth driving coil and the fifth driving coil are respectively arranged on two sides of the first driving coil, and the fourth driving coil and the fifth driving coil are connected in series; the axis of the second drive coil is perpendicular to the axis of the fourth drive coil.
In one embodiment, the signal receiving module further includes a first noise cancellation coil and a second noise cancellation coil for reducing noise, the first noise cancellation coil and the second noise cancellation coil have equal lengths, the radii of the first noise cancellation coil and the second noise cancellation coil are equal to the radius of the receiving coil, the first noise cancellation coil and the second noise cancellation coil are coaxially disposed with the receiving coil, the first noise cancellation coil and the second noise cancellation coil are respectively disposed on two sides of the receiving coil, the distance between the first noise cancellation coil and the receiving coil is equal to the distance between the second noise cancellation coil and the receiving coil, and the first noise cancellation coil and the second noise cancellation coil are sleeved in the first driving coil; and the first noise cancellation coil and the second noise cancellation coil are used for leading currents with equal magnitude and opposite directions.
In one embodiment, the control module comprises an FPGA.
In one embodiment, the control module comprises a driving module, and the FPGA is electrically connected with the driving module; the driving module comprises a first driving signal generating unit, a second driving signal generating unit, a third driving signal generating unit and a digital-to-analog converter, wherein the first driving signal generating unit, the second driving signal generating unit and the third driving signal generating unit are electrically connected with the digital-to-analog converter, the digital-to-analog converter is provided with three driving signal output ends, the first driving signal output end is electrically connected with the first driving coil, the second driving signal output end is electrically connected with the second driving coil, and the third driving signal output end is electrically connected with the fourth driving coil; the digital-to-analog converter is used for performing digital-to-analog conversion on the first initial driving signal generated by the first driving signal generating unit to generate a first target driving signal, and then outputting the first target driving signal to the first driving coil through the first driving signal output end; the digital-to-analog converter is used for performing digital-to-analog conversion on a second initial driving signal generated by the second driving signal generating unit to generate a second target driving signal, and then outputting the second target driving signal to the second driving coil through the second driving signal output end; the digital-to-analog converter is used for performing digital-to-analog conversion on a third initial driving signal generated by the third driving signal generating unit to generate a third target driving signal, and then outputting the third target driving signal to the fourth driving coil through the third driving signal output end; the first target driving signal, the second target driving signal and the third target driving signal are alternating current driving signals with different frequencies.
In one embodiment, the gradient magnetic field generating module comprises a pair of permanent magnets with the same magnetic poles oppositely arranged, and the pair of permanent magnets generates the gradient magnetic field.
In one embodiment, the gradient magnetic field generation module further comprises a distance adjustment mechanism for adjusting the relative distance between the pair of permanent magnets.
In one embodiment, the magnetic nanoparticle imaging system further comprises an upper computer electrically connected with the control module.
Compared with the prior art, the embodiment of the application has the advantages that: in the magnetic nanoparticle imaging system provided by the application, the driving magnetic field generating module generates a variable driving magnetic field, the gradient magnetic field generating module generates a gradient magnetic field, the gradient magnetic field and the driving magnetic field are superposed to form an imaging region, based on the variable driving magnetic field, the position of a zero magnetic field point in the imaging region can be changed, response magnetic field signals of magnetic nanoparticles at different positions in the imaging region can be detected, and the control module receives response magnetic field information to perform imaging. The magnetic nanoparticle imaging system can detect response magnetic field signals of magnetic nanoparticles at different positions in an imaging area based on the gradient magnetic field and the variable driving magnetic field, so that reliable imaging of the magnetic nanoparticles is realized, and the imaging stability of the magnetic nanoparticles is ensured.
Drawings
In order to more clearly illustrate the technical solutions in the embodiments of the present application, the drawings used in the embodiments or the description of the prior art will be briefly described below, and it is obvious that the drawings in the following description are only some embodiments of the present application, and it is obvious for those skilled in the art that other drawings can be obtained according to these drawings without inventive labor.
FIG. 1 is a schematic diagram of a first configuration of a magnetic nanoparticle imaging system provided by an embodiment of the present application;
FIG. 2 is a schematic diagram of a second configuration of a magnetic nanoparticle imaging system provided by an embodiment of the present application;
FIG. 3 is a schematic structural diagram of a driving module of a magnetic nanoparticle imaging system provided by an embodiment of the present application;
FIG. 4 is a diagram of a data transfer process for drive signal generation;
fig. 5 is a perspective view of a first drive coil and a receive coil;
figure 6 is a side view of a first drive coil, a receive coil, a second drive coil, a third drive coil, a fourth drive coil, and a fifth drive coil;
fig. 7 is a schematic perspective view of the first driving coil, the receiving coil, the first permanent magnet, and the second permanent magnet;
fig. 8 is a schematic structural view of a receiving coil, a first noise canceling coil and a second noise canceling coil;
FIG. 9 is an idealized grayscale image of a custom particle distribution;
FIG. 10 is a particle reconstruction map;
figure 11 is a noise map that is actively added to simulate real-world conditions.
Detailed Description
In the following description, for purposes of explanation and not limitation, specific details are set forth, such as particular system structures, techniques, etc. in order to provide a thorough understanding of the embodiments of the present application. It will be apparent, however, to one skilled in the art that the present application may be practiced in other embodiments that depart from these specific details. In other instances, detailed descriptions of well-known systems, devices, circuits, and methods are omitted so as not to obscure the description of the present application with unnecessary detail.
It will be understood that the terms "comprises" and/or "comprising," when used in this specification and the appended claims, specify the presence of stated features, integers, steps, operations, elements, and/or components, but do not preclude the presence or addition of one or more other features, integers, steps, operations, elements, components, and/or groups thereof.
It should also be understood that the term "and/or" as used in this specification and the appended claims refers to and includes any and all possible combinations of one or more of the associated listed items.
It is also to be understood that the terminology used in the description of the present application herein is for the purpose of describing particular embodiments only and is not intended to be limiting of the application. As used in the specification of the present application and the appended claims, the singular forms "a," "an," and "the" are intended to include the plural forms as well, unless the context clearly indicates otherwise.
It should be understood that the order of writing each step in this embodiment does not mean the order of execution, and the order of execution of each process should be determined by its function and inherent logic, and should not constitute any limitation on the implementation process of the embodiment of the present invention.
Reference throughout this specification to "one embodiment" or "some embodiments," or the like, means that a particular feature, structure, or characteristic described in connection with the embodiment is included in one or more embodiments of the present application. Thus, appearances of the phrases "in one embodiment," "in some embodiments," "in other embodiments," or the like, in various places throughout this specification are not necessarily all referring to the same embodiment, but rather "one or more but not all embodiments" unless specifically stated otherwise. The terms "comprising," "including," "having," and variations thereof mean "including, but not limited to," unless expressly specified otherwise.
In order to explain the technical means described in the present application, the following description will be given by way of specific embodiments.
The relationship between the magnetic nanoparticles and the magnetic field will be briefly described below.
The magnetic nano-particles can generate nonlinear magnetization response in an alternating magnetic field, and can generate nonlinear magnetization response only when the magnetic moments of the magnetic nano-particles do not reach a saturated state, so that a gradient magnetic field and a driving magnetic field can be used for generating a zero magnetic field point and a driving zero magnetic field point to scan the magnetic nano-particles, only the magnetic nano-particles near the zero magnetic field point can generate nonlinear magnetization response, and the magnetic moments of the particles far away from the zero magnetic field point reach the saturated state and cannot generate nonlinear magnetization response, so that by using the performance, the nonlinear magnetization response can be detected by using a signal receiving module (such as a receiving coil), the positions of the magnetic nano-particles can be further positioned, and the imaging and further analysis can be performed on the magnetic nano-particles by using the principle.
Scanning modes that traverse the imaging region using zero field point scanning include: spiral trajectories, radial trajectories, lissajou trajectories, cartesian trajectories, and the like.
Referring to fig. 1, a schematic diagram of a first structure of a magnetic nanoparticle imaging system according to an embodiment of the present application is provided.
As shown in fig. 1, the magnetic nanoparticle imaging system provided by the present application includes a control module 101, a gradient magnetic field generation module 102, a driving magnetic field generation module 103, and a signal receiving module 104. As shown in fig. 1, the control module 101 is electrically connected to the driving magnetic field generating module 103 and the signal receiving module 104.
The control module 101 is the core of control and data processing, and controls the operation of the whole system, including the output of driving signals and the reception of receiving signals, and the operation speed of the controller is important, which affects the speed of image reconstruction. The specific implementation manner of the control module 101 is not exclusive, and a Central Processing Unit (CPU) may be used as the controller, or other types of controllers may be used, such as an FPGA (Field Programmable Gate Array). Although the CPU can meet the requirement on signal transmission, the internal structure of the CPU cannot be changed once the CPU leaves a factory, the flexibility of later debugging is reduced, and the internal part of the CPU is of a von Neumann structure and executes instructions one by one in an internal serial mode, so that the program running time is greatly reduced, and the imaging running speed is reduced. If the FPGA with a full-programmable logic array structure is adopted as a controller of the system, compared with a CPU, the FPGA not only has the software programmable capability and the hardware programmable capability, but also realizes stronger system performance, flexibility and expandability by using the advantage of a low-power-consumption system; the FPGA has the advantages of parallel operation algorithm, greatly quickens the operation process of the system, reduces the imaging time, can reduce the peripheral hardware circuit and the research and development cost due to the programmable performance and the function of describing hardware, and quickens the universal utilization rate of the system.
The control module 101 outputs a driving signal to the driving magnetic field generating module 103, and generates a variable driving magnetic field under the action of the driving signal, for example: if the driving signal is a sinusoidal signal, the driving magnetic field is an alternating magnetic field. In general, the driving magnetic field generating module 103 is a driving coil, and the variable driving magnetic field is generated by applying a variable driving signal to the driving coil. The gradient magnetic field generating module 102 is configured to generate a gradient magnetic field, where the gradient magnetic field generating module 102 may be a pair of electromagnetic coils or a pair of permanent magnets (e.g., permanent magnets), and when the gradient magnetic field generating module 102 is a pair of electromagnetic coils, the electromagnetic coils generate a gradient magnetic field with gradient properties by passing a direct current; when the gradient magnetic field generation module 102 is a pair of permanent magnets, the same magnetic poles of the two permanent magnets are arranged oppositely, so that a gradient magnetic field can be generated. The strength of the gradient magnetic field generated by the gradient magnetic field generation module 102 may be fixed or may vary. When the gradient magnetic field generation module 102 needs to generate a variable gradient magnetic field, if the gradient magnetic field generation module 102 is a pair of electromagnetic coils, the strength of the gradient magnetic field can be changed by changing the introduced direct current; if the gradient magnetic field generating module 102 is a pair of permanent magnets, the relative distance between the two permanent magnets needs to be adjusted, at least one of the two permanent magnets can be directly moved manually, or the adjustment can be performed by a distance adjusting mechanism.
The gradient magnetic field and the driving magnetic field are superposed to form an imaging area, and the imaging area can be in a one-dimensional, two-dimensional or three-dimensional structure according to the position relation between the gradient magnetic field and the driving magnetic field. A field-free point can be generated in the imaging region and the position of the field-free point within the imaging region can be varied by means of a variable drive field. The magnetic nanoparticles are within the imaging region.
In the imaging region, the magnetic nanoparticles at each position have a corresponding response magnetic field signal, and then the signal receiving module 104 receives the response magnetic field signals of the magnetic nanoparticles at different positions in the imaging region, where the magnetic field signal is specifically the concentration of the magnetic nanoparticles at each position. The signal receiving module 104 is generally a receiving coil, and the signal receiving module 104 receives the magnetic field signal and converts the magnetic field signal into an electrical signal, specifically: when the signal receiving module 104 is in a changing magnetic field, the change of the magnetic flux in the coil will cause the signal receiving module 104 to generate an induced electromotive force (i.e. an induced voltage). Then, when the signal receiving module 104 receives the magnetic field signal, a corresponding induced voltage signal is output to the control module 101. The control module 101 can output the data signal to an external computer, such as an upper computer, and realize imaging through the external computer; if the control module 101 itself is provided with a display, the control module 101 itself can perform imaging, and the system does not need to be provided with an external computer.
Referring to fig. 2, a schematic diagram of a second structure of a magnetic nanoparticle imaging system according to an embodiment of the present disclosure is shown.
As shown in fig. 2, the magnetic nanoparticle imaging system includes a control module 201, a linear power amplifier 202, a driving magnetic field generation module 203, a signal receiving module 204, a gradient magnetic field generation module 205, a band-pass filter 207, an operational amplifier 208, and an upper computer 209. The control module 201 is electrically connected with the linear power amplifier 202, the linear power amplifier 202 is electrically connected with the driving magnetic field generating module 203, the signal receiving module 204 is electrically connected with the band-pass filter 207, the band-pass filter 207 is electrically connected with the operational amplifier 208, and the operational amplifier 208 is electrically connected with the upper computer 209. The magnetic nanoparticles 206 are in the imaging region formed by the superposition of the gradient magnetic field and the driving magnetic field.
The control module 201 comprises an FPGA (field programmable gate array), and the FPGA not only has the capability of software programming and the capability of hardware programming, but also realizes stronger system performance, flexibility and expandability by using the advantage of a low-power-consumption system; the advantage of using FPGA to run algorithm in parallel greatly accelerates the running process of the system and reduces the imaging time, and because the FPGA has programmable performance and the function of describing hardware, the peripheral hardware circuit and the development cost can be reduced, and the universal utilization rate of the system is accelerated.
The FPGA in the control module 201 outputs a driving signal to drive the driving magnetic field generating module 203 to generate a driving magnetic field, so the control module 201 further includes a driving module for generating the driving signal, and the FPGA is electrically connected to the driving module. In this embodiment, as shown in fig. 3, the driving module includes a first driving signal generating unit 2011, a second driving signal generating unit 2012, a third driving signal generating unit 2013, and a digital-to-analog converter 2015, and further includes a digital-to-analog driving unit 2014 adapted to the digital-to-analog converter 2015, where the first driving signal generating unit 2011, the second driving signal generating unit 2012, and the third driving signal generating unit 2013 are electrically connected to the digital-to-analog driving unit 2014, and the digital-to-analog driving unit 2014 is electrically connected to the digital-to-analog converter 2015.
In this embodiment, a DDS (Direct Digital Synthesis) technique is used to generate the driving signal, and since the purity of the signal generated in this way is high, a filter is not required to be added to the output portion, and the cost can be reduced to a certain extent. The driving signal plays an important role in the whole system design, and the quality of the driving signal directly influences the reconstruction of a later-stage image. Because the accuracy requirement of the system on the accuracy of the signal is very high, and the common way of generating the driving signal by using PWM (Pulse Width Modulation) cannot meet the requirements of accuracy and easy frequency adjustment, the present embodiment adopts the DDS technology to generate the driving signal with different frequencies and different types that can be easily operated and quickly changed. For any one of the first drive signal generation unit 2011, the second drive signal generation unit 2012 and the third drive signal generation unit 2013, the internal DDS part mainly includes: frequency control words, phase accumulators, and phase converters. The frequency value of the driving signal required in the system is calculated as follows:
wherein: f. of OUT For the frequency value of the drive signal required in the system, N is f corresponding to the minimum incremental phase change of the phase accumulator output word clk The number of pulses of f clk For DDS operation clock, M represents the number of bits of the phase accumulator, by adjusting the phase accumulatorAnd the sum pulse number can adjust the output frequency. In this embodiment, the DDS is encapsulated into an IP core, and then the IP core is directly called in a development environment to generate a digital signal, so each driving signal generating unit is essentially an IP core that encapsulates the DDS. By using the DDS IP core with the same three-way structure, the frequency of the driving signal required by the system is changed only by changing the output frequency, and the requirement of the frequency forming a specific scanning track in the system is generated. In general, the frequencies of the three driving signals are different, but there is only a small difference, that is, the difference between the frequencies of any two driving signals is smaller than a preset error threshold, such as a driving signal with a ratio of the three frequencies of 99. By placing in phase, a lissajous scan trajectory can be generated, exciting the magnetic nanoparticles 206 at different locations. The three driving signals output from the respective IP cores are transmitted at byte speed in the case of a transmission clock. The embodiment adopts a digital signal staggered transmission mode, and the transmission has the advantages that multiple signals can be transmitted only by one input, so that the utilization rate of a chip is increased, the power consumption is reduced, and the hardware cost is reduced. As shown in fig. 4, a data transmission process diagram is formed for the driving signal. In this embodiment, the initial driving signal is a sinusoidal amplitude signal, and is input to the digital-to-analog converter 2015 through the digital-to-analog driving unit, and a sinusoidal analog driving signal is output through digital-to-analog conversion, where the sinusoidal analog driving signal is a target driving signal. The method comprises the following specific steps: the first driving signal generating unit 2011 generates a first initial driving signal, and the digital-to-analog converter 2015 performs digital-to-analog conversion on the first initial driving signal to generate a first target driving signal; the second driving signal generating unit 2012 generates a second initial driving signal, and the digital-to-analog converter 2015 performs digital-to-analog conversion on the second initial driving signal to generate a second target driving signal; the third driving signal generating unit 2013 generates a third initial driving signal, and the digital-to-analog converter 2015 performs digital-to-analog conversion on the third initial driving signal to generate a third target driving signal. Thus, the three target drive signals are three sinusoidal analog drive signals of different frequencies.
The digital-to-analog converter 2015 has three driving signal output ends, which are a first driving signal output end, a second driving signal output end and a third driving signal output end, respectively, where the first driving signal output end outputs a first target driving signal, the second driving signal output end outputs a second target driving signal, and the third driving signal output end outputs a third target driving signal.
The linear power amplifier 202 is configured to amplify the first target driving signal, the second target driving signal, and the third target driving signal, and the linear power amplifier 202 may be a conventional linear power amplifier device, which is not described again.
The driving magnetic field generating module 203 includes a driving coil unit, and the signal receiving module 204 includes a receiving coil 2041. The driving coil unit includes a first driving coil 2031, as shown in fig. 5, the radius of the first driving coil 2031 is larger than the radius of the receiving coil 2041, the receiving coil 2041 is coaxially sleeved in the first driving coil 2031, and the difference between the radius of the first driving coil 2031 and the radius of the receiving coil 2041 is set by actual conditions. In order to fix the first driving coil 2031 and the receiving coil 2041, a fixing structure may be provided, such as: a cylindrical structure is provided, and the first driving coil 2031 is fitted around the outer wall of the cylindrical structure, and the receiving coil 2041 is coaxially placed inside the cylindrical structure. The cylindrical structure needs to be made of a material that does not affect the normal operation of the first driving coil 2031 and the receiving coil 2041, such as: the magnetic field generated by the first drive coil 2031 cannot be absorbed or attenuated and the normal operation of the magnetic nanoparticles 206 cannot be affected. In addition to the above-described structure, the first driving coil 2031 and the receiving coil 2041 may be fixed using a fixing bracket.
The drive coil unit further includes a second drive coil 2032, a third drive coil 2033, a fourth drive coil 2034, and a fifth drive coil 2035. As shown in fig. 6, the axes of the second drive coil 2032, the third drive coil 2033, the fourth drive coil 2034, and the fifth drive coil 2035 are on the same plane, and the plane is perpendicular to the axis of the first drive coil 2031. The second driving coil 2032 and the third driving coil 2033 are coaxially disposed, the axis of the second driving coil 2032 (i.e., the axis of the third driving coil 2033) is perpendicular to the axis of the first driving coil 2021, the second driving coil 2032 and the third driving coil 2033 are disposed on both sides of the first driving coil 2031, respectively, and the second driving coil 2032 and the third driving coil 2033 are electrically connected in series. The fourth drive coil 2034 and the fifth drive coil 2035 are coaxially disposed, the axis of the fourth drive coil 2034 (i.e., the axis of the fifth drive coil 2035) is perpendicular to the axis of the first drive coil 2031, the fourth drive coil 2034 and the fifth drive coil 2035 are disposed on both sides of the first drive coil 2031, respectively, and the fourth drive coil 2034 and the fifth drive coil 2035 are electrically connected in series. The axis of the second driving coil 2032 is perpendicular to the axis of the fourth driving coil 2034. If the second driving coil 2032 and the third driving coil 2033 are set as a coil pair a, and the fourth driving coil 2034 and the fifth driving coil 2035 are set as a coil pair b, the axes of the first driving coil 2031, the coil pair a, and the coil pair b are equivalent to the x-axis, the y-axis, and the z-axis in the three-dimensional coordinates, and the three axes are perpendicular to each other two by two. Therefore, the second driving coil 2032 and the third driving coil 2033 in the coil pair a may be a pair of driving coils of a helmholtz structure, and the fourth driving coil 2034 and the fifth driving coil 2035 in the coil pair b may be a pair of driving coils of a helmholtz structure. Further, the second drive coil 2032, the third drive coil 2033, the fourth drive coil 2034 and the fifth drive coil 2035 have the same geometric parameters, i.e., the same length and the same radius. The second drive coil 2032, the third drive coil 2033, the fourth drive coil 2034, and the fifth drive coil 2035 are also equidistant from the first drive coil 2031. In order to fix the second driving coil 2032, the third driving coil 2033, the fourth driving coil 2034, and the fifth driving coil 2035, fixing structures such as: four fixing brackets are arranged, and the driving coils are fixed according to respective positions; or a mounting case is made based on the positional relationship of the four drive coils. The housing needs to be constructed of materials that do not interfere with the proper operation of the four drive coils, such as: the generated magnetic field cannot be absorbed or weakened.
A first driving signal output end of the digital-to-analog converter 2015 is electrically connected with the first driving coil 2031, and transmits a first target driving signal to the first driving coil 2031; a second drive signal output terminal of the digital-to-analog converter 2015 is electrically connected to the second drive coil 2032, and transmits a second target drive signal to the second drive coil 2032 and the third drive coil 2033; a third driving signal output terminal of the digital-to-analog converter 2015 is electrically connected to the fourth driving coil 2034 and transmits a third target driving signal to the fourth driving coil 2034 and the fifth driving coil 2035.
The drive coil unit is used for driving the zero magnetic field point to traverse the whole imaging area. In this embodiment, each drive coil is in the form of a solenoid that drives a field-free point to form a three-dimensional image. The structure of the driving coil unit can form a stable and reliable three-dimensional driving magnetic field. The imaged area can be calculated by: the driving magnetic field is calculated in a finite element mode, the peak value of alternating current is simulated through the direct current electric excitation coil, and then the specific range of an imaging area can be calculated by dividing the driving field value by the gradient field. In practical application, a large current needs to be introduced to achieve a large alternating magnetic field, if a driving coil is driven by directly using voltage generated by a power amplifier, a large voltage value is needed, so that the voltage requirement on the power amplifier is high, and the heat productivity of the driving coil is increased suddenly, so that a matching series resonance method can be used for the driving coil, a large current can be generated only by needing a small amount of current, the requirement of generating a low-voltage large current on the driving coil is met, and meanwhile, in order to avoid the skin effect generated by high-frequency current to increase the power loss of the driving coil, in the embodiment, each driving coil uses litz wire as a winding material. The magnetic field generated for the drive coils can be described using the biot savart theorem, which is given by:
wherein the content of the first and second substances,
is the magnetic induction intensity generated by the current element Idl at one point P in space, I is the current source at the very small part of the line in the driving coil, L is the integral path of the whole driving coil, dl is the tiny line element of the source current, mu 0 Is a vacuum magnetic permeability of 4 π × 10 -7 Tm/A，/>
Is the strength of any current carrying wire at P.
The drive magnetic field formed by the drive coil is obtained by performing line integration on the drive coil.
The gradient magnetic field generating module 205 includes a pair of permanent magnets with the same magnetic poles arranged oppositely, namely a first permanent magnet 2051 and a second permanent magnet 2052, and both the first permanent magnet 2051 and the second permanent magnet 2052 can be conventional permanent magnets. The reason for using permanent magnets to generate gradient fields is that the cost can be reduced, the power supply for the electromagnetic coil is eliminated, and the manufacturing cost of the permanent magnets is much lower than that of the electromagnetic coil. Before the gradient magnetic field is constructed, modeling simulation can be performed on the gradient magnetic field which plays a role in encoding so as to construct a gradient magnetic field which meets the requirement, and the larger the gradient magnetic field is, the smaller the magnetic nanoparticles 206 which can be detected is, and the larger the particle signal intensity is. In this embodiment, the first permanent magnet 2051 and the second permanent magnet 2052 are both cylindrical structures, and the outer diameters and thicknesses of the first permanent magnet 2051 and the second permanent magnet 2052 are 124mm and 20mm, respectively. In the process of constructing the gradient magnetic field, the first permanent magnet 2051 and the second permanent magnet 2052 are coaxially arranged, the same magnetic poles of the first permanent magnet 2051 and the second permanent magnet 2052 are oppositely arranged, specifically, the S poles of the first permanent magnet 2051 and the second permanent magnet 2052 are oppositely arranged, so that the gradient magnetic field is generated, and the generation of the required zero magnetic field point is realized. The driving magnetic field generating module 203 and the signal receiving module 204 are disposed in a gradient magnetic field, and fig. 7 shows a specific arrangement in which the first driving coil 2031, the first permanent magnet 2051, and the second permanent magnet 2052 are coaxially disposed, and the first permanent magnet 2051 and the second permanent magnet 2052 are disposed on both sides of the first driving coil 2031, respectively.
The magnetic field strength is a trend of linear change within a certain range from the zero magnetic field point, and the present embodiment uses the magnetic field of the deformation change to code the position of the magnetic nanoparticles 206. In order to increase or reduce the imaging area of the system, the gradient magnetic field is designed into an adjustable gradient magnetic field, so that the flexibility of the system is improved, and the requirements of different imaging areas are met. Then, a distance adjustment mechanism for adjusting the relative distance between the first permanent magnet 2051 and the second permanent magnet 2052 needs to be provided. The permanent magnets with different distances generate gradient magnetic fields with different gradient values, and the fact that the signal of the particles and the gradient value of the gradient magnetic field are linearly related is known through the langevin theory, the larger the gradient value is, the stronger the signal of the particles with the same diameter is, and the higher the resolution is (the resolution refers to the minimum distance between adjacent particle clusters which can be observed), so that the application range of different particle diameters is enlarged through adjusting the gradient value, and the application occasion of the system is enlarged. When the relative distance between the first permanent magnet 2051 and the second permanent magnet 2052 is adjusted, only one of the first permanent magnet 2051 and the second permanent magnet 2052 may be adjusted, or the first permanent magnet 2051 and the second permanent magnet 2052 may be adjusted at the same time. The distance adjusting mechanism can be implemented in many ways, such as: taking the regulation of first permanent magnet 2051 as an example, distance adjustment mechanism includes the installation base, installation base fixed mounting threaded sleeve, threaded sleeve cooperatees with a threaded rod and constitutes thread drive and is connected, the threaded rod sets up in threaded sleeve, the one end of threaded rod is provided with and is used for driving threaded rod pivoted driving hand wheel, the other end of threaded rod is provided with the mounting, first permanent magnet 2051 sets up with the mounting is fixed, so, when driving hand wheel rotates, the threaded rod rotates, just can drive the mounting back-and-forth movement, and then drive first permanent magnet 2051 back-and-forth movement, realize adjusting and the relative distance between the second permanent magnet 2052. Therefore, if only one permanent magnet needs to be adjusted, only one distance adjusting mechanism needs to be arranged; if two permanent magnets need to be adjusted simultaneously, two distance adjusting mechanisms as described above need to be provided. As other embodiments, the distance adjusting mechanism may also be in other structural forms such as a slide rail. In addition, the first permanent magnet 2051 and the second permanent magnet 2052 can be manually moved directly without a distance adjusting mechanism to change the relative distance between the two.
The driving magnetic field generating module 203 and the gradient magnetic field generating module 205 form a scanner portion of the system, the driving magnetic field generated by the driving magnetic field generating module 203 and the gradient magnetic field generated by the gradient magnetic field generating module 205 are superposed in an orthogonal manner, a zero magnetic field point can traverse a region of interest (the region of interest includes an imaging region, i.e., a region containing the magnetic nanoparticles 206) under the action of the driving magnetic field, and the magnetic nanoparticles 206 near the zero magnetic field point can generate a nonlinear magnetization response under the action of the driving magnetic field.
In order to reduce noise, cancel a part of or even all of the driving signal at the receiving coil 2041, and reduce the introduction of external noise, the signal receiving module further includes a first noise canceling coil 2042 and a second noise canceling coil 2043 in addition to the receiving coil 2041. As shown in fig. 8, the lengths of the first noise canceling coil 2042 and the second noise canceling coil 2043 are equal, the radii of the first noise canceling coil 2042 and the second noise canceling coil 2043 are equal to the radius of the receiving coil 2041, the first noise canceling coil 2042, the second noise canceling coil 2043 and the receiving coil 2041 are coaxially disposed, the first noise canceling coil 2042 and the second noise canceling coil 2043 are respectively disposed on both sides of the receiving coil 2041, the distance between the first noise canceling coil 2042 and the receiving coil 2041 is equal to the distance between the second noise canceling coil 2043 and the receiving coil 2041, and the first noise canceling coil 2042 and the second noise canceling coil 2043 are sleeved in the first driving coil 2031 as the receiving coil 2041. Therefore, the geometric parameters of the first noise canceling coil 2042 and the second noise canceling coil 2043 are identical. The first noise cancellation coil 2042 and the second noise cancellation coil 2043 are used for passing currents with equal magnitude and opposite directions. Therefore, the first noise canceling coil 2042 and the second noise canceling coil 2043 use a winding method for reducing the driving signal to reduce the driving signal generated by the driving magnetic field at the receiving coil 2041. Since the imaging region is relatively distant from the first noise canceling coil 2042 and the second noise canceling coil 2043, the signals received by the first noise canceling coil 2042 and the second noise canceling coil 2043 from the particles on the coils are negligible. By passing currents of the same magnitude and opposite directions, the strong driving signal at the receiver coil 2041 is reduced. In an ideal state, the first noise canceling coil 2042 and the second noise canceling coil 2043 can completely cancel the driving signal, and an external filter can be omitted, so that the cost is reduced, and extra noise generated by the filter itself is also avoided.
In order to fix the receiving coil 2041, the first noise canceling coil 2042, and the second noise canceling coil 2043, fixing structures such as: the fixing structure is a pipe hole (i.e., a solenoid hole) having a certain length, and the receiving coil 2041, the first noise canceling coil 2042, and the second noise canceling coil 2043 are inserted into the pipe hole in the above-described arrangement. The pipe hole needs to be made of a material that does not affect the normal operation of the receiving coil 2041, the first noise canceling coil 2042, and the second noise canceling coil 2043, such as: the normal operation of the magnetic nanoparticles 206 cannot be affected; or the fixing structure is a casing, a plurality of threading holes are formed in the casing, and the receiving coil 2041, the first noise canceling coil 2042 and the second noise canceling coil 2043 penetrate through the corresponding threading holes according to the arrangement mode. The housing needs to be made of a material that does not affect the normal operation of the receiving coil 2041, the first noise canceling coil 2042, and the second noise canceling coil 2043, for example: the normal operation of the magnetic nanoparticles 206 cannot be affected. In addition to the above two implementations, the receiving coil 2041, the first noise canceling coil 2042, and the second noise canceling coil 2043 may be fixed in other configurations.
After the scanner portion is constructed, the imaging region is within the receive coil 2041 and the magnetic nanoparticles 206 are magnetically responsive within the imaging region. A test object, such as a test body membrane, is placed in the receiving coil 2041, and the first noise cancellation coil 2042 and the second noise cancellation coil 2043 contain only the same and small amount of particle signals.
In the imaging region, the magnetic nanoparticles 206 at each position have a corresponding response magnetic field signal, and the receiving coil 2041 receives the response magnetic field signals of the magnetic nanoparticles 206 at different positions in the imaging region, where the magnetic field signal is specifically the concentration of the magnetic nanoparticles 206 at each position.
The receiving coil 2041 receives the magnetic field signal, converts the magnetic field signal into an electrical signal, specifically: when the receiving coil 2041 is in a changing magnetic field, the change in magnetic flux in the coil causes the receiving coil 2041 to generate an induced electromotive force (i.e., an induced voltage). The voltage signal in an analog form is digitized, in this embodiment, a 14M single-channel analog-to-digital conversion chip is selected, and the analog signal passes through the analog-to-digital conversion chip and then is output to the FPGA in the control module 201 through the band-pass filter 207 and the operational amplifier 208 in sequence. The band-pass filter 207 and the operational amplifier 208 are used as data optimization devices, and since the first noise cancellation coil 2042 and the second noise cancellation coil 2043 can greatly reduce noise, the band-pass filter 207 may not be provided; the operational amplifier 208 is used to amplify data, and if the data does not need to be amplified or if the data does not have to be amplified to satisfy the processing requirement, the operational amplifier 208 may not be provided.
Since the analog signal passes through the analog-to-digital conversion chip and then is serial data, in order to achieve the purpose of real-time image display, the embodiment uses the FPGA to construct a First In First Out (FIFO) buffer, and buffers the received data, and the data in the FIFO buffer adopts a First in First out processing mode, so that the data loss caused by that the calculation is not completed and the data is transmitted too fast due to the excessively fast data transmission speed at the front end can be avoided. And the written FIFO program is packaged into an IP core in the FPGA for receiving data, so that the data can be conveniently and directly called later. Since the received data has a certain noise and compensates for the time consumed by the reconstruction part, the present embodiment uses a moving average data method to transmit data, and when new data arrives, 5 bytes that have been saved are selected for average processing, and the calculation formula is as follows:
wherein the content of the first and second substances,
is the current raw data, u, captured by the data acquisition j The frame data which is the immediately preceding frame data next to the currently captured frame data is represented by 5 byte data.
After FFT (Fast Fourier Transform) conversion is performed on the obtained average data, since higher harmonics are more indicative of acquired noise, the signal of particles is mainly concentrated on 3 rd order harmonics and 5 th order harmonics. Thus, the energies of the 3 rd and 5 th harmonics are selected to represent the response signals of the particles, and the intensities of the 3 rd and 5 th harmonics of the particle signals are selected as the intensities of the particle signals using the FPGA.
The FPGA then reconstructs the position and concentration of the magnetic nanoparticles 206 using an ART (Algebra Reconstruction) projection Reconstruction algorithm, which is described below.
After the two magnetic fields are subjected to simulation verification, in order to show whether the magnetic fields meet the requirements of gradient values of the magnetic fields and amplitude values of driving fields required by system operation or not, the 2D platform is verified,
wherein->
Representing a contrast matrix obtained by moving a probe made of known magnetic nanoparticles 206 across the entire imaging area on a three-dimensional stage, m representing the number of points moved in unison, n representing the number of harmonics used per point, or->
Representing the signal of the magnetic nanoparticles 206 measured using the test body film, x being the value to be sought.
An ART projection reconstruction algorithm is applied to the system to reconstruct the position and concentration of the magnetic nanoparticles 206, and the specific steps are as follows: by multiplying the image F of the known matrix and position of m x n and then equaling its value with the measured voltage signal, the reconstruction can be expressed using the following system of equations:
wherein: a is mn Is an element of an m x n matrix, m is the harmonic number of each selected magnetic nanoparticle 206, n is the position of all selected magnetic nanoparticles 206, u is the measured voltage value, and the measured harmonic energy value is used as the intensity value of the image gray scale map.
Since the larger the particle concentration is, the stronger the signal of the particle is, by solving the above equation set, an image having particles linearly related to the concentration can be obtained, and the solved data is stored in the form of a matrix and then subjected to gray scale transformation. Through simulation of a mathematical platform, selecting the minimum gradient value and the driving field amplitude obtained through simulation, calculating the nonlinear response of the particles by using the langevin theory, then simulating the response of the particles in a real environment, and finally obtaining the distribution diagram of the particles.
As shown in FIG. 9, the profile of the custom particles, where the white 20mm by 5mm, white portion of the particles at a spacing of 5mm represents the ideal clumping of particles. By calculating the displayed gray scale map, the distribution of particles can be roughly obtained, as shown in fig. 10. If the gradient value is increased, the resolution is improved, and fig. 11 shows gaussian noise added to simulate noise in a real environment.
Finally, the PFGA uploads the image data to the upper computer 209 in an SPI mode, and the image is displayed in real time through the upper computer 209.
The adjustable permanent magnet gradient field structure expands the use occasions, for example, the possibility of being used on a human body is provided, compared with the mode of using an electromagnet, the gradient field generated by using a permanent magnet greatly reduces the cost because a direct current power supply of the electromagnet is omitted; the inherent advantages of the parallel computation of the FPGA are utilized, the data processing time can be accelerated, the single image reconstruction time of the whole system is reduced, and the programmable logic array of the FPGA is utilized, so that the structure of a peripheral circuit can be reduced, and the system cost and the time cost of research, development and debugging are further reduced.
The above-mentioned embodiments are only used for illustrating the technical solutions of the present application, and not for limiting the same; although the present application has been described in detail with reference to the foregoing embodiments, it should be understood by those of ordinary skill in the art that: the technical solutions described in the foregoing embodiments may still be modified, or some technical features may be equivalently replaced; such modifications and substitutions do not depart from the spirit and scope of the embodiments of the present application, and they should be construed as being included in the present application.
Claims (7)
1. A magnetic nanoparticle imaging system, comprising:
a control module;
a gradient magnetic field generation module;
a driving magnetic field generating module; and
a signal receiving module;
the control module is electrically connected with the driving magnetic field generating module and the signal receiving module, the driving magnetic field generating module generates a variable driving magnetic field under the driving of the control module, the gradient magnetic field generating module is used for generating a gradient magnetic field, the gradient magnetic field and the driving magnetic field are superposed to form an imaging region, the imaging region is used for generating a zero magnetic field point and changing the position of the zero magnetic field point in the imaging region, the imaging region is used for arranging the magnetic nanoparticles, and the signal receiving module is used for receiving response magnetic field signals of the magnetic nanoparticles at different positions and outputting the response magnetic field signals to the control module for imaging;
the driving magnetic field generating module comprises a driving coil unit, and the signal receiving module comprises a receiving coil;
the driving coil unit comprises a first driving coil, the radius of the first driving coil is larger than that of the receiving coil, the receiving coil is coaxially sleeved in the first driving coil, and the imaging area is located in the receiving coil;
the driving coil unit further comprises a second driving coil, a third driving coil, a fourth driving coil and a fifth driving coil, wherein the axes of the second driving coil, the third driving coil, the fourth driving coil and the fifth driving coil are positioned on the same plane, and the plane is vertical to the axis of the first driving coil; the second driving coil and the third driving coil are coaxially arranged, the axis of the second driving coil is perpendicular to the axis of the first driving coil, the second driving coil and the third driving coil are respectively arranged on two sides of the first driving coil, and the second driving coil and the third driving coil are connected in series; the fourth driving coil and the fifth driving coil are coaxially arranged, the axis of the fourth driving coil is perpendicular to the axis of the first driving coil, the fourth driving coil and the fifth driving coil are respectively arranged on two sides of the first driving coil, and the fourth driving coil and the fifth driving coil are connected in series; the axis of the second driving coil is perpendicular to the axis of the fourth driving coil.
2. The magnetic nanoparticle imaging system according to claim 1, wherein the signal receiving module further comprises a first noise cancellation coil and a second noise cancellation coil for reducing noise, the first noise cancellation coil and the second noise cancellation coil have the same length, the first noise cancellation coil and the second noise cancellation coil have the same radius as the receiving coil, the first noise cancellation coil and the second noise cancellation coil are coaxially disposed with the receiving coil, the first noise cancellation coil and the second noise cancellation coil are respectively disposed on two sides of the receiving coil, the distance between the first noise cancellation coil and the receiving coil is the same as the distance between the second noise cancellation coil and the receiving coil, and the first noise cancellation coil and the second noise cancellation coil are sleeved in the first driving coil; and the first noise cancellation coil and the second noise cancellation coil are used for passing currents with equal magnitude and opposite directions.
3. The magnetic nanoparticle imaging system of claim 1, wherein the control module comprises an FPGA.
4. The magnetic nanoparticle imaging system of claim 3, wherein the control module comprises a drive module, the FPGA being electrically connected to the drive module; the driving module comprises a first driving signal generating unit, a second driving signal generating unit, a third driving signal generating unit and a digital-to-analog converter, wherein the first driving signal generating unit, the second driving signal generating unit and the third driving signal generating unit are electrically connected with the digital-to-analog converter, the digital-to-analog converter is provided with three driving signal output ends, the first driving signal output end is electrically connected with the first driving coil, the second driving signal output end is electrically connected with the second driving coil, and the third driving signal output end is electrically connected with the fourth driving coil; the digital-to-analog converter is used for performing digital-to-analog conversion on the first initial driving signal generated by the first driving signal generating unit to generate a first target driving signal, and then outputting the first target driving signal to the first driving coil through the first driving signal output end; the digital-to-analog converter is used for performing digital-to-analog conversion on the second initial driving signal generated by the second driving signal generating unit to generate a second target driving signal, and then outputting the second target driving signal to the second driving coil through the second driving signal output end; the digital-to-analog converter is used for performing digital-to-analog conversion on a third initial driving signal generated by the third driving signal generating unit to generate a third target driving signal, and then outputting the third target driving signal to the fourth driving coil through the third driving signal output end; the first target driving signal, the second target driving signal and the third target driving signal are alternating current driving signals with different frequencies.
5. The magnetic nanoparticle imaging system of claim 1, wherein the gradient magnetic field generating module comprises a pair of oppositely disposed permanent magnets of the same magnetic polarity that generate the gradient magnetic field.
6. The magnetic nanoparticle imaging system of claim 5, wherein the gradient magnetic field generation module further comprises a distance adjustment mechanism for adjusting the relative distance between the pair of permanent magnets.
7. The magnetic nanoparticle imaging system of claim 1, further comprising a host computer electrically connected to the control module.
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