Integration and Application Method of a Magnetic Plasma Sensing System

Abstract

This invention relates to the integration and application of a magnetoplasmic sensing system, belonging to the field of magnetoplasmic biomolecule sensing technology. A magnetoplasmic resonance refractive index sensor, composed of noble metals and magnetic materials, modulates the surface plasmon resonance (SPR) wave vector by applying an external AC electromagnetic field. Combined with dynamic differential phase-locked detection technology, characteristic signals are processed and extracted, significantly improving the stability and signal-to-noise ratio of the sensing system. In the signal detection stage, the transverse magneto-optical Kerr effect signal is collected instead of the traditional reflectivity detection mode, reducing the full width at half maximum (FWHM) of the signal peak by approximately 90% compared to the traditional SPR reflectance spectrum. Simultaneously, the electromagnetic field enhancement effect of localized surface plasmon resonance is introduced. The method of coupling metal nanoparticles with specific targets directly enhances the TMOKE response signal, enabling direct measurement of extremely low concentrations of biomolecules while maintaining a high quality factor, pushing the sensor's detection limit to the sub-picomolar concentration level.

Description

Technical Field

[0001] This invention belongs to the field of magnetic plasma biomolecular sensing technology, specifically involving the preparation of a magnetic plasma sensitive element of noble metal / magnetic multilayer thin film, the construction of a magneto-optical Kerr detection system and supporting data acquisition and processing software, and the use of nanoparticles with magnetic and optical signal responses coupled with biomolecular probes to achieve ultrasensitive sensing and detection, with the detection limit of biomolecules (such as antibodies, tumor markers, etc.) as low as the pM level. Technical Background

[0002] The magneto-optical (MO) effect is a typical physical phenomenon originating from the interaction between light and matter. Its characteristic feature is that when a medium is subjected to an external magnetic field, the propagation characteristics or reflection of incident electromagnetic waves within the medium change. Due to its unique physical mechanism, the magneto-optical effect has yielded numerous scientific achievements in various fields such as data storage, optical isolators, and biosensors. However, the development of the magneto-optical effect has been accompanied by and constrained by the inherently weak magneto-optical response of magnetic materials. For example, in in-plane magnetization-transverse magneto-optical Kerr effect (TMOKE) detection systems, the TMOKE signal is typically between 10⁻³ and 10⁻⁻⁴. 4 This poses a significant challenge to accurate detection. Since the magneto-optical effect essentially originates from the interaction between the electromagnetic field of the incident electromagnetic wave and the electrons inside the material, an effective way to enhance the magneto-optical properties of a material is to strengthen the collective electron resonance on the material surface, thereby achieving strong energy confinement near the interface—specifically, to excite surface plasmon resonance (SPR).

[0003] Surface plasmon resonance (SPR) sensors, as a mature detection technology in the field of optical detection, have become the preferred research method for characterizing the interaction mechanisms of biomolecules due to their advantages such as real-time dynamic monitoring, rapid response characteristics, and high sensitivity. However, studies have found that, limited by the inherent noise of the sensing system, the detection sensitivity of existing SPR sensors is still difficult to exceed 10. -6 The theoretical limit of RIU (Refractive Index Unit). It is worth noting that noble metal thin films, which are the core components of SPR sensors, inevitably have nanoscale surface roughness under conventional preparation processes. This defect will cause the full width at half maximum (FWHM) of the SPR characteristic peaks to be significantly broadened compared to the theoretical value, directly affecting the accuracy of characteristic peak position resolution.

[0004] One effective approach is to construct magnetoplasmic structures and modulate plasma dispersion by applying a magnetic field, resulting in a narrower full width at half maximum (FWHM) of the response signal compared to the reflection signal of traditional surface plasmon resonances, thereby enhancing its performance as a sensing signal. Therefore, sensors based on magneto-optical surface plasmon resonances (MOSPR) have become a focus of research in the field of biomolecular sensing. For example, Blanca Caballero et al. theoretically demonstrated that nanopore arrays integrated with magnetoplasmic structures composed of ferromagnetic and noble metals possess excellent sensing capabilities. Similarly, Li Lixia et al. developed a gold / cobalt bilayer nanoparticle array. Due to the high-quality surface plasmon resonance mode and modulation by an external magnetic field, they achieved a sub-nanometer-scale transverse magneto-optical Kerr effect signal bandwidth with a signal amplitude reaching 0.6 Å. Furthermore, localized surface plasmon resonances (LSPR) generated in metal nanostructures with dimensions much smaller than the wavelength can achieve up to 10 Å. 3 The field enhancement factor significantly enhances the interaction between light and matter. A prominent example is the application of metallic nanomaterials in surface-enhanced Raman scattering (SERS). The use of LSPR to enhance the MO effect has been predicted and observed in various nanostructure systems, such as magnetic nanoparticles, nanorods, and nanodisks. The coupling mechanism between LSPR and MO response has also been extensively studied and reported. Although magnetoplasmic structures exhibit excellent magneto-optical properties and have great potential in sensing applications, their practical application in biomolecular sensing research is relatively limited, with most studies remaining confined to theoretical simulations.

[0005] On the other hand, breakthroughs in the detection limit, a core parameter for measuring sensor performance, are often seen as a significant indicator of technological innovation—a lower detection limit means the instrument can detect trace amounts of analyte concentrations. However, it's worth noting that in practical applications such as clinical diagnostics, improving the sensor's effective operation within the normal physiological concentration range is often more meaningful than pursuing detection limits far below physiological concentrations. For example, in blood glucose monitoring, accurate detection within the physiological concentration range of 3.9-6.1 mmol / L is far more clinically valuable than its ability to detect trace levels as low as 0.1 mmol / L. Therefore, it is necessary to establish a multi-dimensional performance evaluation system when developing biosensors: while pursuing lower detection limits, it is crucial to systematically consider the synergistic optimization of key parameters such as sensitivity, signal-to-noise ratio, and quality factor. Especially in areas such as cost control, operational suitability, and feasibility of mass production, a research and development strategy that prioritizes both basic research and clinical application is essential.

[0006] Existing technologies present significant limitations in traditional surface plasmon resonance sensors, which struggle to simultaneously achieve high sensitivity, high quality factor, and low detection limit.

[0007] 1. Sensing elements made of a single thin film material (such as noble metals Au and Ag) have limited signal broadening (FWHM > 5nm, FWHM > 2°) and quality factor (FOM < 50 RIU⁻¹) due to unavoidable nanoscale surface roughness during the fabrication process;

[0008] 2. The reflectivity detection mode is susceptible to interference from light source fluctuations and interface thermal noise, resulting in a low signal-to-noise ratio and stability in actual tests.

[0009] 3. The electromagnetic field localization enhancement effect is insufficient when detecting low concentrations of biomolecules, which makes it impossible to detect concentrations of picomoles and below. Summary of the Invention

[0010] To achieve synergistic optimization of multiple sensor parameters while maintaining the detection limit at low concentrations, this invention proposes a magnetoplasmic resonant refractive index sensor composed of noble metals and magnetic materials. It modulates the surface plasmon resonance (SPR) wave vector by applying an external AC electromagnetic field, and combines this with dynamic differential phase-locked loop (PLL) detection technology to process and extract characteristic signals, significantly improving the stability and signal-to-noise ratio of the sensing system. In the signal detection stage, the transverse magneto-optical Kerr effect signal is collected instead of the traditional reflectivity detection mode, reducing the full width at half maximum (FWHM) of the signal peak by approximately 90% compared to the traditional SPR reflectance spectrum. This provides a physical basis for improving the sensor's sensitivity and quality factor. Furthermore, the electromagnetic field enhancement effect of localized surface plasmon resonance is introduced, and the TMOKE response signal is directly enhanced by coupling metal nanoparticles to specific targets. This enables the direct measurement of extremely low concentrations of biomolecules while maintaining a high quality factor, pushing the sensor's detection limit to the sub-picomolar level.

[0011] An integrated magnetic plasma sensing and detection system comprises core components including a right-angle prism with a magnetic plasma sensing element (i.e., a multilayer heterojunction thin film), an electromagnet, a silicon photodiode detector (FD11A, Thorlabs) (corresponding to photodetector 1 and photodetector 2 respectively), a laser, a polarizer (including a polarizing prism (GTH10M, Thorlabs) and an aperture), a 5:5 beam splitter, a lens, a lock-in amplifier (SR830), a data acquisition card (USB-6001 NI), a signal generator, the system's mechanical rotation, and a data acquisition system (computer).

[0012] System optical path such as Figure 2As shown, the output beam of the 632.8 nm He-Ne laser is first modulated into TM polarized light using a polarizer, and then split into two beams using a 5:5 beam splitter. One beam serves as the incident reference light and is collected by lens 1 and photodetector 1. The resulting signal is then transmitted to a data acquisition card for intensity normalization to reduce system noise. The other beam from the 5:5 beam splitter is incident on one right-angled surface of a magnetoplasm sensing element prism. The light refracted from the other right-angled surface of the prism is collected by lens 2 and photodetector 2. The resulting signal is then used as the input signal of a lock-in amplifier, serving as the response signal for refractive index sensing. The output of the lock-in amplifier is connected to the data acquisition card, which is connected to a data acquisition system (computer). The mechanical rotation of the system is used to rotate the prism, ensuring that lens 2 and photodetector 2 rotate simultaneously. The device 2 also rotates, causing the light refracted from the other right-angled surface of the plasma sensing element prism to enter the photodetector 2 through the lens 2 for testing. The modulation of the external magnetic field of the magnetic plasma sensing element is provided by a vertically placed AC electromagnet, and the prism with the sensing element is placed in the magnetic field. The function of the signal generator is to output two AC signals of the same frequency: one signal is used by the electromagnet to control the frequency of the external magnetic field of the sensing element, and the other signal is used as the reference input of the lock-in amplifier. The light reflected by the sensing element, i.e., the light refracted from the other right-angled surface of the magnetic plasma sensing element prism, is collected by the lens 2 and the detector 2 and input into the lock-in amplifier for mixing and filtering. The output electrical signal after processing is the TMOKE sensing response signal, which is finally transmitted to the data acquisition system (computer) by the data acquisition card. Figure 4 A physical image of the magnetic plasma sensor detection system is shown.

[0013] The mechanical rotation of the system, such as Figure 3 As shown, the mechanical rotation of the system is used to rotate the prism, including two turntables: coaxial turntable 1 and turntable 2. Turntable 1 fixes the prism, and turntable 2 fixes the detector 2 and lens 2 via an outwardly extending bracket. Turntable 1 and turntable 2 are driven to rotate by stepper motors to ensure that the rotation of turntable 2 when the prism rotates with turntable 1 allows lens 2 and photodetector 2 to receive refracted light. Both turntable 1 and turntable 2 are driven by stepper motors. The system is realized by a θ–2θ coaxial dual rotary stage system driven by stepper motors. According to the law of reflection, the ratio of the change in the incident angle to the change in the reflection angle is always 1:2. Therefore, the ratio of the angular velocity of turntable 1 where the prism is located to the angular velocity of turntable 2 where the detector 2 is located should also be 1:2, so as to realize the measurement of light intensity over the entire range of angles. θ–2θ represents the ratio of the angular velocity of turntable 1 where the prism is located to the angular velocity of turntable 2 where the detector 2 is located, which should also be 1:2.

[0014] The magnetoplasma sensing element is fabricated on the inclined surface of a right-angled triangular prism using physical vapor deposition. A multi-channel PDMS (polydimethylsiloxane) microfluidic chip is attached to the surface of the magnetoplasma sensing element. The multi-channel PDMS microfluidic chip is used to introduce the liquid to be measured onto the surface of the sensor sensing element, such as... Figure 1 As shown, this enables the sensor to perform liquid phase sensing detection; the stepper motor is used to control the rotation of the turntable, thereby driving the right-angle prism to rotate. The stepper motor is controlled by a computer.

[0015] Fabrication of magnetic plasma multilayer heterojunction thin films:

[0016] A magnetic plasma sensing element is fabricated on the inclined surface of a right-angle prism using physical vapor deposition. The magnetic plasma sensing element is a multilayer heterojunction thin film, primarily composed of noble metal materials (such as Au, Ag, Cu, Al, etc.), magnetic materials (ferromagnetic materials such as Fe, Co, Ni, Mo, and their-based metal alloys and intermetallic compounds CoFeB), and garnet ferrite materials (including rare-earth doped and / or non-rare-earth doped materials such as R3Fe5O). 12The rare earth ions R are Y, Gd, Dy, Bi, etc.) and several combinations of strong spin-orbit coupling materials (W, Ta, Pt, Bi, Pb and actinides and their alloys) form a multilayer heterojunction thin film, and the outer surface of the magnetic plasma sensing element is a noble metal thin film material. The specific types can be divided into the following categories: The first category is a sandwich structure composed of noble metal thin film material / ferromagnetic thin film material / noble metal thin film material, such as Au / Co / Au, Ag / CoFeB / Ag, etc. In this type of structure, ferromagnetic material is embedded in noble metal material as the interlayer; The second category is a bilayer heterostructure composed of noble metal thin film material / garnet ferrite thin film material, such as Au (top) / YIG (bottom). In this type of structure, garnet ferrite thin film is used as magnetic material, and noble metal film is covered on the top layer for biofunctionalization; The third category is a trilayer heterostructure composed of noble metal thin film material / rare earth doped garnet ferrite thin film material / garnet ferrite thin film, such as Ag (top) / Ce:YIG / YIG (bottom). This type of structure introduces a magnetic layer with stronger magneto-optical effect. However, because the lattice difference between the doped garnet ferrite and the substrate is too large and it is not easy to crystallize, a layer of YIG is usually deposited in advance as a seed layer during preparation. The fourth type is a three-layer heterostructure composed of noble metal thin film materials, ferromagnetic thin film materials, and strongly spin-orbit coupled thin films. This type of structure introduces a strongly spin-orbit coupled layer to enhance the magneto-optical effect of the ferromagnetic layer. The thickness of each material in the multilayer heterostructure plays a crucial role in the overall performance of the device. The suitable thickness of the noble metal layer is 40-60 nm, mainly used to excite surface plasmon resonance. The thickness of the magnetic thin film material should not be too large, as the high optical loss of this material will affect the performance of the sensor; therefore, the suitable range is 5-20 nm. Garnet ferrite and rare-earth-doped garnet ferrite films have weaker magnetism compared to ferromagnetic materials, but extremely low optical loss; therefore, the suitable thickness range is 50-200 nm. The thickness of the strongly spin-orbit coupled thin film material is 10-20 nm.

[0017] To enhance the adhesion between the heterojunction thin film and the substrate, a Cr or Ti film with a thickness of 2 nm can generally be sputtered onto the substrate first. For noble metal thin film materials and some magnetic thin film materials, direct magnetron sputtering is sufficient. Garnet ferrite thin films need to be annealed to allow them to crystallize and exhibit magnetism.

[0018] All of the above types of multilayer heterojunction thin films can achieve strong coupling between surface plasmon polaritons (SPP) and magneto-plasmons, as well as highly sensitive sensing and detection. Figure 5 Schematic diagrams of four types of multilayer heterojunction thin films are presented.

[0019] Biofunctionalization of the sensing element surface: The purpose of biofunctionalization is to modify the metal surface of the sensor sensing element to enable it to bind detectable molecules more effectively. This mainly involves three steps: First, a self-assembled monolayer is formed on the noble metal surface using mercaptoalkanoic acids (e.g., HS-(CH2)). 11 A self-assembled monolayer is constructed by mixing carboxyl groups (-COOH) and mercaptools (e.g., HS-(CH2)6-OH) in a 1:9 molar ratio, achieving precise control over carboxyl group distribution density and steric hindrance. Next, the self-assembled monolayer is activated using a 4:1 molar ratio EDC / NHS mixture to effectively link biomolecules such as proteins, for example, through localized transport of the EDC / NHS mixture via a microfluidic chip. Finally, analyte recognition units (e.g., IgG) are coupled to the surface of the activated self-assembled monolayer. Furthermore, to enhance the detection limit of the magnetic plasma sensor, biomolecular probes are constructed by coupling analyte biomolecules with noble metal (Au, Ag, Cu, Pt) nanoparticles. The local electromagnetic field enhancement effect of these probes amplifies the sensing response changes caused by differences in biomolecule concentration.

[0020] By setting up the entire system, the usage method of a magnetic plasma sensing system includes the following steps:

[0021] (1) When characterizing the performance of the magnetic plasma sensing system (the sensing element does not need to be biofunctionalized), a characterization solution with a certain refractive index (such as an ethanol solution, preferably with a refractive index not much different from that of the actual substance to be tested) is first introduced. At this time, the signal generator outputs an AC signal of a certain frequency (greater than 50Hz and less than 200Hz) to control the direction of the electromagnetic field (reversal of N and S). At the same time, the direction of the electromagnetic field is parallel to the plane of the sensing element, and the magnetic field strength satisfies the saturation magnetization of the magnetic layer (generally 50mT is sufficient in this invention). The test adopts the angle-TMOKE mode, that is, the turntable is driven by a stepper motor to change the incident angle of the laser to the sensing element (for example, the incident angle is rotated from 40° to 50° with a step size of 0.01°). At each angle of the turntable rotation, the detector 2 records the real-time light intensity and inputs it to the lock-in amplifier for processing to obtain the TMOKE response. After the light intensity signal obtained by the detector 1 is normalized, the relationship between the angle and the TMOKE spectrum can be obtained, that is, the angle-TMOKE spectrum. The optimal incident angle can be obtained from it. Generally, the angle corresponding to the largest slope in the angle-TMOKE spectrum is selected.

[0022] (2) Real-time biological detection adopts the fixed incident angle, i.e. optimal incident angle detection mode. At this time, the stepper motor is not used to drive the turntable to change the incident angle, and the biofunctional sensitive element is replaced. During the test, the electromagnet is also controlled by the AC signal (greater than 50Hz and less than 200Hz) output by the signal generator to control the magnetic field direction. The magnetic field strength is the same as in step (1). Then, a certain concentration of the biomolecule solution to be detected is introduced, and then the detector 2 records the light intensity in real time and inputs it to the lock-in amplifier for processing to obtain the change of TMOKE response over time, i.e. finally obtaining the time-TMOKE spectrum.

[0023] This invention can be applied to biomolecular detection. Furthermore, it can use gold nanoparticles to couple the target biomolecule (e.g., IgG antibody) for electromagnetic field localization enhancement, thereby further improving the sensing response change caused by biomolecule concentration differences and reducing the detection limit. Attached Figure Description

[0024] Figure 1 shows the bonding of the PDMS microfluidic chip of the present invention with a prism having a heterojunction thin film. Figure 2 is a schematic diagram of the magnetic plasma sensor detection system. Figure 3 is a schematic diagram of the θ–2θ coaxial dual rotary table system. Figure 4 Image of a magnetic plasma sensor detection system Figure 5 Schematic diagrams of four types of multilayer heterojunction thin films are given. Figure 6 is a schematic diagram of the sensitive element structure of the magnetic plasma sensor and its specific detection of IgG antibodies. Figure 7 shows the characterization of the sensor's refractive index sensitivity (specific steps are described in Example 1). Figure 8 XRD characterization images of garnet films annealed at different temperatures Figure 9. Localized plasmon resonance coupled magnetoplasmic sensing detection. Figure 10 shows the sensing and detection of antibodies conjugated with nanoparticles of different sizes. Detailed Implementation

[0025] This invention proposes a method for constructing a magnetic plasma biosensor based on a noble metal / magnetic heterostructure thin film, wherein the specific process for thin film preparation is as follows:

[0026] Different materials used in the sensing experiments were prepared using different methods for the heterostructure thin films. Noble metal and magnetic thin films were deposited at room temperature using physical vapor deposition (RF / DC magnetron sputtering and vacuum thermal evaporation) on substrates of different materials and refractive indices. Before coating, the substrate (prism) needed to be cleaned to remove surface impurities. This involved ultrasonic cleaning with acetone (40kHz, 100W) for 10 minutes, followed by rinsing with deionized water, and then immersion in a piranha solution (98% H2SO4 and 30% H2O2 in a 7:3 ratio) at room temperature for 1 hour. After removal, it was rinsed with deionized water and finally dried with nitrogen. To enhance the adhesion between the thin film and the substrate, a 2nm thick Cr or Ti film could be selectively sputtered onto the substrate first. Alternatively, noble metal and some magnetic material thin films (e.g., Au, Ag, Fe, Co) could be deposited using DC magnetron sputtering; specific coating parameters are given in the examples. Ferromagnetic alloy and garnet ferrite thin films (e.g., FeNi, CoFeB, Y3Fe5O) were also prepared. 12 Radio frequency magnetron sputtering is used. The specific parameters for the coating are given in the examples. No heating is performed after coating of noble metal and magnetic thin film materials, while heating is performed after coating of garnet ferrite. After deposition, the film thickness can be tested by profilometer, ellipsometer, etc. The film deposition quality is observed by atomic force microscope (AFM) and scanning electron microscope (SEM). Finally, the sensitive element is placed in an argon atmosphere for storage and later use.

[0027] The detection principle of the sensor signal TMOKE includes the following:

[0028] (1) The core of the dual-channel lock-in amplifier used is a multiplier mixer, which can multiply and mix the optical signal received by detector 2 and the reference signal of the same frequency input by the signal generator. The processed signal is then output as a DC signal after passing through the low-pass filter module built into the lock-in amplifier. The sensor response TMOKE signal is transmitted to the computer via the data acquisition card. The beam after the 5:5 split is used as the incident reference light and is collected by lens 1 and photodetector 1. The obtained signal is then transmitted to the data acquisition card and then to the computer for normalization of light intensity to reduce the influence of system noise.

[0029] (2) The real-time response of the sensing system is used for data collection and processing, which is mainly divided into two test modes: angle-TMOKE mode and time-TMOKE mode.

[0030] Detailed operation steps for Angle-TMOKE mode:

[0031] (1) Place the right-angle prism (with a magnetic plasma multilayer heterojunction thin film and PDMS microfluidic chip attached) on the turntable 1 of the coaxial system and pump in a solution with a certain refractive index (e.g., water or ethanol solution).

[0032] (2) Manually adjust the laser so that it enters the prism from the right-angled face of the prism and enters the inclined face of the prism, i.e. the multilayer heterojunction film, and then exits from the right-angled face of the other side of the prism and is received by the detector 2.

[0033] (3) Set the signal generator to generate two 50Hz AC signals. One signal is used to control the direction of the magnetic field of the electromagnet. The magnetic field strength is set to 50mT. At this time, the prism is in the vertical magnetic field. The other signal is used as the reference signal input of the lock-in amplifier.

[0034] (4) The computer issues instructions to control the stepper motor to drive the turntable to rotate in steps of 0.01°. The total rotation angle depends on the specific situation. At the same time, the output response (TMOKE) of the lock-in amplifier is recorded for each rotation, thus obtaining the angle-TMOKE spectrum.

[0035] Angle-TMOKE mode measurements can reflect the basic sensing performance of a sensitive element without changing the refractive index environment, such as the full width at half maximum (FWHM), peak intensity, and angular offset of the response signal. These are important reference indicators in the process of sensor structure optimization. Simultaneously, this measurement mode can also be used to find the most sensitive incident angle of the device for subsequent real-time sensing and detection at a fixed angle.

[0036] The specific steps for using the Time-TMOKE mode are as follows:

[0037] (1) Similar to step 1 of the angle-TMOKE mode, the pumped liquid is water or buffer solution to obtain the signal baseline when there is no analyte;

[0038] (2) Same as step 2 of the angle-TMOKE mode, but in this mode the incident angle is fixed;

[0039] (3) Same as step 3 of the angle-TMOKE mode;

[0040] (4) The computer issues an instruction to start recording the output response of the lock-in amplifier at each time step (e.g., 1 second), and then pumps the solution containing the biomolecule to be detected into the microfluidic chip. When the biomolecule binds to the functional groups or biorecognition units on the biofunctionalized film, the response signal will change, thereby obtaining the time-TMOKE spectrum.

[0041] The time-reflectivity mode can detect changes in response signals in environments with different refractive indices in real time, and is also an essential means of sensing and detecting low concentrations of biomolecules.

[0042] Biofunctionalization of sensors and the chemical reagents used:

[0043] The buffer solutions used for functionalization of the sensitive element surface and low-concentration biomolecule detection experiments were all 0.01M phosphate-buffered saline (PBS, pH 7.4) diluted with deionized water. Other biochemical materials used in this invention include: anhydrous ethanol, bovine serum albumin (BSA), human immunoglobulin (IgG), goat anti-human immunoglobulin antibody (Goat Anti-Human IgG), N-hydroxysuccinimide (NHS), 1-ethyl-3-(dimethylaminopropyl)carbodiimide (EDC), 6-mercaptohexanol (6-MCH), 11-mercaptoundecanoic acid (11-MUA), colloidal gold-labeled goat anti-human IgG, lipoic acid-PEG-Biotin (LA-PEG-Biotin), and streptavidin. The core objectives of surface functionalization are as follows: First, it aims to achieve the directional immobilization of probe molecules, mainly through chemical bonding to stably anchor biomolecules (such as antigens and antibodies), ensuring the surface density of their active sites and improving the binding efficiency of biomolecules. Second, it aims to reduce non-specific adsorption, for example, by using mixed self-assembled monolayers (such as mercaptoalkyl acids and mercaptohexanol) to construct bioinert interfaces, where short-chain molecules occupy vacancy defects, reducing interference from non-specific signals. Finally, it aims to utilize the local electromagnetic field enhancement effect of interface-modified nanoparticles or biomolecule-coupled nanoparticles to further reduce the detection limit.

[0044] Example 1

[0045] This case study uses a sandwich structure (Ag / Co / Ag) composed of a noble metal / ferromagnetic material as the sensor sensing element, and uses ethanol solutions of different concentrations as the test solutions to illustrate the refractive index sensitivity and other basic sensing performance of the sensor.

[0046] Step 1: Fabrication of the sensor chip

[0047] First, a ZF13 prism was selected as the substrate for the coating, with a refractive index of 1.78 at a wavelength of 632.8 nm. According to the SPP wave vector matching condition, a higher ambient refractive index results in a larger resonant excitation angle; therefore, selecting a high-refractive-index prism as the substrate is beneficial for subsequent biosensing tests in solution. After cleaning the substrate, a 27 nm thick Ag film was deposited using DC magnetron sputtering with the following parameters: background vacuum 8 × 10⁻⁶. -4 The deposition rate was 9 nm / min, obtained from AFM measurement of the thin film step thickness, with an argon working pressure of 1 Pa, an argon flow rate of 80 Sccm, and a DC sputtering power of 10 W. The intermediate magnetic layer Co was deposited using RF magnetron sputtering with a thickness of 8 nm. The deposition parameters were: base vacuum of 8 × 10⁻⁶. -4With an argon working pressure of 0.1 Pa, an argon flow rate of 20 Sccm, and a sputtering power of 100 W, the deposition rate was 6 nm / min as determined by AFM measurement of the film step thickness. The top Ag film was deposited with the same parameters to a thickness of 10 nm. After the film preparation was completed, a self-made polydimethylsiloxane (PDMS) microfluidic chip was attached to the metal film side of the prism.

[0048] Step 2: Preparation of calibration solution and control experimental group

[0049] The refractive index of the solution exhibits an approximately linear relationship with its concentration. Ethanol-water mixed solutions of different volume concentrations (0%, 10%, 20%, 30%, and 40%) were prepared using deionized water. The refractive indices of these solutions were 1.333, 1.338, 1.342, 1.346, and 1.351 (measured using an Abbe refractometer), serving as standard calibration media for evaluating the sensor's refractive index sensitivity. Different concentrations of solution were injected onto the outside of the sensing element using a peristaltic pump to alter the ambient refractive index of the chip surface. For performance comparison, a control experiment was set up, with a 45nm monolayer Ag film deposited on an SPR sensor chip under identical conditions. Initially, the reflectivity and TMOKE curves of the two chips were measured in air. The results showed that the SPR reflectivity of the magnetoplasmic sensor chip in step 1 was significantly narrower (approximately 0.1°) than that of the control experiment, indicating higher resolution.

[0050] Step 3: Determination of refractive index sensitivity and quality factor

[0051] Refractive index sensitivity is an important parameter for evaluating sensor performance; it is defined as... ,in This represents the resonance peak displacement. This represents the change in refractive index. An angle-TMOKE detection mode was used (the turntable rotated in 0.01° steps, the incident angle from 48° to 59°, the AC magnetic field frequency was 50Hz, and the magnetic field strength was 50mT). The angle-TMOKE spectra corresponding to each concentration of ethanol solution were measured, as shown below. Figure 6 As shown in (a), a linear fit was then performed on the resonance points corresponding to each refractive index, yielding a sensitivity of 155.47° / RIU (per unit refractive index) for the magnetic plasma sensor. The reflectance curve of the 45 nm Ag film SPR sensor corresponding to the control experiment was tested under the same conditions, as shown in (a). Figure 6 (b) The sensitivity after linear fitting is 137.75° / RIU, which is due to the additional wave vector modulation provided by the magnetic plasma structure, which can significantly improve the surface sensitivity of the sensor.

[0052] Furthermore, using the time-TMOKE mode, the real-time sensing response curves of ethanol of different concentrations were measured under a fixed angle (incident angle of 52.6°, AC magnetic field frequency of 50Hz, and magnetic field strength of 50mT), and the signal-to-noise ratio of the device at the highest concentration exceeded 2000.

[0053] The magnitude of the resonance peak displacement and the resonance linewidth fundamentally affect the accuracy of resonance peak tracking. Therefore, by defining a quality factor: To evaluate the overall performance of the sensors, Table 1 presents detailed performance parameters of the two sensors at different solution concentrations. The magnetic plasma sensor exhibits a high FOM range of 400.11-631.47 in solution, compared to only 47.44-48.36 for the traditional SPR sensor, representing an order-of-magnitude improvement. The comparative experiments clearly demonstrate the significant advantages of the magnetic plasma sensor in improving sensing performance.

[0054] Table 1 shows the significant advantages that this sensor can achieve through the synergy of the above technologies:

[0055] Half-width at half-height (FWHM) 0.24° 2.85° Reduced by 91.2% Refractive index sensitivity (S) 155.47° 137.75° An increase of 12.9% Quality Factor (FOM) <![CDATA[631.47RIU -1 ]]> <![CDATA[48.2 RIU -1 ]]> Increased by approximately 12 times Limit of Detection (LOD) 5.2pM 0.1nM Reduced by about 20 times

[0056] Table 2 Performance of magnetoplasm and SPR sensors at different refractive indices

[0057] 1.333 0.3413 455.52 2.8487 48.36 1.338 0.3248 478.66 2.9035 47.44 1.342 0.3884 400.11 2.8842 47.76 1.346 0.2462 631.47 2.8538 48.26 1.351 0.3781 411.29 2.8992 47.51

[0058] Example 2

[0059] This case study uses a double-layer heterojunction (Au / YIG) composed of precious metal / garnet ferrite material as the sensor sensing element, and uses BSA as the analyte to verify the sensor's detection performance for biomolecules.

[0060] Step 1: Fabrication of the sensor chip

[0061] Since garnet ferrite YIG films only exhibit magneto-optical properties in their polycrystalline state after high-temperature annealing, unlike Example 1, high-temperature resistant sapphire (Al2O3) was chosen as the coating substrate (i.e., the corresponding right-angle prism), with a refractive index of 1.76 at 632.8 nm. A YIG film with a thickness of approximately 100 nm was deposited using radio frequency magnetron sputtering, with the following coating parameters: base vacuum 8 × 10⁻⁶. -4The parameters were: Pa, argon working pressure 1 Pa, argon flow rate 40 Sccm, and sputtering power 80 W. Annealing after coating is the most critical step, determining the crystallization effect and magneto-optical properties of the film. In this case, rapid annealing was used, annealing the prepared amorphous YIG thin film in air at different temperatures (room temperature, 760℃, 860℃, and 960℃) for 10 min each, followed by XRD and magneto-optical characterization. Figure 7 To characterize the results, a YIG thin film annealed at 860℃ was ultimately selected for subsequent sensing experiments. The top layer of noble metal Au was deposited using DC magnetron sputtering with a thickness of approximately 45 nm, and the deposition parameters were the same as those for the Ag film in Example 1. Finally, the self-made PDMS microfluidic chip was attached to the metal thin film side of the substrate, thus completing the fabrication of the sensing chip.

[0062] Step 2: Biofunctionalization of the sensor chip

[0063] First, the sensor chip prepared in step 1 was immersed in an ethanol mixture containing 20 mM MUA and 180 mM MCH to form a hybrid self-assembled monolayer (SAM) on the Au membrane surface. To maximize the sensitivity of biomolecule detection, the molar ratio of the two thiol molecules was set to 1:9. The long-chain MUA molecules covalently bind to the Au membrane surface through Au-S bonds, and the exposed carboxyl groups can be used for subsequent activation and protein molecule binding. Simultaneously, MCH effectively passivates the Au membrane, reducing the adsorption of non-specific proteins and improving the selectivity of biomolecule recognition. Subsequently, a prepared EDC (200 mM) / NHS (50 mM) mixture was injected at a flow rate of 100 μL / min using a microfluidic pump to activate the carboxyl groups, enabling them to covalently couple with amine-containing biomolecules. Finally, the chip was thoroughly rinsed with PBS to complete the biofunctionalization process.

[0064] Step 3: Real-time dynamic detection of biomolecules

[0065] Following the aforementioned biofunctionalization process, -COOH dangling bonds are formed on the Au membrane surface. To verify the sensor's detection performance for biomolecules, a 10 μg / ml BSA solution was prepared and injected into the sensor chip via a microfluidic pump. The TMOKE signal changes were monitored in real time using time-TMOKE mode (incident angle 52.6°, AC magnetic field frequency 50 Hz, magnetic field strength 50 mT). To compare the effect of interfacial biofunctionalization, a control experiment was set up to test the signal changes of the device without surface modification. The results show that the surface-biofunctionalized sensor chip can achieve real-time detection of biomolecules at the nM level, with a significantly improved signal change at the same concentration.

[0066] Example 3

[0067] This case study utilizes a sandwich structure (Ag / CoFeB / Ag) composed of a noble metal / ferromagnetic alloy as the sensor's sensitive element, and IgG and IgG antibodies as targets for specific biomolecular detection to verify the effectiveness of biomolecular hybridization assays. To further improve the detection limit, Au nanoparticles (Au-NPs) were used to construct a localized plasmonic biomolecular probe for the dynamic detection of Au-NPs-conjugated IgG antibodies. Simultaneously, Au-NPs-conjugated antibodies at different scales (0nm, 10nm, 15nm, 20nm) were set up as experimental control groups to explore the effect of their localized electromagnetic field enhancement on improving detection performance.

[0068] Step 1: Sensor chip fabrication

[0069] Ag / CoFeB / Ag was sequentially deposited on a ZF13 prism. The deposition parameters for the top and bottom Ag films were the same as in Example 1, with deposition thicknesses of 27 nm and 10 nm, respectively. The magnetic layer was a ferromagnetic alloy, CoFeB, deposited using RF magnetron sputtering, with a thickness of 8 nm. The deposition parameters were: base vacuum 8 × 10⁻⁶. -4 The parameters were: Pa, argon working pressure 0.1 Pa, argon flow rate 40 Sccm, and sputtering power 100 W. Finally, the self-made PDMS microchannel chip was attached to the metal thin film side of the substrate.

[0070] Step 2: Biofunctionalization of the sensor chip

[0071] To anchor the IgG protein recognition unit on the chip, the method used was the same as step 2 in Example 2. First, a mixed self-assembled monolayer of MUA and MCH was modified on the chip surface, and the surface carboxyl groups were activated. Then, an IgG solution with a concentration of 100 ug / ml was passed through at a flow rate of 50 μL / min to complete the biofunctionalization of the chip surface. The response signal changes throughout the process were recorded using time-TMOKE mode (incident angle of 52.6°, AC magnetic field frequency of 50 Hz, magnetic field strength of 50 mT). Figure 8 As shown in (a).

[0072] Step 3: Antigen-antibody specific binding detection

[0073] The IgG biorecognition unit can selectively bind to corresponding IgG antibodies, causing a localized change in the refractive index of the chip surface. This highly sensitive optical near-field perturbation significantly enhances the TMOKE sensing response. IgG antibody solutions with concentrations of 1 μg / ml, 2 μg / ml, 5 μg / ml, 10 μg / ml, 20 μg / ml, and 50 μg / ml were prepared and introduced into the sensor at a flow rate of 50 μL / min. The hybridization process of surface-immobilized IgG and IgG antibodies was monitored in real time using a time-TMOKE mode (incident angle 52.6°, AC magnetic field frequency 50 Hz, magnetic field strength 50 mT). Figure 8 As shown in (b), starting from the baseline, the TMOKE response gradually increases with increasing antibody concentration, reflecting an increase in the binding rate. At higher concentrations, the binding rate begins to slow down, indicating that the binding of surface biomolecules gradually approaches saturation. To quantitatively analyze molecular affinity, the Langmuir isotherm adsorption equation was used for fitting: ,in antibody concentration TMOKE response at that time This represents the maximum response when the surface binding sites are fully saturated. The binding constant represents the intermolecular affinity. The fitting results are as follows: Figure 8 As shown in (c), the affinity constant of the obtained magnetic plasma sensor The corresponding affinity range is approximately 6.87 × 10⁻⁶. 6 ~ 1.2 × 10 7 M⁻¹. This value falls within the medium-to-high binding affinity range, indicating that the developed sensor has strong molecular recognition capabilities.

[0074] Step 4: Localized plasmon resonance coupled magnetoplasmic sensing and detection

[0075] To further improve the detection limit of the magnetic plasma sensor, a sandwich-type sensing strategy was employed. This strategy utilizes IgG antibodies coupled with AuNPs (i.e., AuNPs conjugated to the test antibody) to specifically bind to IgG recognition units immobilized on the surface. The AuNPs were 10, 20, and 30 nm in size; a control group consisting of IgG antibodies without AuNP conjugation was included. An antibody solution with a concentration of 50 ng / mL was introduced via a microfluidic pump, and changes in the sensing curve were monitored in real time. Figure 9 As shown in (a), the response change of the 20nm AuNPs-conjugated antibody was the largest, while the signal change was almost unchanged when testing the unconjugated IgG antibody. Figure 9 (b) and (c) demonstrate the local electromagnetic field enhancement effect simulated using COMSOL software. Figure 9(d) is a SEM image of the AuNPs conjugated with the antibody. Based on the three-standard-deviation theory, the detection limit of the device was calculated to be 0.778 ng / mL (5.2 pM). Furthermore, biomolecular probes were constructed using Ag, Cu, and Pt nanoparticles, and the detection limits for IgG antibodies obtained were all below 1 ng / mL.

[0076] Example 4

[0077] This case study utilizes a three-layer heterostructure (Au / CeYIG / YIG) composed of precious metal / rare earth doped garnet ferrite materials as the sensor sensing element and alpha-fetoprotein (AFP) as a tumor marker to achieve clinical diagnosis of liver cancer.

[0078] Step 1: Fabrication of the sensor chip

[0079] Compared to YIG thin films, Ce:YIG thin films exhibit enhanced magneto-optical effects due to the strong spin-orbit coupling effect of Ce, an element with a larger atomic number. However, the substitution of Y ions by Ce ions increases the lattice constant of the crystal. Therefore, a YIG seed layer is deposited first to promote the crystallinity of the Ce:YIG layer and improve the non-reciprocal phase shift. First, a 50 nm YIG layer is deposited on an Al2O3 prism, with the same deposition and annealing parameters as in step 1 of Case 2. Then, a 100 nm Ce:YIG layer is deposited on the crystallized YIG seed layer using RF magnetron sputtering, with the following deposition parameters: base vacuum 8 × 10⁻⁶. -4 The magnetic layer was fabricated using an argon working pressure of 1 Pa, an argon flow rate of 200 Sccm, and a sputtering power of 80 W. After annealing in air at 860℃ for 10 min, the entire magnetic layer was completed. Finally, a 45 nm thick Au layer was deposited on top, with the same deposition parameters as in Case 1, and the PDMS microfluidic chip was encapsulated to complete the chip fabrication.

[0080] Step 2: Biofunctionalization of the sensor chip

[0081] Using the same method as step 2 in Example 2, alpha-fetoprotein antibodies were anchored on the chip surface for subsequent specific detection of alpha-fetoprotein.

[0082] Step 3: Real-time detection of liver tumor markers

[0083] Alpha-fetoprotein (AFP) solutions with concentrations of 0.1 μg / ml, 0.5 μg / ml, 1 μg / ml, 5 μg / ml, and 10 μg / ml were prepared and introduced into the sensor at a flow rate of 50 μL / min. The bonding process between the surface-immobilized AFP antibody and AFP was monitored in real time using a time-TMOKE mode (incident angle of 52.6°, AC magnetic field frequency of 50 Hz, and magnetic field strength of 50 mT). The real-time sensing curve was similar to that in step 3 of Case 3. Using the three-standard-deviation rule, the detection limit of the sensor was defined, ultimately obtaining a detection limit of 2 ng / ml for liver tumor markers, far below the normal range of human AFP. Based on this, Ag / Bi:YIG / YIG and Au / Gd:YIG / YIG sensor chips were also prepared for AFP detection. The preparation conditions of Bi:YIG and Gd:YIG were the same as those of Ce:YIG, and the detection limits obtained were all below 5 ng / mL.

[0084] Example 5

[0085] This case uses a four-layer heterostructure (Ag / CoFeB / W / Ag) composed of noble metal / ferromagnetic material / strong spin-orbit coupling layer / noble metal as the sensor sensing element, and the sandwich structure (Ag / CoFeB / Ag) in Example 3 as a control experiment. IgG and IgG antibodies are used as targets for biomolecular specific detection to verify the improvement of sensor performance by the introduction of the strong spin-orbit coupling layer.

[0086] Step 1: Fabrication of the sensor chip

[0087] The preparation conditions and thickness of the noble metal Ag were the same as in Example 3, and the coating parameters of the ferromagnetic material Fe were: base vacuum degree 8×10 -4 The parameters were: Pa, argon working pressure 0.5 Pa, argon flow rate 40 sccm, and DC sputtering power 15 W. The strong spin-orbit coupling material W was deposited using DC magnetron sputtering with the following parameters: base vacuum 8 × 10⁻⁶ Pa. -4 The conditions were: Pa, argon working pressure 0.5 Pa, argon flow rate 40 Sccm, and DC sputtering power 20 W. Due to the large imaginary part of the refractive index of the W layer, resulting in high light loss, the thickness should not exceed 10 nm. In this case, the deposition thickness was approximately 5 nm, and the overall thickness of the four-layer heterostructure was Ag (27 nm) / CoFeB (8 nm) / W (5 nm) / Ag (10 nm). A sandwich structure Ag (27 nm) / CoFeB (8 nm) / Ag (10 nm) was prepared using the same conditions and methods as a control experiment. Finally, the self-made PDMS microfluidic chip was attached to the metal film side of the substrate for the flow of the biomolecule solution to be tested.

[0088] Step 2: Biofunctionalization of the sensor chip

[0089] The biofunctionalization of the four-layer heterostructure sensitive element in this case is the same as that in Example 3, with IgG protein anchored on the chip as a biorecognition unit for detecting IgG antibodies.

[0090] Step 3: Antigen-antibody specific binding detection

[0091] A 5 μg / ml IgG antibody solution was prepared and introduced into two sensor structures at a flow rate of 50 μL / min. The hybridization process of surface-fixed IgG and IgG antibodies was monitored in real time using time-TMOKE mode (incident angle 52.6°, AC magnetic field frequency 50 Hz, magnetic field strength 50 mT). Results showed that the sensor chip with the added strong spin-orbit coupling layer exhibited a more significant signal change amplitude and a larger signal variation when detecting the same antibody concentration. This demonstrates the improvement of the sensing performance of the magnetic plasmonic structure by the strong spin-orbit coupling thin film material. Based on this, magnetic plasmonic sensor chips with Ag / CoFeB / Ta / Ag, Ag / CoFeB / Pt / Ag, Ag / FeNi / W / Ag, Ag / FeNi / Ta / Ag, and Ag / FeNi / Pt / Ag structures were also prepared. The test results showed that compared with the structure without the strong spin-orbit coupling layer, both the sensitivity and detection limit were significantly improved.

Claims

1. An integrated magnetic plasma sensing and detection system, characterized in that, Its core components mainly include a right-angle prism with a magnetic plasma sensing element, an electromagnet, a silicon photodiode detector, a laser, a polarizer, a 5:5 beam splitter, a lens, a lock-in amplifier, a data acquisition card, a signal generator, the system's mechanical rotation, and a data acquisition system computer; the silicon photodiode detector includes photodetector 1 and photodetector 2; the polarizer includes a polarizing prism and an aperture. First, a polarizer is used to modulate the laser output beam into TM polarized light, and then a 5:5 beam splitter is used to split it into two beams. One beam serves as the incident reference light, collected by lens 1 and photodetector 1, and the resulting signal is transmitted to a data acquisition card for intensity normalization to reduce system noise. The other beam from the 5:5 beam splitter is incident on the right-angled surface of the magnetoplasm sensing element prism, and the light refracted from the other right-angled surface of the prism is collected by lens 2 and photodetector 2. The resulting signal is then used as the input signal of a lock-in amplifier, serving as the response signal for refractive index sensing. The mechanical rotation of the system is used to rotate the prism, ensuring that lens 2 and photodetector 2 rotate simultaneously. The rotation of the prism causes light to refract from the other right-angled surface of the plasma sensing element, allowing it to pass through lens 2 and enter photodetector 2 for testing. The modulation of the external magnetic field of the magnetic plasma sensing element is provided by a vertically placed AC electromagnet, with the prism containing the sensing element placed in the magnetic field. The signal generator outputs two AC signals of the same frequency: one signal is used by the electromagnet to control the frequency of the external magnetic field of the sensing element, while the other signal serves as the reference input for the lock-in amplifier. The reflected light from the sensing element is collected by lens 2 and detector 2 and input into the lock-in amplifier for mixing and filtering. The processed electrical signal is the TMOKE sensing response signal, which is finally transmitted to the computer by the data acquisition card.

2. The integration of a magnetic plasma sensing and detection system according to claim 1, characterized in that, The system's mechanical rotation includes two turntables: coaxial turntable 1 and turntable 2. Turntable 1 holds the prism, while turntable 2 holds the detector 2 and lens 2 via an outwardly extending bracket. Turntables 1 and 2 are driven by stepper motors to ensure that the rotation of turntable 2, as the prism rotates with turntable 1, allows lens 2 and photodetector 2 to receive refracted light. Both turntables 1 and 2 are driven by stepper motors. The system is implemented using a θ–2θ coaxial dual rotary stage system driven by stepper motors. According to the law of reflection, the ratio of the change in the incident angle to the change in the reflection angle is always 1:

2. Therefore, the ratio of the angular velocity of turntable 1 (where the prism is located) to the angular velocity of turntable 2 (where the detector is located) should also be 1:2, thus enabling the measurement of light intensity across the entire angular range. θ–2θ represents the ratio of the angular velocity of turntable 1 (where the prism is located) to the angular velocity of turntable 2 (where the detector is located) and should also be 1:

2.

3. The integration of a magnetic plasma sensing and detection system according to claim 1, characterized in that, The magnetic plasma sensing element is fabricated on the inclined surface of a right-angled triangular prism using physical vapor deposition. A multi-channel PDMS (polydimethylsiloxane) microfluidic chip is integrated on the surface of the magnetic plasma sensing element. The multi-channel PDMS (polydimethylsiloxane) microfluidic chip is used to introduce the liquid to be measured onto the surface of the sensor sensing element, enabling the sensor to perform liquid phase sensing and detection. A stepper motor is used to control the rotation of the turntable, thereby driving the right-angled prism to rotate. The side of the right-angle prism corresponding to the magnetic plasma sensing element is equipped with a container to hold the liquid for testing.

4. The integration of a magnetic plasma sensing and detection system according to claim 1, characterized in that, Fabrication of magnetic plasma multilayer heterojunction thin films: A magnetic plasma sensing element is prepared on the inclined surface of a right-angle prism by physical vapor deposition. The magnetic plasma sensing element is a multilayer heterojunction thin film, which is mainly composed of noble metal materials, magnetic materials (ferromagnetic materials, garnet ferrite materials) and strong spin-orbit coupling materials to form a multilayer heterojunction thin film. The outer surface of the magnetic plasma sensing element is a noble metal thin film material. The specific types can be divided into the following categories: The first category is a sandwich structure composed of noble metal thin film material / ferromagnetic thin film material / noble metal thin film material. This type of structure uses ferromagnetic thin film material as the interlayer embedded in noble metal material. The second category is a bilayer heterostructure composed of noble metal thin film material / garnet ferrite thin film material. This type of structure uses garnet ferrite thin film as magnetic material and covers the top layer with noble metal film for biofunctionalization. The third category is a trilayer heterostructure composed of noble metal thin film material / rare earth-doped garnet ferrite thin film material / garnet ferrite seed layer film. This type of structure introduces a magnetic layer with stronger magneto-optical effect. Generally, a YIG layer is deposited in advance as a seed layer during preparation. The fourth category is a trilayer heterostructure composed of noble metal thin film material / strong spin-orbit coupling thin film material / ferromagnetic thin film material. This type of structure introduces a strong spin-orbit coupling layer to enhance the magneto-optical effect of the ferromagnetic layer. The suitable thickness range for noble metal layers is 40-60 nm, and their main function is to excite surface plasmon resonance; the thickness of ferromagnetic thin film materials is 5-20 nm; the thickness of garnet ferrite and rare earth-doped garnet ferrite thin films is 50-200 nm; and the thickness of strongly spin-orbit coupled thin film materials is 10-20 nm.

5. Noble metal thin film materials are selected from one or more of Au, Ag, Cu, Al, etc.; ferromagnetic thin film materials are selected from Fe, Co, Ni, Mo, Sm, Gd, Nd, Tb and metal alloys and intermetallic compounds based thereon; garnet ferrite thin film materials include rare earth-doped and / or non-rare earth-doped R3Fe5O 12 The rare earth ion R is one or more of Y, Gd, Dy, Bi, etc.; the strong spin-orbit coupling layer is composed of W, Ta, Pt, Bi, Pb and actinide heavy metals and their alloys.

6. The integration of a magnetic plasma sensing and detection system according to claim 4, characterized in that, To enhance the adhesion between the heterojunction film and the substrate, a Cr or Ti film is first sputtered onto the substrate. For noble metal thin film materials and some ferromagnetic thin film materials, direct magnetron sputtering is sufficient. Garnet ferrite thin films need to be annealed to crystallize and exhibit magnetism.

7. The integration of a magnetic plasma sensing and detection system according to claim 1, characterized in that, Biofunctionalization of sensitive element surfaces mainly includes the following steps: The first step is to form a self-assembled monolayer on the surface of the noble metal using mercaptoalkanoic acids (such as HS-(CH2)). 11 The molar ratio of -COOH to mercapto alcohol (e.g., HS-(CH2)6-OH) is 1:9 to self-assemble a monolayer, thereby achieving precise control of carboxyl group distribution density and steric hindrance. Secondly, the self-assembled monolayer is activated by using an EDC / NHS mixture with a molar ratio of 4:1, enabling it to effectively link biomolecules such as proteins. Finally, the activated self-assembled monolayer surface is coupled with the analyte recognition unit.

8. A method of using a magnetic plasma sensing system according to any one of claims 1-7, comprising the following steps: (1) When characterizing the performance of the magnetic plasma sensing system, the sensing element does not need to be biofunctionalized. First, a characterization solution with a certain refractive index is introduced. At this time, the signal generator outputs an AC signal of a certain frequency to control the direction of the electromagnetic field (the reversal of N and the frequency). At the same time, the direction of the electromagnetic field is parallel to the plane of the sensing element, and the magnetic field strength satisfies the saturation magnetization of the magnetic layer. The test adopts the angle-TMOKE mode, that is, the turntable is driven by a stepper motor to change the incident angle of the laser to the sensing element. At each angle of the turntable rotation, the real-time light intensity is recorded by the detector 2 and input to the lock-in amplifier for processing to obtain the TMOKE response. After the light intensity signal obtained by the detector 1 is normalized, the relationship between the angle and the TMOKE spectrum can be obtained, that is, the angle-TMOKE spectrum. The optimal incident angle can be obtained from it. Generally, the angle corresponding to the largest slope in the angle-TMOKE spectrum is selected. (2) Real-time biological detection adopts the fixed incident angle, i.e. optimal incident angle detection mode. At this time, the stepper motor is not used to drive the turntable to change the incident angle, and the biofunctional sensitive element is replaced. During the test, the electromagnet is controlled by the AC signal output by the signal generator to control the magnetic field direction. The magnetic field strength is the same as in step (1). A solution of biomolecules to be detected with a certain concentration is introduced, and then the detector 2 records the light intensity in real time and inputs it to the lock-in amplifier for processing to obtain the change of TMOKE response over time, i.e. finally obtaining the time-TMOKE spectrum.

9. The method according to claim 8, characterized in that... Applications in biomolecular detection.

10. The method according to claim 9, characterized in that, A localized plasmonic biomolecule probe was constructed by coupling noble metal (Au, Ag, Cu, Pt) nanoparticles with the target biomolecules. This probe was used to enhance the sensing response caused by differences in biomolecule concentration through localized electromagnetic field enhancement, thereby reducing the detection limit.

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