CN102349834B - Human body blood sugar concentration noninvasive detection system - Google Patents

Abstract

The invention provides a human body blood sugar concentration noninvasive detection method adopting a sweep frequency optical coherence tomography technology and a system thereof. In the method, infrared optical waves with the sweep frequency light source output wave number of K vertically irradiate the skin through a probe, interference signals of skin tissue back scattering light are measured by an interferometer and a photoelectric detector, in-phase components and quadrature components of OCT (Optical Coherence Tomography) signals are separated from the interference signals in a digital frequency mixing mode, and OCT complex amplitude analytical signals corresponding to the wave number K are reconstructed; when the wave number K of the sweep frequency light source is changed from K0to KN by an equal interval delta K, OCT complex spectrum analytical signals of wave number K space are obtained, IDFT (Inverse Discrete Fourier Transform) is carried out on the OCT complex spectrum analytical signals to obtain OCT signals of coordinate Z space; and the amplitude and the phase information of skin corium layer back scattering light are computed according to the OCT signals of the coordinate Z space, the average refractive index of the corium layer is computed, the blood sugar concentration is computed according to the correlation of the refractive index and the blood sugar concentration, and the noninvasive detection of the human body blood sugar is realized.

Description

Human blood glucose concentration noinvasive detection system

[technical field]

The present invention relates to the technology that the human blood glucose concentration noinvasive detects, be particularly related to method and the system thereof of the blood sugar concentration noinvasive detection of a kind of employing frequency sweep optical coherence tomography (Swept Source Optical Coherence Tomography is called for short SS-OCT).

[background technology]

Along with improving constantly of China's living standards of the people, the sickness rate of diabetes is just in rising trend, the effective ways that do not have at present radical cure, main clinical means is to come blood sugar concentration is closely controlled by frequent detection and injection of insulin to blood glucose concentration value, cooperate corresponding diet and rule of life, control PD.Used detection method generally has wound now, namely detects by carrying out more accurate biochemical enzyme process behind the venous blood sampling, or uses reagent paper to carry out colorimetric measurement after the blood sampling of acupuncture finger tip.This have detection method length detection time, the sampling number of wound many, brought great misery and inconvenience to the patient, also exists and take a blood sample and the cross infection of initiation and the problem of environmental pollution, is not suitable for the continuous monitoring of blood glucose.In the urgent need to developing easy accurately noinvasive blood trouble sugar measuring method.

In recent years, the optical profile type Woundless blood sugar is measured with its unique advantage, the extremely favor of researcher.The optical profile type non-invasive methods mainly comprises polarimetry, light scattering coefficient method, near infrared spectroscopy, pulse optoacoustic method.Chinese patent (application number 200710092791.5) " a kind of orthogonal double polarizing light non-invasive continuous blood sugar measuring apparatus " is with behind the cross-polarization rayed thumb of modulating, measure the light intensity difference of two orthogonal directions of transillumination with analyzer, calculate blood sugar concentration according to this difference; The method adopts transillumination to detect, and polarized light will pass through all too many levels such as finger skin, subcutaneous tissue, blood vessel, skeleton and fingernail, only depends on the light intensity difference of two orthogonal directions measuring transillumination can't eliminate above-mentioned impact, and its measurement accuracy is difficult to guarantee; The method is not only used the polarizer, orthogonal formula electrooptic modulator in addition, also needs Faraday polarization apparatus, thumb thickness measuring unit, focuses on convex lens, lock-in amplifier, square-wave generator and some feedback driver elements, and system is comparatively numerous and jumbled, and the method haves much room for improvement.

The major obstacle of near infrared spectroscopy is owing to contain a large amount of moisture in the biological tissue, water causes the decay of light serious to Optical Absorption, the content of glucose in tissue fluid and blood is low and excursion is little, the absorptance of glucose also is far smaller than the absorptance of water, it is very faint that blood sugar concentration changes the useful signal that causes, the physiological change of tissue, blood flow, the light intensity that measuring condition etc. cause changes much larger than the light intensity that causes because of sugared concentration change and changes, if do not have stable measuring condition or effectively eliminate the method for measuring noise, the information that near infrared spectroscopy accurately extracts blood sugar concentration be difficult to (referring to document: Liu Rong, the people such as Xu Kexin, " subject matter and progress in the detection of optics Woundless blood sugar ", Chinese science, 2007, the 37th volume supplementary issue: the 124-131 page or leaf).

The light scattering coefficient general theory of law is the back-scattering light of human body tissue, calculates the tissue reduced scattering coefficient.Owing to component content in this scattering coefficient satellite changes, therefore can obtain the situation of change of component content in the body by the variation of following the trail of reduced scattering coefficient.Such as Xu Min, Li Shengli is write articles the blood sugar concentration optical scattering measuring system of introduction (referring to document: Xu Min, Li Shengli, " the blood sugar detecting method research of Scattering Measurement ", sensing technology journal, in February, 2006, the 9th volume first phase, the 100-103 page or leaf), LASER Light Source and a photodetector that this system is 805nm by four wavelength are measured scattered light intensity, obtain the scattering coefficient of tissue, thereby calculate blood sugar concentration; Because the method can't be from the impact on measurement result of the strength information separating muscle of scattered light, skeletal tissue, therefore, the measurement effect of blood sugar concentration light scattering coefficient method is still waiting further improvement too.

[summary of the invention]

In order to solve the problems of the prior art, the invention provides a kind of interference that does not relate to all multiple locations such as muscle, blood vessel, skeleton and fingernail, the human blood glucose concentration noinvasive detection technique that measurement links is less.

The invention provides a kind of human blood glucose concentration noinvasive detection method, it may further comprise the steps:

Step 1: frequency-sweeping laser source output infrared waves, its wave number represents with K, by detection probe irradiation human body skin, the back-scattering light that is returned by skin histology is sent to interferometer by described detection probe, and described back-scattering light comprises reflected light and refracted light;

Step 2: this interferometer interferes back-scattering light and reference light, and introduces phase-modulator in reference path, and the adding frequency is ω _CThe cosine carrier driving voltage, make additional variation of phase place of reference light

Adopt photodetector to measure the overall strength I of interference light _T(K), I _T(K) comprise the self correlation intensity, the cross-correlation intensity between each back-scattering light of back-scattering light of self correlation intensity, each layer tissue of skin of reference light, four contents of the cross-correlation intensity of each back-scattering light and reference light, the 4th OCT signal that is called each layer tissue of skin;

Step 3: with A/D converter to above interference light overall strength signal I _T(K), take the sampling period as T _S, carry out the time discretization sampling, obtain its digital signal sequences I ^DIG _TAnd be kept at signal processing unit (K);

Step 4: for from interference light overall strength signal I ^DIG _T(K) isolate the OCT signal of each layer tissue of skin in, above digital signal sequences I ^DIG _T(K) with frequency 2 ω _CDigital local oscillator signal NCO1=cos (2 ω _CNT _S) do digital mixing, behind the FIR wave digital lowpass filter, obtain the in-phase component I of OCT digital signal ^I_DIG _OCT(K), again with signal sequence I ^DIG _T(K) and frequencies omega _CDigital local oscillator signal NCO2=cos (ω _CNT _S) do digital mixing, behind the FIR wave digital lowpass filter, obtain the quadrature component I of OCT digital signal ^Q_DIG _OCT(K), wherein n is positive integer;

Step 5: the in-phase component I that gets the OCT signal ^I_DIG _OCT(K) be the complex amplitude real part, get the quadrature component I of OCT signal ^Q_DIG _OCT(K) be the complex amplitude imaginary part, the OCT complex amplitude analytic signal that reconstruct is corresponding with wave number K, that is:

${I^{\mathrm{DIG}}}_{\mathrm{OCT}}\left(K\right)=\frac{{I^{I\_\mathrm{DIG}}}_{\mathrm{OCT}}\left(K\right)}{J_2\left(M_C\right)}-\frac{{{\mathrm{jI}}^{Q\_\mathrm{DIG}}}_{\mathrm{OCT}}\left(K\right)}{J_1\left(M_C\right)}$

J ₁(M _C), J ₂(M _C) be 1 grade and 2 grades of Bessel functions,

M _CBe the reference light phase-modulation degree of depth,

Wherein $j=\sqrt{-1};$

Step 6: when the wave number K of swept light source output light-wave by K ₀To K _NWhen changing with δ K uniformly-spaced, i.e. K=K ₀+ n δ K,

OCT complex amplitude analytic signal I ^DIG _OCT(K) by I ^DIG _OCT(K ₀) change to I ^DIG _OCT(K _N), obtain one group of OCT complex-specturm analytic signal in the wave number K space, with I ^DIG _OCT(K) write as sequence signal form: I ^DIG _OCT(K)=I ^DIG _OCT(n δ K), wherein N is positive integer;

Step 7: when the wave number K of swept light source output light-wave by K ₀To K _NWhen changing with δ K uniformly-spaced, in the mode of synchronous monitoring, obtain and wave number K=K ₀The light source output power value P (n δ K) that+n δ K is corresponding utilizes P (n δ K) to above OCT complex-specturm analytic signal I ^DIG _OCT(n δ K) does normalized, obtains normalized OCT complex-specturm analytic signal I in the wave number K space ^Norm _OCT(n δ K):

${I^{\mathrm{Norm}}}_{\mathrm{OCT}}\left(\mathrm{n\δK}\right)=\frac{{I^{\mathrm{DIG}}}_{\mathrm{OCT}}\left(\mathrm{n\δK}\right)}{P\left(\mathrm{n\δK}\right)};$

Step 8: to above normalized OCT complex-specturm analytic signal I ^Norm _OCT(n δ K) makes discrete inverse Fourier transform IDFT, obtains the OCT signal in coordinate Z space, and Z is along the light wave incident direction, and write as sequence signal form: U (Z)=U (n δ Z), Z=Z ₀+ n δ Z,

And U (n δ Z) comprises real part U _r(n δ Z) and imaginary part U _i(n δ Z):

$U\left(\mathrm{n\δZ}\right)=\frac{1}{N+1}\underset{i=0}{\overset{N}{\mathrm{\Σ}}}{I^{\mathrm{Nom}}}_{\mathrm{OCT}}\left(\mathrm{n\δK}\right)\exp \left(\frac{2\mathrm{\π}*i*n}{N+1}\right)$

$=U_r\left(\mathrm{n\δZ}\right)+{\mathrm{jU}}_i\left(\mathrm{n\δZ}\right)$

$U_r\left(\mathrm{n\δZ}\right)=\frac{1}{N+1}\underset{i=0}{\overset{N}{\mathrm{\Σ}}}{I^{\mathrm{Nom}}}_{\mathrm{OCT}}\left(\mathrm{n\δK}\right)\cos \left(\frac{2\mathrm{\π}*i*n}{N+1}\right)$

$U_i\left(\mathrm{n\δZ}\right)=\frac{1}{N+1}\underset{i=0}{\overset{N}{\mathrm{\Σ}}}{I^{\mathrm{Nom}}}_{\mathrm{OCT}}\left(\mathrm{n\δK}\right)\sin \left(\frac{2\mathrm{\π}*i*n}{N+1}\right)$

Z wherein ₀Be the most surperficial space coordinates of skin, Z _NFor light wave arrives at the innermost space coordinates of skin;

Step 9: with the OCT signal U in the coordinate Z space (n δ Z), write as exponential form by following formula:

$\left|U\left(\mathrm{n\δZ}\right)\right|=\sqrt{{U_r}^2\left(\mathrm{n\δZ}\right)+{U_i}^2\left(\mathrm{n\δZ}\right)}$

Wherein | U (n δ Z) | and

Be respectively mould and the argument of OCT signal U (n δ Z), the two comprises Z=Z ₀Amplitude and the phase information of the skin histology back-scattering light that+n δ Z is corresponding;

Step 10: by to mould

Do the difference coefficient of coordinate Z and calculate, find the position of back-scattering light amplitude hit point, thereby determine the space coordinates scope of dermal layer of the skin, namely determine the upper interface of skin corium, the space coordinates at lower interface;

Step 11: be respectively Z=Z if determine the upper interface of skin corium, the space coordinates at lower interface ₀+ l δ Z, Z=Z ₀+ m δ Z, N 〉=m 〉=l, m, l are positive integer, and then the phase contrast of the upper and lower interface of skin corium back-scattering light is calculated as follows:

Step 12: by the phase contrast of the upper and lower interface of skin corium back-scattering light, calculate the mean refractive index of dermal layer of the skin

Wherein

Be the average wave number of swept light source,

Step 13: last mean refractive index according to dermal layer of the skin With the relevance formula of blood sugar concentration Cg, calculate human blood glucose concentration C _g:

$\stackrel{\‾}{n}=1.348+1.17*{10}^{-6}*C_g.$

The infrared waves of the method take swept light source output wave number as K by probe vertical irradiation skin, measured the interference signal of skin histology back-scattering light with interferometer and photodetector, with A/D converter to above interference signal, take the sampling period as T _S, carry out the time discretization sampling, obtain its digital signal sequences, be kept at signal processing unit; From interfere digital signal sequences, separate in-phase component and the quadrature component of OCT signal in the digital mixing mode, and reconstruct the OCT complex amplitude analytic signal corresponding with wave number K; When the wave number K of swept light source by K ₀To K _N, when changing with δ K uniformly-spaced, obtain the OCT complex-specturm analytic signal in wave number K space, this OCT complex-specturm analytic signal is discrete inverse Fourier transform IDFT, obtain the OCT signal of coordinate Z space (Z is along the light wave incident direction); According to amplitude and the phase information of the OCT calculated signals dermal layer of the skin back-scattering light in coordinate Z space, calculate the mean refractive index of skin corium, according to the dependency of refractive index and blood sugar concentration, calculate blood sugar concentration, realize the noinvasive detection of blood sugar for human body.

The OCT technology has time domain OCT (Time Domain OCT) and frequency domain OCT (Spectral DomainOCT) two large classes.Time domain OCT is in order to obtain the axial depth information of sample, reference arm must carry out axial depth scanning, and it is that data acquisition is carried out in pointwise in the scope of coherence length, has so just greatly limited its scanning speed, therefore, time domain OCT system scanning means complex structure; Image taking speed is not high; Common frequency domain OCT system mainly is comprised of weak relevant wideband light source, interferometer and spectrogrph etc.The axial depth scanning information (tomographic map) of tested sample obtains as inverse Fourier transform by the interference spectrum to spectrogrph output, and the scanning information of different depth can simultaneously or walk abreast and obtain.Relative time domain OCT, common frequency domain OCT need not the axial scan of reference arm, has got rid of the device of axial scan, has improved image taking speed and detectivity; Yet, because the spectrogrph of frequency domain OCT system generally adopts optical dispersion element (prism or scattered grating) and surface imaging ccd detector to make up, utilize dispersion element with the different wave length ripple in the locus separately, and in the ccd detector imaging.Therefore, common frequency domain OCT systems technology is complicated, expensive.Can't satisfy the cost requirement of human blood glucose concentration noinvasive detection field.

SS-OCT belongs to a kind of technology distortion of frequency domain OCT (Spectral Domain OCT), the SS-OCT system mainly is comprised of frequency-sweeping laser source, interferometer and photodetector, when the wave number K of swept light source by when uniformly-spaced changing, measure by interferometer, photodetector and A/D converter, obtain in-phase component and quadrature component with the corresponding OCT signal of wave number K, and be built into the OCT spectrum signal in wave number K space, this OCT spectrum signal is discrete inverse Fourier transform IDFT, obtains the axial depth scanning information of tested sample.SS-OCT had both got rid of the reference arm axial scan device of time domain OCT; The spectrogrph parts of having got rid of again frequency domain OCT system, light requirement is not learned dispersion element and surface imaging ccd detector; Photodetector that SS-OCT adopts can be normal optical fulgurite (point probe), and its cost is far below the CCD of surface imaging, and the technical complexity of its interface circuit is more far below the CCD device.Therefore, at human blood glucose concentration noinvasive detection field, SS-OCT has time domain OCT and the incomparable technical advantage of common frequency domain OCT.

Adopt the back-scattering light of SS-OCT commercial measurement skin histology, polarimetry, light scattering coefficient method, near infrared spectroscopy do not relate to the interference of all multiple locations such as muscle, blood vessel, skeleton and fingernail relatively, and measurement links is less; In-phase component and the quadrature component of from interference signal, separating the OCT signal in the digital mixing mode, OCT complex-specturm analytic signal in the reconstruct wave number K space, OCT complex-specturm analytic signal is discrete inverse Fourier transform IDFT, obtain amplitude and the phase information of dermal layer of the skin back-scattering light, fully utilized amplitude and the phase information of back-scattering light.So the method that blood sugar concentration noinvasive provided by the invention detects has stronger technical advantage.

As a further improvement on the present invention, shining human body skin by detection probe in the step 1 is vertical irradiation.

The advantage of vertical irradiation is: (1) incident optical power loss is little, guarantees that most of light enters human body skin; (2) return path of back-scattering light approaches and vertically returns, and has greatly reduced light and has reflected or the diffuse-reflectance component plunderring of skin surface, and scattering back light and detection probe coupling efficiency are high.

Human blood glucose concentration noinvasive detection system, it comprises frequency-sweeping laser source, the light source synchronous tracker, interferometer, detection probe, the first photodetector, A/D converter, signal processing unit, control unit, described frequency-sweeping laser source is provided with optical fiber interface A and electric interfaces B, described light source synchronous tracker comprises a 1x2 fiber coupler and the second photodetector, the A of fiber coupler, B, the C arm successively with frequency-sweeping laser source, the second photodetector, interferometer links to each other, described interferometer has three optical fiber interface A, B, C and electric interfaces D, the optical fiber interface A of interferometer, B, C successively with the light source synchronous tracker, the first photodetector, detection probe links to each other, the electric interfaces D of interferometer links to each other with control unit, the first photodetector, the second photodetector all links to each other with A/D converter, described control unit also with frequency-sweeping laser source, A/D converter, signal processing unit links to each other.

The infrared waves of frequency-sweeping laser source output, its wave number represents with K, at first enter the light source synchronous tracker, be divided into two by wherein a 1x2 fiber coupler, one road light send the second photodetector, this second photodetector is with A/D converter, control unit etc., mode with synchronous monitoring, obtain the luminous power output valve corresponding with wave number K of light source, the optical fiber interface A of another Lu Guangjing interferometer enters interferometer, optical fiber interface C by interferometer is sent to detection probe, is used for shining human body skin; The back-scattering light that is returned by skin histology is detected again the probe collection and is recycled to interferometer, under the effect of interferometer, forms interference light, and the overall strength signal of its interference light is received by the first photodetector by the optical fiber interface B of interferometer; The overall strength signal of the interference light that the first photodetector is exported is by A/D converter, take the sampling period as T _SCarry out the time discretization sampling, obtain its digital signal sequences, signal processing unit carries out a series of calculation process to this digital signal sequences, comprise the space coordinates location of strength signal normalized, digital local oscillator generation, digital mixing, FIR digital low-pass filtering, analytic signal reconstruct, FFT, skin histology, the refractive index inverting scheduling algorithm module of skin corium, realize that the noinvasive of blood sugar for human body detects.Therefore, the hardware module of native system is few, and the analog circuit that particularly is subject to external interference is few, and system is convenient, flexible, is easy to realize, reliability is high.

As a further improvement on the present invention, described interferometer comprises optical fiber circulator, the 2nd 1x2 fiber coupler, fibre optic phase modulator, optical fiber type faraday reflecting mirror, three optical fiber interfaces 1 of optical fiber circulator, 2,3 successively with the light source synchronous tracker, the A arm of the 2nd 1x2 fiber coupler, the first photodetector links to each other, the B arm of the first fiber coupler links to each other with detection probe, the C arm of the 2nd 1x2 fiber coupler links to each other with the optical fiber interface 1 of phase-modulator, the optical fiber interface 2 of this phase-modulator links to each other with optical fiber type faraday reflecting mirror, and the electric interfaces 3 of this phase-modulator links to each other with control unit.

Infrared light from the light source synchronous tracker, optical fiber interface 1 through optical fiber circulator enters optical fiber circulator first, because the current characteristics of optical fiber circulator light annular, this infrared light can only be from the optical fiber interface 2 of optical fiber circulator out, and enter the 2nd 1x2 fiber coupler by the A arm of the 2nd 1x2 fiber coupler, be divided into two by the 2nd 1x2 fiber coupler, one road light is delivered to detection probe through the B of the first fiber coupler arm, be used for shining human body skin, the back-scattering light of skin histology again detected probe is collected, and be back to the B arm of the first fiber coupler as flashlight; Another light path arrives at optical fiber type faraday reflecting mirror behind the C of the first fiber coupler arm and phase-modulator, the reflected light of faraday's reflecting mirror is recycled to the C arm of the first fiber coupler as reference light.Above flashlight and reference light converge at the A arm of the 2nd 1x2 fiber coupler, because the formed interference light of interference effect enters optical fiber circulator by the optical fiber interface 2 of optical fiber circulator again, because the current characteristics of optical fiber circulator light annular, this interference light can only be from the optical fiber interface 3 of optical fiber circulator out, and received by the first photodetector.Like this, guaranteed that interference light can not return light source synchronous tracer and frequency-sweeping laser source, thereby avoided causing the optical power monitoring error of light source synchronous tracker and the performance inconsistency of LASER Light Source because of the impact of interference light.Therefore, native system adopts the interferometer of optical fiber circulator design that good stability is arranged.

As a further improvement on the present invention, described detection probe comprises optical fiber GRIN Lens, convex lens 1, convex lens 2 and convex lens 3, four be centered close on the same optical axis, wherein GRIN Lens links to each other with interferometer by optical fiber, the light output end of GRIN Lens is positioned at the front focus place of convex lens 1, and the back focus of convex lens 2 overlaps with the front focus of convex lens 3.

As a further improvement on the present invention, the length L of described GRIN Lens is between (12) P and (3/4) P, and P is intercept.

As a further improvement on the present invention, the focal length of convex lens 1 and convex lens 2 are all f1, and protruding 3 focal length is f2, and f2 is greater than f1.

Described GRIN Lens links to each other with interferometer by optical fiber, and the length L of GRIN Lens is between (12) P and (3/4) P.So in the process of infrared waves by probe irradiation skin, after GRIN Lens, the bright dipping of its end face becomes with optical axis less than 45 ° of angles from the light of interferometer, its whole light all are radiated on the convex lens 1; Because the end face of GRIN Lens is positioned at the front focus place of convex lens 1, behind the light planoconvex lens 1, form the directional light that is parallel to optical axis that expands for the first time; This directional light planoconvex lens 2 converges at again the back focus place of convex lens 2; Because the back focus of convex lens 2 overlaps with the front focus of convex lens 3, behind the light planoconvex lens 3, form at last the directional light that expands for the second time.

Therefore, in the process of infrared waves by probe irradiation skin, described detection probe becomes directional light with the bright dipping of interferometer, and has the effect that the light secondary expands, and makes infrared waves vertical irradiation skin become possibility.

When infrared light wave vertical irradiation skin, the back-scattering light of skin histology can be considered as vertically returning approx, specifically return process: the light that returns at first enters convex lens 3, its most of light converges at the front focus place of convex lens 3, because the back focus of convex lens 2 overlaps with the front focus of convex lens 3, becomes the directional light that is parallel to optical axis behind the light planoconvex lens 2; The directional light planoconvex lens 1 that this returns converges at again the front focus place of convex lens 1; Again because the end face of GRIN Lens is positioned at the front focus place of convex lens 1, this converges the end face that light positive arrives at well GRIN Lens, under the effect of GRIN Lens, the light that returns is as flashlight, B arm by optical fiber, the first fiber coupler enters interferometer, interfere with reference light, form interference light.

Therefore, this detection probe can be utilized the back-scattering light of skin histology fully in the process of collecting flashlight, makes the maximization of its light that returns and the coupling efficiency between the fibre optic interferometer.

As a further improvement on the present invention, described signal processing unit mainly is comprised of microprocessor or DSP or FPGA and internal algorithm software.

As a further improvement on the present invention, this algorithm software comprises the space coordinates location of interference signal sampling, strength signal normalized, digital local oscillator NCO generation, digital mixing, FIR digital low-pass filtering, analytic signal reconstruct, FFT, skin histology, the modules such as refractive index inversion algorithm of skin corium.

Software radio is a kind of brand-new communication design concept of rising gradually later in 90 years 20th century, its core concept is: load different communication software at an opening, modular general hardware platform, and the as far as possible close radio-frequency antenna of digitized processing (A/D and D/A conversion), realize communication function by software as much as possible.The software implementation of function can reduce the hardware circuit of very flexible, especially reduces the simulation link, makes communication system possess opening, comprehensive programmability, is convenient to system upgrade, Function Extension.

The present invention has used the design concept of " software radio " just, the design philosophy of " software radio " is introduced the research of " human blood glucose concentration noinvasive detection method and system ", the hardware platform of the described signal processing unit of system is mainly to be realized that by microprocessor or DSP or FPGA the internal algorithm software of described signal processing unit has the interference signal sampling, the strength signal normalized, digital local oscillator produces, digital mixing, the FIR digital low-pass filtering, analytic signal reconstruct, FFT, the space coordinates location of skin histology, the modules such as the refractive index inversion algorithm of skin corium form.Each functions of modules is as follows: interference signal sampling: the control A/D converter, take the sampling period as T _S, to the interference light intensity signal of the first photodetector output, carry out the time discretization sampling, obtain its digital signal sequences.Because the present invention introduces phase-modulator in the interferometer reference path, the adding frequency is ω _CThe cosine carrier driving voltage, make additional variation of phase place of reference light, thereby, comprise zero-frequency, carrier wave fundamental frequency omega in the frequency spectrum of interference light intensity signal _CAnd ω _CHigh order frequency is if the frequency component of interference signal is 10 ω _CFrequency multiplication is according to the Nyquist sampling theory, for undistorted recovery original analog, then sampling rate f _s $(f_S=\frac{1}{T_S}),$ Should satisfy: $f_S\&GreaterEqual;20\frac{{\mathrm{\&omega;}}_{\mathrm{<figure-callout\; id="2"\; label="C"\; filenames="121024144228.png,121024144243.png"\; state="\{\{state\}\}">C</figure-callout>}}}{2\mathrm{\&pi;}}.$

The strength signal normalized: because when the wave number K of swept light source output light-wave when uniformly-spaced changing, the light source output power value may change thereupon; Adopt the mode of synchronous monitoring, obtain the light source output power value corresponding with wave number K, utilize the synchronous monitoring value, interference signal is done normalized, the output that can eliminate swept light source fluctuates with the variation of K, and then the detection accuracy of Effective Raise system.

Digital local oscillator produces: produce near desirable cosine wave with digital method, comprise NCO1=cos (2 ω _CNT _S) and NCO2=cos (ω _CNT _S); NCO1, NCO2 obtain with the ROM look-up table, namely the phase value according to cosine wave calculates cosine value, and press this phase value as address variable (pointer variable), store the cosine value corresponding with this phase value in the ROM device, call when doing digital mixing for signal processing unit (microprocessor or DSP or FPGA).The process that digital local oscillator produces comprises: phase-accumulated, phase place addition, cosine are tabled look-up etc.There are the problems such as frequency shakiness, phase noise, temperature drift in simulation local oscillation signal circuit.Relative simulation local oscillation signal circuit, the performance of digital local oscillator signal will stablize, more reliable, convenient, flexible during use, cost is cheaper.Digital mixing: the purpose of digital mixing: in-phase component and the quadrature component of from interference light overall strength signal, isolating the OCT signal of each layer tissue of skin.Its process is its digital signal sequences and digital local oscillator signal NCO1=cos (2 ω _CNT _S) multiply each other, behind the FIR wave digital lowpass filter, obtain the in-phase component of OCT digital signal, again with signal sequence and digital local oscillator signal NCO2=cos (ω _CNT _S) multiply each other, behind the FIR wave digital lowpass filter, obtain the quadrature component of OCT digital signal.Therefore, there is not the nonlinear problem of Analogue mixer in digital mixing.

The FIR digital low-pass filtering: the present invention repeatedly uses the FIR wave digital lowpass filter in the digital mixing process.The FIR wave digital lowpass filter only has the forward direction branch road, does not have feedback circuit, has the linear phase shift characteristic.

Analytic signal reconstruct: after the in-phase component that obtains the OCT signal and quadrature component, the in-phase component of getting the OCT signal is the complex amplitude real part, the quadrature component of getting the OCT signal is the complex amplitude imaginary part, the OCT complex amplitude analytic signal that reconstruct is corresponding with wave number K, i.e. OCT spectrum signal.

The purpose of FFT:FFT algorithm is: OCT complex-specturm analytic signal is made discrete inverse Fourier transform IDFT, obtain the OCT signal in coordinate Z space, Z is along the light wave incident direction.And the OCT signal in coordinate Z space is complex signal, and the mould of this complex signal and argument have characterized respectively amplitude and the phase information of skin histology back-scattering light.

The space coordinates of skin histology location: calculate by the difference coefficient of the mould of the OCT signal in coordinate Z space being coordinate Z, can find the position of back-scattering light amplitude hit point, thereby determine the space coordinates scope of dermal layer of the skin, also namely determine the upper interface of skin corium, the space coordinates at lower interface;

The refractive index inversion algorithm of skin corium: this process mainly comprises: (1) calculates the phase contrast of skin corium upper and lower interface back-scattering light according to the upper interface of skin corium, the space coordinates at lower interface; (2) according to the phase contrast of skin corium upper and lower interface back-scattering light, calculate the mean refractive index n of dermal layer of the skin.

In sum, the present invention has used the design concept of " software radio ", the design philosophy of " software radio " is introduced the research of " human blood glucose concentration noinvasive detection method and system ", A/D converter is placed on after the first photodetector, as early as possible with the interference signal digitized, and replace traditional analogy method with digital method as far as possible, take microprocessor or DSP or FPGA as the hardware platform of signal processing unit, adopt the many algorithms software module to realize the interference signal sampling, the strength signal normalized, digital local oscillator produces, digital mixing, the FIR digital low-pass filtering, analytic signal reconstruct, FFT, the space coordinates location of skin histology, the functions such as the refractive index inverting of skin corium.Significantly reduce the simulation link of very flexible, avoid the analog circuit problem, the problem includes: non-linear, the problems such as frequency is unstable, phase noise and temperature drift make cost performance, the overall performance of system that good improvement be arranged.

[description of drawings]

Fig. 1 is that SS-OCT blood glucose noinvasive detection system of the present invention is always schemed;

Fig. 2 is interferometer structure figure of the present invention;

Fig. 3 is detection probe structure chart of the present invention;

Fig. 4 is digital mixing flow chart of the present invention;

Fig. 5 is complex-specturm analytic signal process chart of the present invention.

[specific embodiment]

The present invention is further described below in conjunction with description of drawings and the specific embodiment.

Ultimate principle of the present invention:

Human body skin mainly is made of epidermis, corium and subcutaneous tissue etc., and the light refractive index of every layer tissue is different, skin corium has flourishing blood microcirculation network, blood sugar for human body is by the blood microcirculation network, form dextrose components at skin corium and pile up, the increase of concentration of glucose can cause the average optical refractive index of skin corium to increase.With light-wave irradiation human body skin surface, when light wave enters skin corium, in the at the interface generation scattering of skin corium and return back-scattering light; And the variation of light refractive index causes the variation of the backscattering light intensity of skin corium on the one hand, causes on the other hand the variation of the backscattering light phase of skin corium.Therefore, amplitude and the phase information of electric field intensity that can be by measuring the back-scattering light that skin corium returns, calculate the phase contrast of the upper and lower interface of skin corium back-scattering light, calculate again the average optical refractive index of skin corium, according to the relevance formula of refractive index and blood sugar concentration, calculate human blood glucose concentration at last.

Technological core of the present invention:

Frequency-sweeping laser source output wave number is the infrared waves of K, by detection probe irradiation human body skin, if skin is considered as the lamellar tissue of N shell structure, because the light refractive index of every layer tissue is different, when light wave enters skin inside, in the at the interface generation scattering of every adjacent layer and return back-scattering light.The electric field that is located at the backscattering light wave at i interface is

The electric field of the total scattering light wave of N shell tissue is

If the electric field intensity of reference light wave is

When the total electric field vector of total scattering light wave by the interference light wave of interferometer and reference light wave formation

$\stackrel{\&RightArrow;}{E_T}=\stackrel{\&RightArrow;}{E_R}+\stackrel{\&RightArrow;}{E_S}=\stackrel{\&RightArrow;}{E_R}+\underset{i=0}{\overset{i=N}{\mathrm{\&Sigma;}}}\stackrel{\&RightArrow;}{E_i}---\left(1\right).$

Adopt photodetector can measure the overall strength I of interference light _T(K), because photodetector is square law device, have $I_T\left(K\right)\&Proportional;\stackrel{\&RightArrow;}{E_T}*\stackrel{\&RightArrow;}{E_T^{*}},$

For

The complex conjugate vector.

$\stackrel{\&RightArrow;}{E_T}*\stackrel{\&OverBar;}{E_T^{*}}={{\mathrm{<figure-callout\; id="2"\; label="E\; R"\; filenames="121024144228.png,121024144243.png"\; state="\{\{state\}\}">E</figure-callout>}}_R}^2+\underset{i=0}{\overset{\mathrm{<figure-callout\; id="2"\; label="N\; E\; i"\; filenames="121024144228.png,121024144243.png"\; state="\{\{state\}\}">N</figure-callout>}}{\mathrm{\&Sigma;}}}{E_i}^{}{{}_{}}^2+2\underset{i\&NotEqual;j}{\overset{N}{\mathrm{\&Sigma;}}}{E}_i{E}_j\cos {\mathrm{\&Phi;}}_{\mathrm{ij}}+2\underset{i=0}{\overset{N}{\mathrm{\&Sigma;}}}{E}_R{E}_i\cos {\mathrm{\&Phi;}}_{\mathrm{Ri}}---\left(2\right)$

E wherein _RElectric field intensity for reference light wave

Mould, E _i, E _jRespectively the mould of the scattered light electric field intensity of skin i layer, j bed boundary, Φ _IjBe the phase contrast between skin i layer scattered light and the j layer scattered light, Φ _RiBe the phase contrast between skin i layer scattered light and the reference light.First reference light in the following formula is self correlation intensity, and second is respectively the self correlation intensity of the back-scattering light of each layer tissue of skin; Cross-correlation intensity between the 3rd all back-scattering light; The 4th is the cross-correlation intensity of the back-scattering light of reference light and each layer tissue of skin, and this is exactly the OCT signal summation of skin histology.

Core content of the present invention provides complete technical scheme, isolate the OCT signal of each layer tissue of skin the overall strength signal of the interference light of measuring from photodetector, obtain back-scattering light amplitude and the phase information of each layer tissue of skin, comprise the upper interface of determining skin corium, the space coordinates at lower interface, obtain the upper interface from skin corium, lower back-scattering light amplitude and phase information at the interface, by the phase contrast of the two, calculate the light refractive index of skin corium, dependency according to this refractive index and blood sugar concentration, calculate human blood glucose concentration, thereby realize that noinvasive detects.

The technical solution adopted for the present invention to solve the technical problems is: a kind of human blood glucose concentration noinvasive detection method that adopts the frequency sweep optical coherence tomography, and it may further comprise the steps:

Step 1: the wave number of frequency-sweeping laser source output is the infrared waves of K, and by detection probe vertical irradiation human body skin, the back-scattering light by skin histology returns is sent to interferometer by this detection probe again;

Step 2: this interferometer interferes back-scattering light and reference light, and introduces phase-modulator in reference path, and the adding frequency is ω _CThe cosine carrier driving voltage, make additional variation of phase place of reference light

The interference light overall strength I that measures of photodetector then _T(K) be:

$I_T\left(K\right)\&Proportional;{{\mathrm{<figure-callout\; id="2"\; label="E\; R"\; filenames="121024144228.png,121024144243.png"\; state="\{\{state\}\}">E</figure-callout>}}_R}^2+\underset{i=0}{\overset{\mathrm{<figure-callout\; id="2"\; label="N\; E\; i"\; filenames="121024144228.png,121024144243.png"\; state="\{\{state\}\}">N</figure-callout>}}{\mathrm{\&Sigma;}}}{E_i}^{}{{}_{}}^2+2\underset{i\&NotEqual;j}{\overset{N}{\mathrm{\&Sigma;}}}{E}_i{E}_j\cos {\mathrm{\&Phi;}}_{\mathrm{ij}}\left(K\right)$

$+2\underset{i=0}{\overset{N}{\mathrm{\&Sigma;}}}{E}_R{E}_i\cos [M_C\cos \left({\mathrm{\&omega;}}_{C}t\right)+{\mathrm{\&Phi;}}_{\mathrm{Ri}}\left(K\right)]---\left(3\right)$

By (3) formula as can be known, I _T(K) comprise the self correlation intensity, the cross-correlation intensity between each back-scattering light of back-scattering light of self correlation intensity, each layer tissue of skin of reference light, four contents of the cross-correlation intensity of each back-scattering light and reference light, the 4th OCT signal that is called each layer tissue of skin.

(3) are launched with the Bessel functional form:

$I_T\left(K\right)\&Proportional;{{\mathrm{<figure-callout\; id="2"\; label="E\; R"\; filenames="121024144228.png,121024144243.png"\; state="\{\{state\}\}">E</figure-callout>}}_R}^2+\underset{i=0}{\overset{\mathrm{<figure-callout\; id="2"\; label="N\; E\; i"\; filenames="121024144228.png,121024144243.png"\; state="\{\{state\}\}">N</figure-callout>}}{\mathrm{\&Sigma;}}}{E_i}^{}{{}_{}}^2+2\underset{i\&NotEqual;j}{\overset{N}{\mathrm{\&Sigma;}}}{E}_i{E}_j\cos {\mathrm{\&Phi;}}_{\mathrm{ij}}\left(K\right)$

$+2\underset{i=0}{\overset{N}{\mathrm{\&Sigma;}}}{E}_R{E}_i\left\{\begin{array}{c}{[J}_0\left(M_C\right)+2\underset{l=1}{\overset{\&infin;}{\mathrm{\&Sigma;}}}{(-1)}^l{J}_{2l}\left(M_C\right)\cos 2l{\mathrm{\&omega;}}_{C}t]{\cos \mathrm{\&Phi;}}_{\mathrm{Ri}}\left(K\right)\\ -2\left[\underset{l=0}{\overset{\&infin;}{\mathrm{\&Sigma;}}}{(-1)}^l{J}_{2l+1}\left(M_C\right)\cos (2l+1){\mathrm{\&omega;}}_{C}t\right]\sin {\mathrm{\&Phi;}}_{\mathrm{Ri}}\left(K\right)\end{array}\right\}---\left(4\right)$

Find out from (4) formula, comprise zero-frequency, carrier wave fundamental frequency omega in the frequency spectrum of only the 4th (the OCT signal of each layer tissue of skin) _CAnd ω _CTherefore high order frequency, can adopt the mixing mode that the 4th and first three items (cross-correlated signal between the autocorrelation signal of relative reference light, the autocorrelation signal of each back-scattering light, each back-scattering light) are separated.

Step 3: with A/D converter to above interference light overall strength signal I _T(K), take the sampling period as T _S, carry out the time discretization sampling, obtain its digital signal sequences I ^DIG _TAnd be kept at signal processing unit (K);

Step 4: for from interference light overall strength signal I ^DIG _T(K) isolate the OCT signal of each layer tissue of skin in, above digital signal sequences I ^DIG _T(K) with frequency 2 ω _CDigital local oscillator signal NCO1=cos (2 ω _CNT _S) do digital mixing, behind the FIR wave digital lowpass filter, obtain the in-phase component I of OCT digital signal ^I_DIG _OCT(K), again with signal sequence I ^DIG _T(K) and frequencies omega _CDigital local oscillator signal NCO2=cos (ω _CNT _S) do digital mixing, behind the FIR wave digital lowpass filter, obtain the quadrature component I of OCT digital signal ^I_DIG _OCT(K), wherein n is positive integer, referring to accompanying drawing 4;

${I^{I\_\mathrm{DIG}}}_{\mathrm{OCT}}\left(K\right)\&Proportional;\underset{i=0}{\overset{N}{\mathrm{\&Sigma;}}}{E}_R{E}_i{J}_2\left(M_C\right)\cos {\mathrm{\&Phi;}}_{\mathrm{Ri}}\left(K\right)$

${I^{Q\_\mathrm{DIG}}}_{\mathrm{OCT}}\left(K\right)\&Proportional;\underset{i=0}{\overset{N}{\mathrm{\&Sigma;}}}{E}_R{E}_i{J}_1\left(M_c\right)\sin {\mathrm{\&Phi;}}_{\mathrm{Ri}}\left(K\right)---\left(5\right)$

J wherein ₁(M _C), J ₂(M _C) 1 grade respectively, 2 grades of Bessel functions;

Step 5: the in-phase component I that gets the OCT signal ^I_DIG _OCT(K) be the complex amplitude real part, get the OCT signal quadrature component I ^Q_DIG _OCT(K) be the complex amplitude imaginary part, the OCT complex amplitude analytic signal that reconstruct is corresponding with wave number K, that is:

${I^{\mathrm{DIG}}}_{\mathrm{OCT}}\left(K\right)=\frac{{I^{I\_\mathrm{DIG}}}_{\mathrm{OCT}}\left(K\right)}{J_2\left(M_C\right)}-\frac{{{\mathrm{jI}}^{Q\_\mathrm{DIG}}}_{\mathrm{OCT}}\left(K\right)}{J_1\left(M_C\right)}---\left(6\right)$

Wherein

Referring to accompanying drawing 5;

Step 6: when the wave number K of swept light source output light-wave by K ₀To K _NWhen changing with δ K uniformly-spaced, namely

OCT complex amplitude analytic signal I ^DIG _OCT(K) by I ^DIG _OCT(K ₀) change to I ^DIG _OCT(K _N), obtain one group of OCT complex-specturm analytic signal in the wave number K space, with I ^DIG _OCT(K) write as the sequence signal form:

I ^DIG _OCT(K)=I ^DIG _OCT(nδK)

Wherein N is positive integer, referring to accompanying drawing 5;

Step 7: when the wave number K of swept light source output light-wave by K ₀To K _NWhen changing with δ K uniformly-spaced, in the mode of synchronous monitoring, obtain and wave number K=K ₀The light source output power value P (n δ K) that+n δ K is corresponding utilizes P (n δ K) to above OCT complex-specturm analytic signal I ^DIG _OCT(n δ K) does normalized, obtains normalized OCT complex-specturm analytic signal I in the wave number K space ^Norm _OCT(n δ K):

${I^{\mathrm{Norm}}}_{\mathrm{OCT}}\left(\mathrm{n\&delta;K}\right)=\frac{{I^{\mathrm{DIG}}}_{\mathrm{OCT}}\left(\mathrm{n\&delta;K}\right)}{P\left(\mathrm{n\&delta;K}\right)}---\left(7\right)$

Referring to accompanying drawing 5;

Step 8: to above normalized OCT complex-specturm analytic signal I ^Norm _OCT(n δ K) makes discrete inverse Fourier transform IDFT, and obtain the OCT signal of coordinate Z space (Z is along the light wave incident direction), and write as sequence signal form: U (Z)=U (n δ Z), $Z=Z_0+\mathrm{n\&delta;Z},\mathrm{\&delta;Z}=\frac{Z_N-Z_0}{N}---\left(8\right)$ And U (n δ Z) comprises real part U _r(n δ Z) and imaginary part U _i(n δ Z):

$U\left(\mathrm{n\&delta;Z}\right)=\frac{1}{N+1}\underset{i=0}{\overset{N}{\mathrm{\&Sigma;}}}{I^{\mathrm{Nom}}}_{\mathrm{OCT}}\left(\mathrm{n\&delta;K}\right)\exp \left(\frac{2\mathrm{\&pi;}*i*n}{N+1}\right)$

$=U_r\left(\mathrm{n\&delta;Z}\right)+{\mathrm{jU}}_i\left(\mathrm{n\&delta;Z}\right)$

$U_r\left(\mathrm{n\&delta;Z}\right)=\frac{1}{N+1}\underset{i=0}{\overset{N}{\mathrm{\&Sigma;}}}{I^{\mathrm{Nom}}}_{\mathrm{OCT}}\left(\mathrm{n\&delta;K}\right)\cos \left(\frac{2\mathrm{\&pi;}*i*n}{N+1}\right)$

$U_i\left(\mathrm{n\&delta;Z}\right)=\frac{1}{N+1}\underset{i=0}{\overset{N}{\mathrm{\&Sigma;}}}{I^{\mathrm{Nom}}}_{\mathrm{OCT}}\left(\mathrm{n\&delta;K}\right)\sin \left(\frac{2\mathrm{\&pi;}*i*n}{N+1}\right)---\left(9\right)$

Z wherein ₀Be the most surperficial space coordinates of skin, Z _NFor light wave arrives at the innermost space coordinates of skin, referring to accompanying drawing 1 and accompanying drawing 5;

Step 9: with the OCT signal U in the coordinate Z space (n δ Z), write as exponential form by following formula:

$\left|U\left(\mathrm{n\&delta;Z}\right)\right|=\sqrt{{U_r}^2\left(\mathrm{n\&delta;Z}\right)+{U_i}^2\left(\mathrm{n\&delta;Z}\right)}---\left(10\right)$

Wherein | U (n δ Z) | and Be respectively mould and the argument of OCT signal U (n δ Z), the two comprises Z=Z ₀Amplitude and the phase information of the skin histology back-scattering light that+n δ Z is corresponding are referring to accompanying drawing 5;

Step 10: by to mould

Do the difference coefficient of coordinate Z and calculate, can find the position of back-scattering light amplitude hit point, thereby determine the space coordinates scope of dermal layer of the skin, also namely determine the upper interface of skin corium, the space coordinates at lower interface;

Step 11: be respectively Z=Z if determine the upper interface of skin corium, the space coordinates at lower interface ₀+ l δ Z, Z=Z ₀+ m δ Z, N 〉=m 〉=l, m, l are positive integer, and then the phase contrast of the upper and lower interface of skin corium back-scattering light is calculated as follows:

Step 12: by the phase contrast of the OCT signal of the skin corium of skin, calculate the mean refractive index of dermal layer of the skin

Wherein

Be the average wave number of swept light source,

Step 13: last mean refractive index according to dermal layer of the skin

With the relevance formula of blood sugar concentration Cg (referring to document: " the blood sugar detecting method research of Scattering Measurement ", sensing technology journal, in February, 2006, the 9th volume first phase, 100-103 page or leaf), calculate human blood glucose concentration C _g:

$\stackrel{\&OverBar;}{n}=1.348+1.17*{10}^{-6}*C_g---\left(12\right);$

The scheme that the present invention solves further technical problem is: it comprises frequency-sweeping laser source, light source synchronous tracker, interferometer, detection probe, the first photodetector, A/D converter, signal processing unit, control unit, described frequency-sweeping laser source is provided with optical fiber interface A and electric interfaces B, described light source synchronous tracker comprises a 1x2 fiber coupler and the second photodetector, and the A of fiber coupler, B, C arm link to each other with frequency-sweeping laser source, the second photodetector, interferometer successively.Referring to accompanying drawing 1.

Described interferometer has three optical fiber interface A, B, C and electric interfaces D, optical fiber interface A, the B of interferometer, C link to each other with light source synchronous tracker, the first photodetector, detection probe successively, the electric interfaces D of interferometer links to each other with control unit, first, second photodetector of photodetector all links to each other with A/D converter, and described control unit also links to each other with frequency-sweeping laser source, A/D converter, signal processing unit.Referring to accompanying drawing 1.

Described interferometer comprises optical fiber circulator, the 2nd 1x2 fiber coupler, fibre optic phase modulator, optical fiber type faraday reflecting mirror, three optical fiber interfaces 1 of optical fiber circulator, 2,3 successively with the light source synchronous tracker, the A arm of the 2nd 1x2 fiber coupler, the first photodetector links to each other, the B arm of the first fiber coupler links to each other with detection probe, the C arm of the 2nd 1x2 fiber coupler links to each other with the optical fiber interface 1 of phase-modulator, the optical fiber interface 2 of this phase-modulator links to each other with optical fiber type faraday reflecting mirror, and the electric interfaces 3 of this phase-modulator links to each other with control unit.Referring to accompanying drawing 2.

Described detection probe links to each other with the signal arm of interferometer, the bright dipping of interferometer is become directional light, and has an effect that light expands, this detection probe is by the optical fiber GRIN Lens, convex lens 1, convex lens 2 and convex lens 3 form, four be centered close on the same optical axis, wherein GRIN Lens links to each other with interferometer by optical fiber, the length L of this GRIN Lens is between (12) P and (3/4) P, P is intercept, the light output end of GRIN Lens is positioned at the front focus place of convex lens 1, the back focus of convex lens 2 overlaps with the front focus of convex lens 3, the light of interferometer is first after GRIN Lens and convex lens 1 form the directional light that expands for the first time, converge to again the front focus of convex lens 3 through convex lens 2, finish for the second time finally by convex lens 3 that directional light expands, the focal length of convex lens 1 and convex lens 2 are all f1, and the focal length of convex lens 3 is f2, and f2 is greater than f1, referring to accompanying drawing 3.

Described signal processing unit mainly is comprised of microprocessor or DSP or FPGA and internal algorithm software, and this algorithm software is comprised of the space coordinates location of interference signal sampling, strength signal normalized, digital local oscillator generation, digital mixing, FIR digital low-pass filtering, analytic signal reconstruct, FFT, skin histology, the refractive index inverting scheduling algorithm module of skin corium.

Above content is the further description of the present invention being done in conjunction with concrete preferred implementation, can not assert that implementation of the present invention is confined to these explanations.For the general technical staff of the technical field of the invention, without departing from the inventive concept of the premise, can also make some simple deduction or replace, all should be considered as belonging to protection scope of the present invention.

Claims (5)

1. human blood glucose concentration noinvasive detection system, it is characterized in that: it comprises frequency-sweeping laser source, the light source synchronous tracker, interferometer, detection probe, the first photodetector, A/D converter, signal processing unit, control unit, described frequency-sweeping laser source is provided with optical fiber interface A and electric interfaces B, described light source synchronous tracker comprises a 1x2 fiber coupler and the second photodetector, the A of the one 1x2 fiber coupler, B, the C arm successively with frequency-sweeping laser source, the second photodetector, interferometer links to each other, described interferometer has three optical fiber interface A, B, C and electric interfaces D, the optical fiber interface A of interferometer, B, C successively with the light source synchronous tracker, the first photodetector, detection probe links to each other, the electric interfaces D of interferometer links to each other with control unit, the first photodetector, the second photodetector all links to each other with A/D converter, described control unit also with frequency-sweeping laser source, A/D converter, signal processing unit links to each other.

2. human blood glucose concentration noinvasive detection system according to claim 1, it is characterized in that: described interferometer comprises optical fiber circulator, the 2nd 1x2 fiber coupler, fibre optic phase modulator, optical fiber type faraday reflecting mirror, three optical fiber interfaces 1 of optical fiber circulator, 2,3 successively with the light source synchronous tracker, the A arm of the 2nd 1x2 fiber coupler, the first photodetector links to each other, the B arm of the first fiber coupler links to each other with detection probe, the C arm of the 2nd 1x2 fiber coupler links to each other with the optical fiber interface 1 of phase-modulator, the optical fiber interface 2 of this phase-modulator links to each other with optical fiber type faraday reflecting mirror, and the electric interfaces 3 of this phase-modulator links to each other with control unit.

3. human blood glucose concentration noinvasive detection system according to claim 1 and 2, it is characterized in that: described detection probe comprises optical fiber GRIN Lens, convex lens 1, convex lens 2 and convex lens 3, four be centered close on the same optical axis, wherein GRIN Lens links to each other with interferometer by optical fiber, the end face of GRIN Lens is positioned at the front focus place of convex lens 1, and the back focus of convex lens 2 overlaps with the front focus of convex lens 3.

4. human blood glucose concentration noinvasive detection system according to claim 3, it is characterized in that: the length L of described GRIN Lens is between (1/2) P and (3/4) P, and P is intercept.

5. human blood glucose concentration noinvasive detection system according to claim 3, it is characterized in that: the focal length of convex lens 1 and convex lens 2 are all f1, and the focal length of convex lens 3 is f2, and f2 is greater than f1.

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