Noninvasive quantitative blood concentration monitoring system and method

Abstract

The invention relates to the technical field of medical diagnosis, and discloses a noninvasive quantitative blood concentration monitoring system. The system comprises an ultrasonic data acquisition platform, an ultrasonic transducer, a light source, a digital delay generator, an energy regulator, an optical fiber and a computer, the ultrasonic data acquisition platform is used for receiving a signal from the ultrasonic transducer, converting the signal into digital data for subsequent analysis and processing, and outputting a trigger signal to the digital delay generator to control the two light sources to perform automatic staggered data acquisition so as to ensure synchronization of laser output and photoacoustic signal acquisition; and the ultrasonic transducer can convert a sound wave signal generated by the photoacoustic effect into an electric signal. According to the novel blood drug monitoring method based on the photoacoustic pumping technology, the blood drug concentration can be accurately obtained without considering the scattering characteristic of tissue, and meanwhile, the novel blood drug monitoring method has the advantages of being high in sensitivity, rapid in response and noninvasive, so that real-time, continuous and accurate blood drug concentration monitoring is achieved.

Description

Technical Field

[0001] The present invention relates to the technical field of medical diagnosis, and particularly to a non-invasive quantitative blood drug concentration monitoring system and method. Background Art

[0002] Precision medicine is to provide treatment accurately at the best time point and with the most appropriate dose according to the individual pharmacokinetic characteristics of patients, so as to maximize the curative effect and minimize the side effects, thereby ensuring the safety and effectiveness of treatment. Blood drug concentration is an important indicator for evaluating drug efficacy and metabolism degree. Blood drug monitoring helps doctors optimize drug dosage, prevent drug poisoning, improve treatment effect, and ensure individualized and precise treatment by measuring the drug concentration in blood. In addition, during the new drug R & D process, blood drug monitoring provides important information on drug metabolism characteristics, pharmacokinetic parameters, etc., supporting the safety and efficacy evaluation of new drugs. Therefore, continuous blood drug monitoring has important clinical significance in drug treatment.

[0003] Conventional drug monitoring methods, such as chromatographic analysis, immunoassay, etc., usually require blood sampling and sample analysis to determine the drug concentration in blood. These methods are time-consuming and laborious, and usually take several hours or days to obtain monitoring results, which may cause doctors to be unable to obtain the drug concentration information of patients in time, thus affecting subsequent diagnosis and treatment decisions. In addition, traditional in vitro monitoring methods usually cannot achieve continuous monitoring, restricting their application in clinics. In contrast, electrochemical sensors for blood drug monitoring can provide continuous real-time data streams, which helps timely intervention in disease treatment. However, the method of implanting the sensor into the subcutaneous tissue of the skin not only may cause pain, but also has certain infection risks and immune rejection reactions, restricting the feasibility of this technology in wide applications. Therefore, a non-invasive quantitative real-time blood drug concentration monitoring method is needed to improve the acceptance of hospitals and patients, thereby greatly popularizing blood drug monitoring.

[0004] Optical molecular detection has the characteristics of non-invasiveness, high sensitivity, high specificity, etc., which well meet the accuracy requirements of circulating blood drug monitoring. However, traditional optical technologies are affected by strong tissue scattering and face challenges in inaccuracy in quantitatively analyzing the molecular concentration in deep tissues. Quantitative photoacoustic tomography is a promising method for detecting the molecular concentration in deep tissues. However, most of these methods rely on various light propagation models to analyze the light intensity distribution. This optical compensation is usually mathematically uncertain, and small input errors easily lead to a significant decrease in accuracy. More importantly, due to the strong absorption of background tissues such as blood, the photoacoustic signals of a few key molecules will be submerged in strong background signals.

[0005] In photoacoustic pumping imaging, the pump light excites the probe molecules to the singlet state, which will enter the triplet state through intersystem crossing, and then generate photoacoustic signals through non-radiative transitions induced by the probe light. Such probe molecules with long-lived triplet states can be regarded as a kind of photoacoustic probe with a repeatedly activatable optical switch. Within its triplet lifetime, its absorption spectrum can be changed by the pump light, while biological background tissues do not have this effect because the excited state lifetime is shorter than the optical pulse. Therefore, by differentiating the photoacoustic signals generated by the probe light with and without the pump light, highly specific molecular detection can be achieved. In short, photoacoustic pumping is expected to become an important tool for drug monitoring in personalized medicine, helping doctors optimize treatment plans according to individual drug responses. Summary of the Invention

[0006] To solve the technical problems that the existing blood drug monitoring methods have certain limitations and cannot meet the requirements of real-time, accurate, and non-invasive, and that it is difficult to compensate for the strong and specific light scattering from the tissues above the target blood vessels in the existing optical and photoacoustic blood drug detection, the present invention provides a non-invasive quantitative blood drug concentration monitoring system and method.

[0007] The present invention is implemented by the following technical solutions: A non-invasive quantitative blood drug concentration monitoring system, comprising:

[0008] An ultrasonic data acquisition platform, an ultrasonic transducer, a light source, a digital delay generator, an energy regulator, an optical fiber, and a computer;

[0009] The ultrasonic data acquisition platform is used to receive the signals from the ultrasonic transducer, convert them into digital data for subsequent analysis and processing, and at the same time output a trigger signal to the digital delay generator to control the two light sources for automatic interleaved data acquisition to ensure the synchronization of laser output and photoacoustic signal acquisition;

[0010] The ultrasonic transducer can convert the acoustic wave signals generated by the photoacoustic effect into electrical signals. When light irradiates the drug molecules, the photoacoustic effect will occur. The drug absorbs light energy and expands, thus generating acoustic waves, and these acoustic waves can be captured and transmitted by the transducer;

[0011] The light source is two pulsed lasers, and the two pulsed lasers respectively generate two light pulses with the same pulse repetition frequency and different central wavelengths, namely the pump pulsed laser and the probe pulsed laser.

[0012] As a further improvement of the above solution, the digital delay generator is controlled by an external trigger from the ultrasonic data acquisition platform to output pulses for each channel. These output pulses are used to control the relative time delay of the two pulsed lasers, that is, the time difference between the pump light and the probe light, to ensure the synchronization of laser output and photoacoustic signal acquisition, specifically as follows:

[0013] (1). Obtain S pummp+probe，Δt When obtaining S pummp+probe，Δt , control the probe pulse to arrive with a very small time delay after the pump pulse reaches the tissue, and the output time of the probe pulse is the time when the ultrasonic data acquisition platform acquires the signal;

[0014] (2). Obtain S pump，Δt When obtaining S pump，Δt , appropriately delay the probe pulse on the original basis;

[0015] (3). Obtain S probe When obtaining S probe , appropriately delay the pump pulse on the original basis.

[0016] As a further improvement of the above solution, the energy regulator is placed at the front end of the outlet of the pump light source to adjust the intensity of the pump light source. The optical fiber is used to couple the output signals of the two pulsed lasers to ensure that the two beams of light irradiate the same position of the tissue.

[0017] As a further improvement of the above solution, the computer is used to receive and process ultrasonic data, perform image reconstruction, analysis, and final blood drug concentration estimation. The computer can control the operation of the system and output the final monitoring result.

[0018] As a further improvement of the above solution, the ultrasonic data acquisition platform starts the digital delay generator through a trigger signal to control the two pulsed lasers to alternately emit pump light and probe light. After the intensity of the pump light is adjusted by the energy regulator, it is coupled with the probe light through the optical fiber and focused on the same position of the target tissue. The drug molecules in the tissue absorb light energy and expand, generating a photoacoustic signal, which propagates in the form of sound waves and is captured by the ultrasonic transducer.

[0019] As a further improvement of the above solution, the ultrasonic transducer converts the received acoustic signal into an electrical signal and transmits it to the ultrasonic data acquisition platform for digital processing. The acquisition platform transmits the data to the computer, and the computer performs image reconstruction and analysis, extracts the pump light signal, probe light signal, and the superimposed signal of the two respectively in combination with the time delay data, and through the analysis and comparison of these signals, realizes the accurate estimation of the blood drug concentration and generates the final monitoring result. During the whole process, the laser output and the photoacoustic signal acquisition are strictly synchronized to ensure the accuracy and timeliness of the data.

[0020] The present invention also provides a non-invasive quantitative blood drug concentration monitoring method, including the following steps:

[0021] (1). By differentiating the photoacoustic signals generated by the probe light with and without the pump light, high-specificity molecular detection is achieved. Therefore, the TTD signal of the probe molecule can be obtained using the following formula:

[0022] S TTD，Δt = Spup+probe，Δt -S pump，Δt -S probe (1)

[0023] Among them, S TTD，Δt is the TTD signal obtained by removing the background through differentiation, S pup+probe，Δt is the photoacoustic signal when the pump and probe lights are simultaneously excited, S pump，Δt is the photoacoustic signal generated by only the excitation of the pump light, S probe is the photoacoustic signal generated by only the excitation of the probe light, and Δt represents the delay time of the probe laser relative to the pump laser;

[0024] (2). Under the excitation of the pump light, the molecule transitions from the ground state (S0) to the first excited singlet state (S1). Due to the spin-orbit coupling effect, the molecule undergoes intersystem crossing and enters the first excited triplet state (T1). Since the transition from the T1 state to the ground state S0 is spin-forbidden, the T1 state has a very long lifetime (τ ∼ 10 -6 -10 s). The molecule in the T1 state can be excited to the T2 state by the probe light, then return to the T1 state through non-radiative relaxation and generate a photoacoustic signal. Therefore, by writing the rate equations of the populations N1, N2, and N3 of the S0, S1, and T1 states during the pump excitation process, the following can be obtained:

[0025]

[0026] Among them, A 21 represents the spontaneous emission coefficient from S1 to S0, B is the light absorption coefficient, I(t) is the laser intensity distribution, A 23 represents the intersystem crossing rate from the S1 state to the T1 state. Since the T1 state has a long lifetime, the spontaneous emission from the T1 state to the ground state S0 is not considered.

[0027] As a further improvement of the above scheme, the following steps are also included:

[0028] (3). Since the lifetime of the S1 state is short, significantly shorter than the typical laser pulse width of about 5 ns, it is assumed that the population of the S1 state does not accumulate and is in a transient equilibrium state, that is, N2 ≈ 0, and dN 2 / dt = 0. Therefore, the following formula can be obtained:

[0029]

[0030] Based on the above, the following can also be obtained:

[0031]

[0032] Among them, N 0 is the initial number of particles in the S0 state, T is the pulse duration, ΦP is the pulse energy of the pump laser. Thus, it can be seen that due to the existence of the metastable state T1, population inversion can be achieved between the T1 state and the S0 state, and the number of particles N in the T1 state 3 increases exponentially with the pump laser energy Φ P and gradually saturates. In contrast, due to the short lifetime of the T2 state, the molecule may undergo multiple excitation - relaxation cycles, so the probe laser will not cause a significant change to N in the T1 state. 3 Therefore, the amplitude of the TTD signal is proportional to N 3 and the energy Φ of the probe laser T and can be written as:

[0033]

[0034] where C 0 is the concentration of the target molecule, Φ P0 is the value when the pump pulse energy reaches 63% saturation, A TTD，Δs is the TTD amplitude at full saturation, μ a is the absorption coefficient, η th is the thermal conversion efficiency, Γ is the Grueneisen parameter, Δt is the delay time of the pump laser relative to the probe laser, τ is the lifetime of the triplet state, and Z us represents the acoustic influence factor, including the acoustic scattering of the tissue, the size of blood vessels, and the detection angle of inclination;

[0035] (4). Assuming that the photoacoustic signal mainly comes from hemoglobin, then it will have a linear relationship with the energy of the probe laser. Therefore, formula (8) can be rewritten as:

[0036]

[0037] where A BV is the photoacoustic amplitude of the probe laser in the target blood vessel, K is a constant, which can be calibrated by samples with known C 0 and hemoglobin concentration. Therefore, the energy of the pump laser can be adjusted, and the maximum ratio R max can be obtained by fitting formula (9). Then, the concentration of the target molecule can be determined by C 0 = R max / K.

[0038] As a further improvement of the above - mentioned scheme, the triplet state is usually formed when a drug molecule is excited by a pump laser and transitions to the singlet excited state, and then enters the triplet state through a non - radiative transition process. Therefore, the absorption spectrum of the drug molecule is selectively changed in response to the pump laser irradiation, allowing the probe laser to detect the drug molecule without background interference through signal differentiation.

[0039] As a further improvement of the above solution, the TTD signal of the triplet drug molecule increases exponentially with the intensity of the pump laser and linearly with the intensity of the probe laser. The intensity of the pump laser determines the proportion of drug molecules excited to the triplet state, and the drug concentration is quantitatively proportionalized using the vascular detection photoacoustic signal as a reference. Therefore, this method can overcome the challenge of inaccurate optical molecular measurements caused by strong and variable light scattering in tissues, thereby achieving non-invasive quantitative monitoring of the concentration of specific target molecules.

[0040] Compared with the prior art, the beneficial effects of the present invention are as follows:

[0041] 1. The new method for blood drug monitoring based on photoacoustic pumping technology of the present invention can accurately obtain the blood drug concentration without considering the scattering characteristics of tissues, and at the same time has the advantages of high sensitivity, fast response, and non-invasiveness, thereby achieving real-time, continuous, and accurate blood drug concentration monitoring. BRIEF DESCRIPTION OF THE DRAWINGS

[0042] Figure 1 Schematic diagram of the blood drug monitoring principle based on photoacoustic pumping detection of the present invention;

[0043] Figure 2 Schematic diagram of the blood drug monitoring system based on photoacoustic pumping of the present invention;

[0044] Figure 3 Schematic diagram of the automatic interleaved data acquisition of the present invention;

[0045] Figure 4 In (a), it is a schematic diagram showing that as the intensity of the pump laser increases, the intensities of the pump image and the TTD image increase, while the intensity of the probe image remains unchanged. (b) is a schematic diagram showing the change of the TTD / probe ratio with the intensity of the pump laser in tubes with different drug molecule concentrations. (c) is a schematic diagram showing the change of the TTD / probe ratio with the drug molecule concentration. (d) is a schematic diagram showing the change of the TTD signal with the intensity of the probe laser;

[0046] Figure 5 In (a), it is a schematic diagram of the ultrasonic image of two closely placed tubes, with the drug molecule concentration of the left tube being 100 μM and the right one being 400 μM. (b) is a schematic diagram of the photoacoustic image of the two tubes. (c) is a schematic diagram showing the change of the TTD / probe ratio with the intensity of the pump laser. (d)-(f) are corresponding schematic diagrams when covering the chicken breast tissue layer. DETAILED DESCRIPTION OF THE INVENTION

[0047] Next, in combination with the drawings and specific embodiments, the present invention will be further described.

[0048] As Figure 2 shown, a non-invasive quantitative blood drug concentration monitoring system includes

[0049] An ultrasonic data acquisition platform, an ultrasonic transducer, a light source, a digital delay generator, an energy regulator, an optical fiber, and a computer;

[0050] The ultrasonic data acquisition platform is used to receive signals from the ultrasonic transducer, convert them into digital data for subsequent analysis and processing, and output a trigger signal to the digital delay generator to control two light sources for automatic interleaved data acquisition, so as to ensure the synchronization of laser output and photoacoustic signal acquisition;

[0051] The ultrasonic transducer can convert the acoustic wave signals generated by the photoacoustic effect into electrical signals. When light irradiates on drug molecules, the photoacoustic effect will occur. The drug absorbs light energy and expands, thus generating acoustic waves, and these acoustic waves can be captured and transmitted by the transducer;

[0052] The light source is two pulsed lasers, and the two pulsed lasers respectively generate two light pulses with the same pulse repetition frequency and different central wavelengths, namely the pump pulse laser and the probe pulse laser.

[0053] The digital delay generator is externally triggered by the ultrasonic data acquisition platform to control the pulse output of each channel. These output pulses are used to control the relative time delay between the two pulsed lasers, that is, the time difference between the pump light and the probe light, so as to ensure the synchronization of laser output and photoacoustic signal acquisition. Specifically as follows:

[0054] (1). When obtaining S pummp+probe，Δt , control the probe pulse to arrive with a tiny time delay after the pump pulse reaches the tissue, and the moment when the probe pulse is output is the moment when the ultrasonic data acquisition platform acquires the signal;

[0055] (2). When obtaining S pump，Δt , appropriately delay the probe pulse on the original basis;

[0056] (3). When obtaining S probe , appropriately delay the pump pulse on the original basis.

[0057] The energy regulator is placed at the front end of the pump light source outlet and is used to adjust the intensity of the pump light source. The optical fiber is used to couple the output signals of the two pulsed lasers to ensure that the two beams of light irradiate on the same position of the tissue.

[0058] The computer is used to receive and process ultrasonic data, perform image reconstruction, analysis, and final blood drug concentration estimation. The computer can control the operation of the system and output the final monitoring results.

[0059] The ultrasonic data acquisition platform starts the digital delay generator through a trigger signal, controls two pulsed lasers to alternately emit pump light and probe light. After the intensity of the pump light is adjusted by an energy regulator, it is coupled with the probe light through an optical fiber and focused on the same position of the target tissue. The drug molecules in the tissue absorb the light energy and expand, generating a photoacoustic signal. This signal propagates in the form of sound waves and is captured by the ultrasonic transducer.

[0060] The ultrasonic transducer converts the received acoustic signal into an electrical signal, transmits it to the ultrasonic data acquisition platform for digital processing, and the acquisition platform transmits the data to a computer. The computer performs image reconstruction and analysis, extracts the pump light signal, probe light signal and their superimposed signal respectively in combination with the time delay data, and through the analysis and comparison of these signals, realizes the accurate estimation of blood drug concentration and generates the final monitoring result. During the whole process, the laser output and the photoacoustic signal acquisition are strictly synchronized to ensure the accuracy and timeliness of the data.

[0061] The ultrasonic data acquisition platform starts the digital delay generator through a trigger signal, controls two pulsed lasers to alternately emit pump light and probe light. After the intensity of the pump light is adjusted by an energy regulator, it is coupled with the probe light through an optical fiber and focused on the same position of the target tissue. The drug molecules in the tissue absorb the light energy and expand, generating a photoacoustic signal. This signal propagates in the form of sound waves and is captured by the ultrasonic transducer. The ultrasonic transducer converts the received acoustic signal into an electrical signal, transmits it to the ultrasonic data acquisition platform for digital processing, and the acquisition platform transmits the data to a computer. The computer performs image reconstruction and analysis, extracts the pump light signal, probe light signal and their superimposed signal respectively in combination with the time delay data, and through the analysis and comparison of these signals, realizes the accurate estimation of blood drug concentration and generates the final monitoring result. During the whole process, the laser output and the photoacoustic signal acquisition are strictly synchronized to ensure the accuracy and timeliness of the data.

[0062] The present invention also provides a non-invasive quantitative blood drug concentration monitoring method, including the following steps:

[0063] (1). By differentiating the photoacoustic signals generated by the probe light with and without the pump light, high-specificity molecular detection can be achieved. Therefore, the TTD signal of the probe molecule can be obtained using the following formula:

[0064] S TTD，Δt =S pup+probe，Δt -S pump，Δt -S probe (1)

[0065] Wherein, S TTD，Δt is the TTD signal obtained by removing the background through differentiation, S pup+probe，Δt is the photoacoustic signal when the pump light and the probe light are simultaneously excited, Spump，Δt is the photoacoustic signal generated only by the pump light excitation, S probe is the photoacoustic signal generated only by the probe light excitation, and Δt represents the delay time of the probe laser relative to the pump laser;

[0066] (2). Under the excitation of the pump light, the molecule transitions from the ground state (S0) to the first excited singlet state (S1). Due to the spin-orbit coupling effect, the molecule undergoes intersystem crossing and enters the first excited triplet state (T1). Since the transition from the T1 state to the ground state S0 is spin-forbidden, the T1 state has a very long lifetime (τ ∼ 10 -6 −10 s). The molecules in the T1 state can be excited to the T2 state by the probe light, then return to the T1 state through non-radiative relaxation and generate photoacoustic signals. Therefore, by writing the rate equations of the populations N1, N2, and N3 in the S0, S1, and T1 states during the pump excitation process, we can obtain:

[0067]

[0068] Among them, A 21 represents the spontaneous emission coefficient from S1 to S0, B is the light absorption coefficient, I(t) is the laser intensity distribution, A 23 represents the intersystem crossing rate from the S1 state to the T1 state. Since the T1 state has a long lifetime, the spontaneous emission from the T1 state to the ground state S0 is not considered.

[0069] (3). Since the lifetime of the S1 state is short, significantly shorter than the typical laser pulse width of about 5 ns, assuming that the population of the S1 state does not accumulate and is in a transient equilibrium state, that is, N2 ≈ 0, and dN 2 / dt = 0. Therefore, the following formula can be obtained:

[0070]

[0071] Based on the above, it can also be obtained that:

[0072]

[0073] Among them, N 0 is the initial number of particles in the S0 state, T is the pulse duration, Φ P is the pulse energy of the pump laser. Therefore, it can be seen that due to the existence of the metastable state T1, population inversion can be achieved between the T1 state and the S0 state. The population N 3 of the T1 state increases exponentially with the pump laser energy Φ P and gradually saturates. In contrast, since the lifetime of the T2 state is short, the molecule may experience multiple excitation-relaxation cycles. Therefore, the probe laser will not affect N 3Produces significant changes. Therefore, the amplitude of the TTD signal is proportional to N 3 and the energy Φ of the probe laser T and can be written as:

[0074]

[0075] where C 0 is the concentration of the target molecule, Φ P0 is the value when the pump pulse energy reaches 63% saturation, A TTD，Δt is the TTD amplitude at full saturation, μ a is the absorption coefficient, η th is the heat conversion efficiency, Γ is the Grueneisen parameter, Δt is the delay time of the pump laser relative to the probe laser, τ is the lifetime of the triplet state, and Z us represents the acoustic influence factor, including the acoustic scattering of the tissue, the size of the blood vessel, and the detection angle of the inclination;

[0076] (4). Assuming that the photoacoustic signal mainly comes from hemoglobin, then it will have a linear relationship with the energy of the probe laser. Therefore, formula (8) can be rewritten as:

[0077]

[0078] where A BV is the photoacoustic amplitude of the probe laser in the target blood vessel, K is a constant, which can be calibrated by known samples of C 0 and hemoglobin concentration. Therefore, the energy of the pump laser can be adjusted, and the maximum ratio R max can be obtained by fitting formula (9), and then the concentration of the target molecule can be determined by C 0 = R max / K.

[0079] As Figure 1 shown, the triplet state is usually when the drug molecule is excited by the pump laser and transitions to the singlet excited state, and then enters the triplet state through a non-radiative transition process. Therefore, the absorption spectrum of the drug molecule is selectively changed in response to the pump laser irradiation, allowing the probe laser to detect the drug molecule without background interference through signal difference.

[0080] The TTD signal of the triplet drug molecule increases exponentially with the intensity of the pump laser and linearly with the intensity of the probe laser. The intensity of the pump laser determines the proportion of drug molecules excited to the triplet state, and the drug concentration is proportionally quantified using the vascular detection photoacoustic signal as a reference. Therefore, this method can overcome the challenge of inaccurate optical molecular measurements caused by strong and variable light scattering in tissues, thereby realizing non-invasive quantitative monitoring of the concentration of specific target molecules.

[0081] It should be noted that, on the premise of no conflict, in order to verify the effect of the present invention, relevant phantom experiments were carried out, which proved that the method can accurately quantify the drug concentration under different light scattering conditions.

[0082] Example 1

[0083] S1. In the experiment, methylene blue (MB) was used as the target drug molecule. MB is an FDA-approved water-soluble dye, which is widely used in clinical and biomedical research. Its T1 lifetime in water is about 2 μs, and its maximum absorption peaks in the S0 and T1 states are about 650 nm and 820 nm respectively. Therefore, these two wavelengths were set as the pump light and probe light wavelengths respectively, and the delay time Δt was 0.3 μs.

[0084] S2. S pup+probe，Δt is the photoacoustic signal obtained when the pump and probe lasers are turned on simultaneously, and S pump，Δt is the photoacoustic signal generated by only the pump light excitation, and S prove is the photoacoustic signal obtained when only the probe laser is turned on. The three data acquisition sequences were obtained using an interleaved method and averaged 300 times to improve the signal-to-noise ratio.

[0085] S3. To obtain the pump intensity dependence of the TTD signal, the pump laser intensity was controlled by an energy regulator, and data were collected at about 10 different laser energy levels. As Figure 4 (a) shows, as the pump laser intensity increases, the intensities of the pump image and the TTD image increase, while the probe image remains stable.

[0086] S4. A transparent plastic tube with a diameter of 3 mm and negligible wall thickness was used to simulate a blood vessel. The tube was filled with 12% hemoglobin solution and different concentrations of MB. In the first phantom experiment, the changes in the TTD signal under different pump and probe light intensities were monitored. Through the signal peak value of the image, the relationship between TTD / TR(Φ P ) and the pump laser intensity was plotted. As Figure 4 (b) shows, initially, as the pump laser intensity Φ P increases, this ratio rises rapidly and then stabilizes at the saturation value R max . As Figure 4 (c) shows, the increase in R max depends on the MB concentration, so the MB concentration in the blood vessel can be determined by measuring R max .

[0087] S5. The energy regulator was placed in front of the probe laser, and the TTD signal was measured when the probe laser intensity changed. As Figure 4As shown in (d), the TTD signal increases linearly with the detection laser intensity. In short, the TTD signal increases exponentially with the pump laser intensity and linearly with the detection laser intensity;

[0088] Example 2

[0089] S1. In the second phantom experiment, two tubes with different MB concentrations were placed under a 2-mm-thick chicken breast tissue slice to simulate the light attenuation caused by the tissue and evaluate the compensation ability of the method in this study for the unknown light attenuation from the tissue. Figure 5 (a) and 5(b) show the ultrasound and photoacoustic images of two closely arranged tubes containing 100 and 400 μM MB, respectively. Figure 5 (c) shows the variation of the TTD / detection ratio with the pump laser intensity;

[0090] S2. By fitting these two curves, the R max values of the 400 μM and 100 μM MB tubes were determined to be 0.2707 and 0.084, respectively. Figure 5 (d)-(f) show the corresponding results when the tubes are covered with a 2-mm-thick chicken breast tissue. As can be seen from Figure 5 (e), due to the light attenuation of the chicken breast tissue, the photoacoustic amplitude decreases significantly. Figure 5 The R max value fitting results in (f) are 0.2326 and 0.0783, corresponding to the 400 μM and 100 μM MB tubes, respectively.

[0091] S3. Based on the above results, it was confirmed that the method in this study can quantitatively and accurately evaluate the blood drug concentration in deep tissues without prior knowledge of the tissue scattering characteristics. It should be noted that as the MB concentration increases, the rising speed of R max gradually slows down. As shown in Figure 4 (b), this may be due to the formation of dimers at high MB concentrations, resulting in a decrease in the effective concentration. Therefore, in Figure 5 (c) and 5(f), although the R max values of the 400 μM and 100 μM tubes only differ by about three times, these values are consistent with the results in Figure 4 (b). This consistency further verifies the high repeatability and reliability of the method.

[0092] The above embodiments are only the preferred embodiments of the present invention and cannot be used to limit the scope of protection of the present invention. Any non-substantial changes and substitutions made by those skilled in the art based on the present invention fall within the scope of protection required by the present invention.

Claims

1. A non-invasive quantitative blood drug concentration monitoring system, characterized in that: include Ultrasonic data acquisition platform, ultrasonic transducer, light source, digital delay generator, energy regulator, optical fiber and computer; The ultrasonic data acquisition platform is used to receive signals from the ultrasonic transducer and convert them into digital data for subsequent analysis and processing, and at the same time output a trigger signal to the digital delay generator to control the two light sources to perform automatic interleaved data acquisition to ensure synchronization between laser output and photoacoustic signal acquisition; The ultrasonic transducer can convert the acoustic wave signal generated by the photoacoustic effect into an electrical signal. When light is irradiated onto drug molecules, a photoacoustic effect is generated, and the drug absorbs the light energy and expands, thereby generating acoustic waves, which can be captured and transmitted by the transducer; The light sources are two pulse lasers, which respectively generate two light pulses with the same pulse repetition frequency and different central wavelengths, namely a pump pulse laser and a detection pulse laser.

2. The non-invasive quantitative blood drug concentration monitoring system according to claim 1, characterized in that: The digital delay generator is controlled by an external trigger from the ultrasonic data acquisition platform to output pulses of each channel. The output pulses are used to control the relative delay of the two pulse lasers, that is, the time difference between the pump light and the detection light, to ensure the synchronization of the laser output and the photoacoustic signal acquisition, as follows: (1) Obtain S pummp+probe，Δt When the detection pulse is controlled to arrive a few microseconds after the pump pulse arrives at the tissue, the detection pulse is output at the same time as the ultrasonic data acquisition platform collects the signal; (2) Obtain S pump，Δt When the detection pulse is properly delayed on the original basis; (3) Obtain S probe When the pump pulse is delayed appropriately on the original basis.

3. The non-invasive quantitative blood drug concentration monitoring system according to claim 1, characterized in that: The energy regulator is placed at the front end of the pump light source outlet to adjust the intensity of the pump light source. The optical fiber is used to couple the output signals of two pulse lasers to ensure that the two beams of light irradiate the same position of the tissue.

4. The non-invasive quantitative blood drug concentration monitoring system according to claim 1, characterized in that: The computer is used to receive and process ultrasound data, perform image reconstruction, analysis and final blood drug concentration estimation. The computer can control the operation of the system and output the final monitoring results.

5. The non-invasive quantitative blood drug concentration monitoring system according to claim 1, characterized in that: The ultrasonic data acquisition platform starts the digital delay generator through a trigger signal, controls two pulse lasers to emit pump light and detection light alternately, and after the intensity of the pump light is adjusted by the energy regulator, it is coupled with the detection light through the optical fiber and focused on the same position of the target tissue. The drug molecules in the tissue absorb the light energy and expand to generate a photoacoustic signal, which is transmitted in the form of sound waves and captured by the ultrasonic transducer.

6. The non-invasive quantitative blood drug concentration monitoring system according to claim 1, characterized in that: The ultrasonic transducer converts the received sound wave signal into an electrical signal and transmits it to the ultrasonic data acquisition platform for digital processing. The acquisition platform transmits the data to a computer, which performs image reconstruction and analysis, and extracts the pump light signal, the detection light signal and the superposition signal of the two in combination with the time delay data. By analyzing and comparing these signals, accurate estimation of blood drug concentration is achieved and the final monitoring results are generated. During the whole process, the laser output and the photoacoustic signal acquisition are strictly synchronized to ensure the accuracy and timeliness of the data.

7. A non-invasive quantitative blood drug concentration monitoring method, characterized in that: The method comprises the non-invasive quantitative blood drug concentration monitoring system as claimed in claim 1 and the following steps: (1) Highly specific molecular detection is achieved by differentiating the photoacoustic signal generated by the probe light with and without pump light. Therefore, the TTD signal of the probe molecule can be obtained using the following formula: S TTD，Δt =S pup+probe，Δt -S pump，Δt -S probe (1) Among them, S TTD，Δt is the TTD signal obtained by subtracting the background, S pup+probe，Δt is the photoacoustic signal when the pump and probe light are excited simultaneously, S pump，Δt is the photoacoustic signal generated by pump light excitation only, S probe Only the photoacoustic signal generated by optical excitation is detected, and Δt represents the delay time of the detection laser relative to the pump laser; (2) Under the excitation of pump light, the molecule transitions from the ground state (S0) to the first excited singlet state (S1). Due to the spin-orbit coupling effect, the molecule undergoes intersystem crossing and enters the first excited triplet state (T1). Since the transition from the T1 state to the ground state S0 is spin-forbidden, the T1 state has a very long lifetime (τ~10 -6 -10s), the molecules in the T1 state can be excited to the T2 state by the detection light, and then return to the T1 state by non-radiative relaxation, and generate photoacoustic signals. Therefore, by writing the rate equations of the number of particles N1, N2 and N3 in the S0, S1 and T1 states during the pump excitation process, we can get: Among them, A 21 represents the spontaneous radiation coefficient from S1 to S0, B is the light absorption coefficient, I(t) is the laser intensity distribution, A 23 represents the intersystem crossing rate from the S1 state to the T1 state. Since the T1 state has a long lifetime, the spontaneous radiation from the T1 state to the ground state S0 is not considered.

8. The non-invasive quantitative blood drug concentration monitoring method according to claim 7, characterized in that: The following steps are also included: (3) Since the lifetime of the S1 state is short, significantly shorter than the typical laser pulse width of about 5ns, assuming that the particles in the S1 state do not accumulate and are in a transient equilibrium state, that is, N2≈0, and dN2 / dt=0, the following formula can be obtained: Based on the above, we can also get: Where N0 is the initial number of particles in the S0 state, T is the pulse duration, Φ P is the pulse energy of the pump laser. Therefore, it can be seen that due to the existence of the metastable state T1, the energy level inversion can be achieved between the T1 state and the S0 state. The particle N3 in the T1 state increases with the pump laser energy Φ P In contrast, due to the short lifetime of the T2 state, the molecule may undergo multiple excitation-relaxation cycles, so the detection laser will not cause significant changes to N3 in the T1 state. Therefore, the amplitude of the TTD signal is related to N3 and the energy of the detection laser Φ T is proportional to, and can be written as: Where C0 is the concentration of the target molecule, Φ P0 is the value when the pump pulse energy reaches 63% saturation, A TTD，Δt is the TTD amplitude at full saturation, μ a is the absorption coefficient, η th is the thermal conversion efficiency, Γ is the Grueneisen parameter, Δt is the delay time of the pump laser relative to the probe laser, τ is the lifetime of the triplet state, and Z us Represents acoustic influencing factors, including acoustic scattering of tissue, vessel size, and tilted detection angle; (4) Assuming that the photoacoustic signal mainly comes from hemoglobin, it will be linearly related to the energy of the detection laser. Therefore, formula (8) can be rewritten as: Among them, A BV is the photoacoustic amplitude of the detection laser in the target blood vessel, K is a constant that can be calibrated by samples with known C0 and hemoglobin concentration. Therefore, the energy of the pump laser can be adjusted, and the maximum ratio R can be obtained by fitting formula (9): max , then through C0 = R max / K to determine the concentration of the target molecule.

9. The non-invasive quantitative blood drug concentration monitoring method according to claim 7, characterized in that: The triplet state is usually when the drug molecule transitions to the singlet excited state after being excited by the pump laser, and then enters the triplet state through a non-radiative transition process. Therefore, the absorption spectrum of the drug molecule is selectively changed in response to the pump laser irradiation, thereby allowing the probe laser to detect the drug molecule through signal differential without background interference.

10. The non-invasive quantitative blood drug concentration monitoring method according to claim 7, characterized in that: The TTD signal of the triplet drug molecule increases exponentially with the pump laser intensity and increases linearly with the detection laser intensity. The intensity of the pump laser determines the proportion of the drug molecules excited to the triplet state, and the drug concentration is proportionally quantified using the vascular detection photoacoustic signal as a reference. Therefore, this method can overcome the challenge of inaccurate optical molecular measurement caused by strong and variable light scattering in tissues, thereby achieving non-invasive quantitative monitoring of the concentration of specific target molecules.

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