WO2022165863A1 - 一种用于优化连续流人工心脏脉动工作模式的内皮细胞体外培养系统 - Google Patents
一种用于优化连续流人工心脏脉动工作模式的内皮细胞体外培养系统 Download PDFInfo
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Definitions
- the invention belongs to the technical field of artificial organs, and relates to an endothelial cell culture chamber and an extracorporeal circulation system for optimizing the pulsation working mode of an artificial heart.
- a miniature extracorporeal simulated circulatory system in which changes in hemodynamic signals induced by different pulsatile working modes of continuous-flow artificial heart pump speed affect endothelial cell function.
- Artificial heart is the current non-drug mechanical treatment and rehabilitation method for end-stage heart failure.
- the continuous flow artificial heart has been greatly promoted and applied due to its advantages of small size, high reliability, easy implantation and operation.
- Its main body is an impeller blood pump, which outputs a steady flow when the impeller rotates at a constant speed, and outputs a pulsating flow when the rotational speed changes periodically.
- the impeller pump is usually set to a constant speed, and this working mode will significantly reduce the pulsatility of arterial blood flow and blood pressure, resulting in dysfunction of vascular endothelial cells, thereby inducing arteriovenous malformation and hemorrhagic stroke. , as well as a large number of adverse events such as kidney and other organ damage.
- endothelial cell function such as vascular vasodilator nitric oxide (Nitric Oxide, NO) and vasoconstrictor factor endothelin-1 (Endothelin-1, ET-1) and other blood vessels
- vascular vasodilator nitric oxide Nitric Oxide, NO
- vasoconstrictor factor endothelin-1 Endothelin-1, ET-1
- TNF- ⁇ tumor necrosis factor- ⁇
- interleukin-6 interleukin-6
- interleukin-8 interleukin-8
- the amplitude adjustment of the waveform of the periodic pump speed change is mainly based on the hemodynamic mechanism of the coupling effect of the ventricle and afterload.
- the frequency adjustment includes synchronous adjustment (the frequency of the pump speed change is consistent with the heart rate) and asynchronous adjustment (the frequency of the pump speed change is independent of the heart rate). It is expected to improve the pulsatility of the hemodynamic signal by changing the Vascular endothelial function, reducing the incidence of adverse events in peripheral blood vessels and organs.
- microfluidic chips have the advantages of less sample required, easy integration, easy optical detection, and good biological adaptability.
- ECCM in vitro endothelial cell culture model
- cell Mechanobiology Research System the ECCM currently established to study the influence of the hemodynamic microenvironment corresponding to the different pulsatile working modes of the artificial heart pump on the function of the arterial endothelial cells cannot truly reproduce the real human arterial endothelial microenvironment under the action of the artificial heart. hemodynamic signals.
- the purpose of the present invention is to provide a method capable of truly simulating the blood pressure, wall shear stress and stretch strain (stress) signals in the arterial endothelial hemodynamic microenvironment caused by the pulsatile working mode of the artificial heart.
- This method cleverly combines hemodynamic principles, microfluidic chip technology and intelligent feedback control technology.
- the multi-component lumped parameter model builds an in vitro fluid simulation circulatory system to reproduce the combined effects of pressure, shear stress and stretch strain on body arterial endothelial cells under different pulsatile working modes of pump speed after implantation of an artificial heart in patients with heart failure. It can be used to study the quantitative relationship between hemodynamic signals and mechanobiological effects of arterial endothelial cells and its molecular biological mechanism.
- An in vitro culture system for endothelial cells for optimizing continuous flow artificial heart pulsation working mode, including three basic units: the first basic unit is the cell culture chamber on the microfluidic chip and the off-chip Multi-element aortic arch afterload hydrodynamic loop (shown in Figure 2); off-chip multi-element aortic arch afterload hydrodynamic loop includes a flu, resistance valve, elastic cavity 1, elastic cavity 2 in series with the cell culture cavity; elastic cavity 1.
- the elastic chambers 2 are respectively arranged on both sides of the cell culture chamber.
- the second basic unit is a pulsatile fluid loading device and an artificial heart device that simulate the power source of the cardiovascular system.
- the fluid loading device is realized by a pulsed blood pump (Q 1 (t) in Figure 3), which can Simulate the waveforms of blood pressure, wall shear stress and stretch strain on the body arterial endothelial cells of normal and heart failure patients;
- the artificial heart device is connected in parallel to both ends of the pulsed blood pump (Q 2 (t) in Figure 3), and then Connecting the two in series into the above-mentioned hydrodynamic circulation loop can simulate the blood pressure, wall shear stress and stretch strain signals under different pump speed pulsatile modulation modes under the body arterial endothelial cells;
- the third basic unit is a peripheral detection and feedback control system, as shown in Figure 1, including an inverted fluorescence microscope, a CCD high-speed camera system, a pressure sensor and a flow sensor, and a proportional-integral-derivative feedback control system, a pressure sensor and a flow rate sensor.
- the sensors are arranged on both sides of the cell culture chamber, and are used to monitor and collect the pressure and flow waveforms at the input and output ends of the cell culture chamber in real time.
- the fluorescence microscope is located above the cell culture chamber Rc, and the CCD high-speed camera system is connected to the fluorescence microscope.
- the CCD high-speed camera system, pressure sensor and flow sensor are all connected with the proportional-integral-derivative feedback control system.
- the acquisition and feedback of cell morphological structure data, the proportional-integral-derivative feedback control system can quantitatively control the changes of related hemodynamic signals, in the microfluidic chip cell culture chamber to generate different pulsation working modes based on artificial heart pump speed in vivo Combined effects of pressure, shear stress, and stretch-strain signals on arterial endothelial cells.
- the above endothelial cell culture cavity is a cavity with a concave cross-section, and an elastic film similar to the elastic modulus of the artery is bonded to a hard light-transmitting polymethyl methacrylate (PMMA) material
- PMMA polymethyl methacrylate
- the cell culture cavity below the lower surface of the elastic film is filled with circulating liquid; air is introduced into the cavities on both sides of the upper surface of the elastic film to provide sufficient space for the deformation of the film on both sides of the cavity under the action of pulsating fluid pressure.
- the upper surface of the elastic film has smooth circular arcs at both ends, and the middle of the upper surface of the film is close to the inner surface of the cavity in the horizontal direction, so that the elastic film attached to the endothelial cells below can be pulled on both sides.
- the choice of the elastic modulus of the elastic film should be determined according to the principle of elasticity and by accurately simulating the actual needs of blood pressure, shear stress and stretch strain waveforms in the endothelial microenvironment of different parts of the aorta.
- the cell culture chamber provides circulating liquid for the cells in the cell culture chamber through the cooperation of the one-way valve and the liquid reservoir.
- cardiopulmonary bypass system can be equivalent to a circuit model, wherein: the flow resistance of the endothelial cell culture chamber is equivalent to resistance (Rc in Figure 3), and the compliance of the film on the culture chamber is equivalent to capacitance (Figure 3).
- C 1 the compliance, flow resistance and inductance of the vascular bed downstream of the aortic arch are equivalent to capacitance, resistance and inductance (C 2 , R and L in Figure 3).
- the design of the off-chip multi-component aortic arch afterload hydrodynamic loop requires that the pressure, wall shear stress, and stretch strain of the endothelial cells cultured on the cell culture cavity membrane are related to the corresponding parts of the endothelium of the heart failure patient after implantation of the artificial heart.
- the waveforms of blood pressure, shear stress, and stretch strain experienced by cells are consistent:
- the waveforms of blood pressure p(t), wall shear stress ⁇ ⁇ (t) and stretch strain ⁇ (t) obtained by human or animal experiments in the vicinity of body arterial endothelial cells were simulated as targets.
- the blood pressure and shear stress waveforms of endothelial cells cultured on the luminal membrane are equal to the blood pressure and wall shear stress in the body arterial endothelial microenvironment, and the blood flow q(t) and pressure drop ⁇ p(t) must satisfy:
- ⁇ is the viscosity of the cell culture medium
- Hc, Wc and Lc are the height, width and length of the cell culture chamber, respectively.
- the aortic arch afterload hemodynamic behavior is equivalent to a circuit system.
- the input impedance of the hydrodynamic circuit is expressed as the ratio of the input pressure waveform p(t) and the blood flow waveform q(t) in the frequency domain, and is characterized by the amplitude and phase of the blood pressure and blood flow harmonic components corresponding to the angular frequency ⁇ n :
- a multi-component extracorporeal fluid simulation circulatory system was constructed to simulate the afterload hemodynamic characteristics of the aortic arch according to the numerical values of influenza L, resistance valve R, elastic cavity C 1 and elastic cavity C 2 .
- the circulating fluid in the system is in vitro vascular endothelial cell culture fluid
- the elastic lumen simulates arterial compliance (flow volume)
- the resistance valve simulates viscous resistance (flow resistance)
- the flu element simulates flow inertia.
- the above-mentioned device for simulating the power source of the cardiovascular system is realized by connecting the artificial heart to both ends of the pulsatile fluid loading device in parallel.
- the pulsatile fluid loading device combined with the PID feedback control device can be used to simulate the blood pressure, wall shear stress and stretch strain signals in the hemodynamic microenvironment of the body arterial endothelial cells under normal and physiological conditions of heart failure, and the artificial heart can be simulated.
- the device is connected in parallel with the fluid loading device, and then the two are connected in series to the above-mentioned hydrodynamic circulation loop, and combined with the PID feedback control device, the artificial heart pump speed can be generated under different pulsatile working modes.
- the fluid dynamics signal waveform feeds the collected signal back to the PID control device, which can further adjust the fluid loading device and artificial heart, thereby quantitatively regulating the pressure, flow signal amplitude and frequency changes acting on the multi-element simulated circulatory system.
- the cell culture chamber of the fluidic chip produces the combined effects of blood pressure, shear stress and stretch strain under different pulsatile working modes based on the pumping speed of the artificial heart.
- Step 1 The primary cultured endothelial cells were subcultured, and the second to fifth passages were used for the experiment. Adjust the values of influenza L, resistance valve R, elastic cavity C 1 , and elastic cavity C 2 in the extracorporeal simulated circulatory system, and pass circulating liquid into the cell culture cavity. Under different pulsatile working modes of artificial heart pumping speed, the combined stimulation of various hemodynamic signals for arterial endothelial cells is loaded.
- Step 2 Continue to load the hemodynamic signal stimulation corresponding to the above working mode, and then perform activity detection on the cells to ensure the effectiveness of the above system.
- Step 3 Collect cell samples from the microfluidic chip cell culture chamber, and detect gene and protein expression levels to analyze blood pressure, shear stress, and stretch strain caused by different pulsatile working modes of artificial heart pumping speed. Effects of hemodynamic signals on gene and protein expression levels of vasoactive substances and pro-inflammatory cytokines.
- the present invention can successfully reproduce the blood pressure, wall shear stress and stretch strain signals corresponding to different pulsatile working modes of the artificial heart pump speed, and uses a more integrated and less consumable
- the microfluidic chip cell culture chamber studies the differential effects of the combined stimulation of hemodynamic signals on the function of arterial endothelial cells under the above working modes; it provides an efficient method for studying the quantitative relationship between hemodynamic signals and the function of arterial endothelial cells. It is a reasonable experimental platform and provides a scientific basis for screening out the continuous flow artificial heart pumping mode that is more conducive to improving and maintaining normal endothelial function.
- Figure 1 is a schematic diagram of the structure of the in vitro endothelial cell culture chamber and the peripheral monitoring system.
- Figure 2 is a schematic diagram of the hydrodynamic circulation loop of the in vitro endothelial cell culture chamber.
- Figure 3 Schematic diagram of the equivalent circuit model of aortic arch afterload hemodynamic behavior.
- Figure 4 is a schematic diagram of the cell culture chamber of the microfluidic stretch chip.
- Figure 5 shows the blood pressure and shear stress waveforms of vascular endothelial cells in the aortic arch obtained through in vivo experiments in normal, heart failure and asynchronous modulation modes, and the blood pressure in the culture cavity obtained by inverse solution based on the shear stress waveform and the cell culture cavity size.
- Figure 6 shows the fitting results of the actual input impedance amplitude and phase angle using Matlab/Simulink for the above-mentioned equivalent circuit model under the three modes of normal, heart failure and asynchronous modulation respectively, and the blood flow information is used as the current excitation,
- the output voltage of the model is used as the simulation target, and the schematic diagram of the results compared with the blood pressure in Figure 5;
- (a-1) the amplitude-frequency curve of the input impedance under normal physiological conditions;
- (a-2) the input impedance under normal physiological conditions The phase angle-frequency curve of the The phase angle-frequency curve of the input impedance under the condition of heart failure;
- (b-3) the comparison diagram of the model output voltage and blood pressure under the condition of heart failure,
- (c-1) the amplitude of the input impedance under the asynchronous pulsatile working mode of the artificial heart pumping speed - Frequency curve;
- (c-2) Phase angle-frequency curve of input impedance in asynchronous pulsati
- A(ii) and B(ii) are the pressure and flow sensors located at both ends of the microfluidic chip cell culture chamber;
- Rc is the microfluidic chip Flow resistance of cell culture chamber;
- C 1 is film compliance;
- R is flow resistance of connecting pipe;
- C 2 is elastic air cavity that characterizes compliance;
- L is flu of connecting pipe during liquid circulation.
- Design microfluidic chip cell culture chamber height Hc, width Wc and length Lc are respectively 0.3mm, 6mm and 15mm, and the viscosity ⁇ of cell culture fluid is usually 0.001Pa s;
- Three target input impedances z( ⁇ i ) are calculated respectively for the target blood pressure and blood flow;
- the hemodynamic characteristics of the extracorporeal simulated circulatory system can be characterized by the five-element equivalent circuit model shown in Figure 3.
- the input impedance of the circuit can be can be expressed as:
- Equation 4 the equivalent input impedance of the five-element lumped parameter model Combined with the target input impedance z( ⁇ i ), the parameter values of each element can be obtained through the system identification method.
- the input impedance curve corresponding to the five-element lumped parameter model basically agrees with the target input impedance curve (circle in FIG. 6 ).
- the blood pressure waveform obtained by Matlab/simulink simulation is basically consistent with the corresponding blood pressure waveform in Figure 5, as shown in Figure 6,
- the root mean square error is 0.237, 0.401 and 0.625 in turn; then the hydrodynamic circulation loop based on the microfluidic chip cell culture chamber is constructed as shown in Figure 2;
- the chip is manufactured by a standardized micromachining method.
- An elastic film similar to the elastic modulus of the artery is bonded to a cavity made of hard light-transmitting PMMA material.
- the cross-section of the cavity is concave, and the elastic film is
- the cell culture cavity below the surface is filled with circulating liquid; air is introduced into the cavities on both sides of the upper surface of the elastic membrane to provide sufficient space for the membrane on both sides to deform under the action of pulsating fluid pressure; the middle cavity on the upper surface of the elastic membrane
- Both ends of the membrane are smooth arc-shaped, and the middle part of the upper surface of the membrane is close to the inner surface of the cavity in the horizontal direction, which can make the elastic membrane attached to the endothelial cells below only produce horizontal stretch strain under the action of pulling on both sides.
- the thickness of the cavity cavity should be designed to ensure that the endothelial cells can be focused when observing the morphology and structure of the endothelial cells with a microscope, and will not be deformed under the action of pulsating pressure;
- the principle of elasticity is determined by accurately simulating the actual needs of blood pressure, shear stress and stretch strain waveforms in the endothelial microenvironment of different parts of the aorta;
- the signal acquisition and processing system (ii) includes an inverted fluorescence microscope, a CCD high-speed camera system, a pressure sensor and a flow sensor for real-time monitoring and acquisition of pressure and flow waveforms at the input end A and output end B of the cell culture chamber, as well as microfluidic control The actual morphological structure of cells in the cell culture chamber on a chip.
- the pulsed blood pump (i) combined with the PID feedback control device (iii) can accurately simulate the physiological conditions of normal and heart failure. Simulate the blood pressure, wall shear stress and stretch strain signals of vascular endothelial cells in specific parts of the aortic arch under different pulsatile working modes of continuous flow artificial heart pumping speed, and finally load quantitative and controllable pulsatile flow signals to the multi-element simulated circulatory system; The signal is fed back to the PID control device (iii), which can further adjust the pulsed blood pump (i) and the artificial heart (i), thereby quantitatively regulating the changes in the pressure, flow signal amplitude and frequency acting on the multi-element simulated circulatory system, and finally The combined effects of blood pressure, shear stress, and stretch strain under different pulsatile working modes based on the pumping speed of the artificial heart are generated in the cell culture chamber of the microfluidic chip.
- the pressure in the microfluidic chip can be measured by the pressure sensor; the shear stress can be calculated from the flow waveform measured by the flow sensor and the geometric size of the cell culture chamber; and for the horizontal direction generated by the elastic film in the chip under different pressures.
- the stretch strain can be first calibrated on the film with fluorescent microspheres and measured by fluorescence microscopy.
- the tensile strain corresponding to the elastic film is obtained by giving different pressures, and then the relational expression between the pressure and the tensile strain is established. According to this approximate expression, the tensile strain of the elastic film in the actual experiment is determined under the condition that the pressure is known.
- the morphological structure of endothelial cells was detected and recorded by a microscope combined with a CCD high-speed camera system and saved to an industrial computer.
- step 1 the primary cultured endothelial cells are subcultured with EGM medium, and passages 2-5 are used for experiments.
- endothelial cells were planted on the elastic membrane of the cell culture chamber of the microfluidic chip coated by Fibronection, so that the cells adhered to the wall and the confluency reached more than 90%.
- step 2 the arterial endothelial cells are loaded with a combination of hemodynamic signals corresponding to different pulsatile working modes of the artificial heart pump; NucViewTM-488 cell activity detection reagent is used to detect cell activity to ensure the effectiveness of simulating the circulatory system in vitro.
- Step 3 Collect cell samples from the microfluidic chip cell culture chamber to detect gene and protein expression levels, so as to obtain the blood pressure, shear stress and stretch strain corresponding to the different pulsatile working modes of the artificial heart pump speed. Effects of hydrodynamic signaling on gene and protein expression levels of vasoactive substances and proinflammatory cytokines.
- the invention can successfully reproduce the blood pressure, wall shear stress and stretch strain signals borne by the body arterial endothelial cells under different pulsating working modes of the artificial heart pumping speed, and can simulate the arterial endothelium cultured under the combined stimulation of the above-mentioned hemodynamic signals. Differential effects of cellular function are monitored in real time.
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Abstract
一种用于优化连续流人工心脏脉动工作模式的内皮细胞体外培养系统，属于人工器官技术领域。该系统包括三部分：1)微流控芯片上的细胞培养腔及芯片外多元件主动脉弓后负荷流体力学循环回路。2)模拟心血管系统动力源的装置：流体加载装置由脉冲式血液泵实现；人工心脏装置并联接入到脉冲式血液泵的两端。3)外围检测与反馈控制系统，包括压力、流量传感器，荧光显微镜，CCD高速摄像系统及比例-积分-微分反馈控制系统。该系统可精确模拟真实的主动脉弓不同部位血管内皮细胞血流动力学微环境，为研究人工心脏泵速的不同脉动工作模式与局部动脉内皮微环境血流动力学信号之间的定量关系提供微型化、客观化、标准化和定量化的实验平台。
Description
本发明属于人工器官技术领域，涉及一种用于优化人工心脏脉动工作模式的内皮细胞培养腔及体外循环系统，是基于血流动力学原理、微流控芯片及智能反馈控制技术，用于研究连续流人工心脏泵速的不同脉动工作模式引起的血流动力学信号改变影响内皮细胞功能的微型体外模拟循环系统。
人工心脏是目前针对终末期心衰的非药物性机械治疗方法和康复手段。连续流人工心脏因其体积小、可靠性高、易于植入和操作等优点得到了极大推广和应用。它的主体是叶轮式血泵，叶轮恒速旋转时输出定常流，转速周期性改变时输出脉动流。临床使用时为了操作方便，通常将叶轮泵设置为恒定转速，而这种工作模式会导致动脉血流和血压的脉动性明显降低，造成血管内皮细胞功能失调，从而诱发动静脉畸形、出血性中风、以及肾脏等其他器官受损等大量不良事件的发生。
近20多年的基础和临床研究发现，人工心脏能够通过血流动力学机制影响动脉内皮功能，从而进一步调节外周血管的结构和功能。血管内皮作为介于血流和血管壁组织之间的一道屏障，位于血管壁的最内层，因此处于复杂的血流动力学微环境中，同时直接承受血流产生的壁面剪应力、血压以及血压导致的血管周向牵张应力(或应变)等血流动力学信号作用。内皮细胞能通过细胞膜表面受体及感受器识别细胞外微环境中的血流动力学信号及其变化，并将力学信号通过一系列的信号通路级联反应传递到细胞内部引起基因和蛋白表达的变化，即力学生物学(mechanobiology)机制，最终影响内皮细胞功能的变化，如 舒血管因子一氧化氮(Nitric Oxide,NO)和缩血管因子内皮素-1(Endothelin-1，ET-1)等血管活性物质的分泌、以及肿瘤坏死因子-α(tumor necrosis factor-α,TNF-α)、白细胞介素-6(interleukin-6,IL-6)、白细胞介素-8(interleukin-8,IL-8)等促炎性细胞因子表达水平的变化。这些物质的短期效应会影响血管壁舒张和收缩功能、引起炎性反应等，长期效应则导致血管结构和功能的重建，即引起血管壁厚、管径大小、血管弹性等发生变化。
目前的研究已尝试将连续流人工心脏的血泵转速设置为各种周期性变化的脉动工作模式，周期性泵速变化波形的幅值调节主要根据心室和后负荷耦合作用的血流动力学机制进行设定，频率调节则包括同步调节(泵速变化的频率与心率一致)和异步调节(泵速变化的频率与心率无关)两种方式，期望通过改变血流动力学信号的脉动性从而改善血管内皮功能，降低外周血管及器官的不良事件发生率。然而，连续流人工心脏泵速的不同脉动工作模式如何影响血流动力学信号的规律及如何差异化调控动脉内皮细胞的功能，迄今为止未有全面、系统的研究，因而限制了人工心脏泵速的不同脉动工作模式在临床的精准实施及在治疗与康复中科学合理的使用。
动物模型与人体临床实验是将连续流人工心脏应用于临床之前开展动脉内皮血流动力学微环境特性分析的最直接方式。然而动物和人体的在体动脉内皮细胞所处的血流动力学微环境非常复杂，且极易受呼吸和神经调节等其他因素的影响，同时存在个体差异大、血流动力学参数监测精度低、成本高、周期长、以及伦理学上的争议等问题。针对上述局限性，目前的研究已使用机械泵模拟心室，各种分布参数(硅胶弹性管等)或集中参数元件(血管顺应性、血流惯性、外周阻力等)模拟动脉后负荷输入阻抗，建立包括人工心脏在内的体外模拟循环系统(Mock Circulatory System,MCS)模型。然而这些研究还尚未对动 脉内皮细胞附近局部微环境中的血压、壁面剪应力和牵张应变进行详细分析，同时普遍存在尺寸偏大、循环液体量多、不包含体外细胞培养腔等缺陷，不便于开展细胞力学生物学研究。
相比较而言，微流控芯片具有所需样品量少、易集成、易于光学检测及良好的生物适应性等优势。近年来的研究表明，基于微流控芯片的体外内皮细胞培养腔(Endothelial Cell CultureModel,ECCM)是能够模拟且易于监控血流动力学微环境信号的微型化、客观化、标准化和定量化的内皮细胞力学生物学研究系统。但是，当前所建立的用于研究人工心脏泵速不同脉动工作模式对应的血流动力学微环境对动脉内皮细胞功能影响的ECCM，未能真实再现人工心脏作用下人体动脉内皮微环境中真实的血流动力学信号。因此，迫切需要设计和构建能够精准模拟在体动脉内皮血流动力学微环境的微型体外模拟循环系统，既能实现人工心脏泵速在不同脉动工作模式下血流动力学信号的精准加载和控制，又能对微流控芯片细胞培养腔内的动脉内皮细胞力学生物学效应进行在线、实时的监测，以便于更好地分析不同脉动工作模式下动脉内皮细胞附近局部微环境中的血压、壁面剪应力、牵张应力(或应变)等关键血流动力学信号，以此为人工心脏泵速脉动工作模式的优化选择提供科学依据，从而提高人工心脏对心衰疾病的治疗与康复能力。
发明内容
本发明的目的在于：提供一种能够真实模拟人工心脏脉动工作模式引起的动脉内皮血流动力学微环境中血压、壁面剪应力和牵张应变(应力)信号的方法。该方法将血流动力学原理、微流控芯片技术及智能反馈控制技术巧妙结合，通过集成度更高、耗材更少的微流控芯片细胞培养腔及表征其后负荷血流动力学特性的多元件集中参数模型搭建体外液体模拟循环系统，再现心衰患者在植 入人工心脏后泵速的不同脉动工作模式下在体动脉内皮细胞所承受的压力、剪应力以及牵张应变的组合作用，可用于研究血流动力学信号与动脉内皮细胞力学生物学效应之间的定量关系及其分子生物学机制。
本发明的技术方案如下：
一种用于优化连续流人工心脏脉动工作模式的内皮细胞体外培养系统(如图1所示)，包括三个基本单元：第一个基本单元为微流控芯片上的细胞培养腔及芯片外多元件主动脉弓后负荷流体力学循环回路(如图2所示)；芯片外多元件主动脉弓后负荷流体力学循环回路包括与细胞培养腔串联的流感、阻力阀、弹性腔1、弹性腔2；弹性腔1、弹性腔2分别设置于细胞培养腔的两侧。
第二个基本单元为模拟心血管系统动力源的脉动式流体加载装置及人工心脏装置，如图2所示，流体加载装置利用脉冲式血液泵实现(图3中Q
1(t))，能够模拟正常和心衰病人在体动脉内皮细胞承受的血压、壁面剪应力和牵张应变波形；人工心脏装置并联接入到脉冲式血液泵的两端(图3中Q
2(t))，再将二者串联接入到上述流体力学循环回路之中，可模拟不同泵速脉动调制模式下在体动脉内皮细胞承受的血压、壁面剪应力和牵张应变信号；
第三个基本单元为外围检测与反馈控制系统，如图1所示，包括倒置的荧光显微镜、CCD高速摄像系统、压力传感器和流量传感器、以及比例-积分-微分反馈控制系统，压力传感器和流量传感器均设置在细胞培养腔的两侧，用于实时监测和采集细胞培养腔输入端和输出端的压力和流量波形，荧光显微镜位于细胞培养腔Rc的上方，CCD高速摄像系统与荧光显微镜相连，用于采集微流控芯片细胞培养腔内细胞的实际形态结构，CCD高速摄像系统、压力传感器、流量传感器均与比例-积分-微分反馈控制系统相连，通过对细胞培养腔两端的压力和流量波形以及细胞形态结构数据的采集及反馈，比例-积分-微分反馈控制系 统可定量调控相关血流动力学信号的变化，在微流控芯片细胞培养腔产生基于人工心脏泵速不同脉动工作模式下在体动脉内皮细胞承受的压力、剪应力及牵张应变信号的组合作用。
如图4所示，上述内皮细胞培养腔是断面为凹字形的腔体，与动脉弹性模量类似的弹性薄膜键合在硬质透光的聚甲基丙烯酸甲酯(polymethyl methacrylate，PMMA)材料制成的腔体上，弹性薄膜下表面以下的细胞培养腔充满循环液体；弹性薄膜的上表面两侧空腔中通入空气，为两侧腔处薄膜在脉动流体压力作用下发生形变提供足够的空间；弹性薄膜上表面中部腔体的两端为光滑圆弧形，且薄膜上表面中部紧贴腔体凹部水平方向的内表面，能够使下方附着内皮细胞的弹性薄膜在两侧牵拉作用下仅发生水平方向的牵张应变，同时腔体凹部厚度的设计须确保在使用显微镜观察内皮细胞形态结构时可聚焦，且在脉动压力作用下不会发生形变；下层细胞培养腔的几何尺寸、弹性薄膜的弹性模量的选择需根据弹性力学原理，并通过精准模拟在主动脉不同部位内皮微环境中的血压、剪应力和牵张应变波形的实际需要确定。
进一步的，所述的细胞培养腔是通过单向阀和贮液池相配合来为细胞培养腔中的细胞提供循环液体的。
进一步的，所述的体外循环系统可等效为电路模型，其中：内皮细胞培养腔的流阻等效为电阻(图3中Rc)、培养腔上薄膜的顺应性等效为电容(图3中C
1)、主动脉弓下游血管床的顺应性、流阻和流感等效为电容、电阻和电感(图3中C
2、R和L)。
进一步的，芯片外多元件主动脉弓后负荷流体力学循环回路的设计需使细胞培养腔薄膜上培养的内皮细胞承受的压力、壁面剪应力、牵张应变与植入人工心脏后心衰病人相应部位内皮细胞承受的血压、剪应力、牵张应变波形一致：
首先，以人体或动物实验检测分析得到的在体动脉内皮细胞局部附近的血压p(t)、壁面剪应力τ
ω(t)和牵张应变ε(t)波形为模拟目标，为使细胞培养腔薄膜上培养的内皮细胞承受的血压和剪应力波形与在体动脉内皮微环境中的血压和壁面剪应力相等，血流量q(t)和压力降Δp(t)必须满足：
式中，η为细胞培养液粘度，Hc、Wc和Lc分别为细胞培养腔的高度、宽度和的长度。
其次，根据流体力学回路与电路之间的相似关系，将主动脉弓后负荷血流动力学行为等效为电路系统。其中流体力学回路输入阻抗表示为输入压力波形p(t)和血流量波形q(t)在频域上比值，用角频率ω
n对应的血压和血流量谐波分量的幅值和相位表征：
∠z(ω
n)＝∠P(ω
n)-∠Q(ω
n) (2b)
式中，|P(ω
n)|和|Q(ω
n)|分别为血压和血流量经傅里叶变换后，在角频率ω
n处的幅值；∠P(ω
n)和∠Q(ω
n)分别为血压和血流量经傅里叶变换后，在角频率ω
n处的相角；|z(ω
n)|和∠z(ω
n)分别为主动脉弓下游后负荷输入阻抗在ω
n处的幅值和相角。等效电路模型的输入阻抗表示为图3中电路元件组合而成的复函数，可基于上述流体力学回路中输入阻抗的幅频曲线和相频曲线，通过式3中系统辨识方法来确定等效集中参数电路模型各元器件的参数值，
式中，
和
分别为集中参数电路模型等效输入阻抗在ω
n处等效阻抗的幅值和相角。
最后，根据流感L、阻力阀R、弹性腔C
1、弹性腔C
2数值搭建模拟主动脉弓后负荷血流动力学特性的多元件体外液体模拟循环系统。
如图1和图2所示，系统中的循环液体为体外血管内皮细胞培养液，弹性腔模拟动脉顺应性(流容)，阻力阀模拟粘性阻力(流阻)，流感元件模拟流动惯性。值得指出的是，由于主动脉弓不同部位壁面附近的血压、剪应力和牵张应变波形不同，描述主动脉弓这些不同部位的后负荷系统血流动力学行为的等效回路可能不是唯一的，需要根据实际波形的不同进行调整。由脉冲式血液泵和人工心脏产生期望的血流量波形。一旦知道细胞培养腔的输入血流量波形，则根据前述等效回路唯一确定压力波形。
上述模拟心血管系统动力源的装置是由人工心脏并联在脉动式流体加载装置两端实现的。可使用脉动式流体加载装置结合PID反馈控制装置实现正常和心衰生理条件下对在体动脉内皮细胞血流动力学微环境中的血压、壁面剪应力和牵张应变信号的模拟，将人工心脏装置与流体加载装置并联，再将二者串联接入到上述流体力学循环回路之中，并结合PID反馈控制装置可产生人工心脏泵速在不同脉动工作模式下主动脉弓不同部位血管内皮细胞承受的血流动力学信号波形，将采集的信号反馈到PID控制装置，可进一步调节流体加载装置及人工心脏，从而定量调控作用于多元件模拟循环系统的压力、流量信号幅度和频率的变化，最终在微流控芯片细胞培养腔产生基于人工心脏泵速不同脉动工作模式下的血压、剪应力和牵张应变的组合作用。
应用上述系统研究连续流人工心脏泵速的不同脉动工作模式与局部动脉内皮微环境血流动力学信号之间定量关系的实验步骤如下：
步骤一：将原代培养的内皮细胞进行传代培养，第2-5代用于实验。调整体外模拟循环系统中流感L、阻力阀R、弹性腔C
1、弹性腔C
2数值的大小，往细胞培养腔中通入循环液体，通过调控脉冲式血液泵和人工心脏装置，在连续流人工心脏泵速不同脉动工作模式下，加载对于动脉内皮细胞各种血流动力学信号的组合刺激。
步骤二：继续加载上述工作模式对应的血流动力学信号刺激，之后对细胞进行活性检测，以确保上述系统的有效性。
步骤三：从微流控芯片细胞培养腔内收集细胞样本，并进行基因和蛋白表达水平的检测，以此分析人工心脏泵速的不同脉动工作模式引起的血压、剪应力、以及牵张应变等血流动力学信号对血管活性物质及促炎性细胞因子基因和蛋白表达水平的影响。
本发明的有益效果：本发明基于上述体外模拟循环系统可成功再现人工心脏泵速的不同脉动工作模式对应的血压、壁面剪应力和牵张应变信号，并使用集成度更高、耗材更少的微流控芯片细胞培养腔研究上述工作模式下血流动力学信号的组合刺激对动脉内皮细胞功能的差异化影响；为研究血流动力学信号与动脉内皮细胞功能之间的定量关系提供了高效合理的实验平台，并为筛选出更有利于改善和维持正常内皮功能的连续流人工心脏泵速工作模式提供了科学依据。
图1为体外内皮细胞培养腔及外围监控系统结构示意图。
图2为体外内皮细胞培养腔流体力学循环回路示意图。
图3为主动脉弓后负荷血流动力学行为等效电路模型示意图。
图4为微流控拉伸芯片细胞培养腔示意图。
图5为正常、心衰及异步调制模式下，通过在体实验获得的主动脉弓部位血管内皮细胞承受的血压和剪应力波形，及根据剪应力波形和细胞培养腔尺寸逆向求解获得的培养腔中血流量波形示意图。
图6为利用Matlab/Simulink对上述等效电路模型分别在正常、心衰和异步调制三种模式下，对实际输入阻抗幅值和相角的拟合结果，并以血流量信息作为电流激励，模型输出电压作为仿真目标，与图5中的血压进行对比后的结果示意图；(a-1)：正常生理条件下输入阻抗的幅度-频率曲线；(a-2)：正常生理条件下输入阻抗的相角-频率曲线；(a-3)：正常生理条件下模型输出电压与血压对比图，(b-1)：心衰条件下输入阻抗的幅度-频率曲线；(b-2)：心衰条件下输入阻抗的相角-频率曲线；(b-3)：心衰条件下模型输出电压与血压对比图，(c-1)：人工心脏泵速的异步脉动工作模式下输入阻抗的幅度-频率曲线；(c-2)：人工心脏泵速的异步脉动工作模式下输入阻抗的相角-频率曲线；(c-3)：人工心脏泵速的异步脉动工作模式下模型输出电压与血压对比图。
图中：包括流体加载装置-脉冲式血液泵(i)和人工心脏(i)；倒置的荧光显微镜、CCD高速摄像系统、压力传感器和流量传感器在内的信号采集处理系统(ii)和比例-积分-微分(Proportional+Integral+Derivative，PID)反馈控制系统(iii)；A(ii)和B(ii)为位于微流控芯片细胞培养腔两端的压力及流量传感器；Rc为微流控芯片细胞培养腔流阻；C
1为薄膜顺应性；R为连接管道流阻；C
2为表征顺应性的弹性空气腔；L为液体循环过程中连接管道的流感。
现针对人工心脏泵速不同脉动模式下的动脉内皮血流动力学微环境中血压的模拟，阐述具体实施方案：
(1)设计微流控芯片细胞培养腔高度Hc、宽度Wc和长度Lc分别为0.3mm、 6mm和15mm，通常细胞培养液的粘度η为0.001Pa·s；利用图5中三种生理状态下的目标血压和血流量分别计算三种目标输入阻抗z(ω
i)；
(2)该体外模拟循环系统的血流动力学特性可使用如图3所示的五元件等效电路模型进行表征，根据相关电路理论可知，该电路的输入阻抗
可表示为：
(3)如公式4所示为该五元件集中参数模型的等效输入阻抗
再结合目标输入阻抗z(ω
i)，通过系统辨识方法可得到各元件参数值，正常生理状态下对应的流体力学回路中各元件参数值分别为：Rc＝8.6kPa·s/ml、R＝113.06kPa·s/ml、C
1＝0.0053ml/kPa、C
2＝0.0097ml/kPa、L＝19.2972kPa·s
2/ml；心衰状态及植入人工心脏后泵速的异步脉动工作模式下对应的流体力学回路中各元件参数值分别为：Rc＝13kPa·s/ml、R＝109kPa·s/ml、C
1＝0.005ml/kPa、C
2＝0.009ml/kPa、L＝1kPa·s
2/ml。如图6所示，五元件集中参数模型对应的输入阻抗曲线(图6中的实线)与目标输入阻抗曲线(图6中的圆圈)基本吻合。基于上述三种不同生理状态下的元件参数值，给定对应的输入血流量波形后，经Matlab/simulink仿真得到的血压波形与图5中相对应的血压波形基本一致，如图6所示，其均方根误差依次为0.237、0.401和0.625；之后构建如图2所示的基于微流控芯片细胞培养腔的流体力学循环回路；
(4)该芯片采用标准化微加工方法进行制作，与动脉弹性模量类似的弹性薄膜键合在硬质透光的PMMA材料制成的腔体上，该腔体断面为凹字形，弹性薄膜下表面以下的细胞培养腔充满循环液体；弹性薄膜的上表面两侧空腔中通入空气，为两侧腔处薄膜在脉动流体压力作用下发生形变提供足够的空间；弹性薄膜上表面中部腔体的两端为光滑圆弧形，且薄膜上表面中部紧贴腔体凹部 水平方向的内表面，能够使下方附着内皮细胞的弹性薄膜在两侧牵拉作用下仅发生水平方向的牵张应变，同时腔体凹部厚度的设计须确保在使用显微镜观察内皮细胞形态结构时可聚焦，且在脉动压力作用下不会发生形变；下层细胞培养腔的几何尺寸、弹性薄膜的弹性模量的选择需根据弹性力学原理，并通过精准模拟在主动脉不同部位内皮微环境中的血压、剪应力和牵张应变波形的实际需要确定；
(5)建立如图1所示的体外内皮细胞培养腔及外围监控系统，包括脉冲式血液泵(i)、人工心脏(i)、各种器件组成的信号采集处理系统(ii)和比例-积分-微分反馈控制系统(iii)。信号采集处理系统(ii)包括倒置的荧光显微镜、CCD高速摄像系统、压力传感器和流量传感器，用于实时监测和采集细胞培养腔输入端A和输出端B的压力及流量波形，以及微流控芯片细胞培养腔内细胞的实际形态结构。脉冲式血液泵(i)结合PID反馈控制装置(iii)可精确模拟正常和心衰生理条件，人工心脏装置(i)并联在脉冲式血液泵两端，可结合PID反馈控制装置(iii)精确模拟连续流人工心脏泵速在不同脉动工作模式下主动脉弓特定部位血管内皮细胞承受的血压、壁面剪应力和牵张应变信号，最终对多元件模拟循环系统加载定量可控的脉动流信号；将采集的信号反馈给PID控制装置(iii)，可进一步调节脉冲式血液泵(i)及人工心脏(i)，从而定量调控作用于多元件模拟循环系统的压力、流量信号幅度和频率的变化，最终在微流控芯片细胞培养腔内产生基于人工心脏泵速不同脉动工作模式下的血压、剪应力、以及牵张应变的组合作用。
微流控芯片内的压力可通过压力传感器测得；剪应力可通过流量传感器测得的流量波形和细胞培养腔几何尺寸计算得到；而对于芯片内弹性薄膜在不同压力下所产生的水平方向的牵张应变则可先利用荧光微球在薄膜上进行标定， 并通过荧光显微镜测量。通过给定不同的压力来获得弹性薄膜相对应的牵张应变，进而建立压力与牵张应变之间的关系表达式。依据该近似表达式，在压力已知的情况下来确定实际实验中弹性薄膜的牵张应变。此外，内皮细胞的形态结构由显微镜结合CCD高速摄像系统检测记录并保存至工控机。
(6)用于研究连续流人工心脏泵速的不同脉动工作模式与局部动脉内皮微环境血流动力学信号之间定量关系的具体实验步骤如下：
步骤一，将原代培养的内皮细胞用EGM培养基进行传代培养，第2-5代用于实验。实验时，将内皮细胞种植于经Fibronection包被的微流控芯片细胞培养腔的弹性膜上，使细胞贴壁且融合度达到90％以上。
步骤二，对动脉内皮细胞加载人工心脏泵速不同脉动工作模式对应的血流动力学信号的组合刺激；使用NucViewTM-488细胞活性检测试剂进行细胞活性检测，以确保体外模拟循环系统的有效性。
步骤三，从微流控芯片细胞培养腔内收集细胞样本进行基因和蛋白表达水平的检测，以此得出人工心脏泵速的不同脉动工作模式下对应的血压、剪应力和牵张应变等血流动力学信号对血管活性物质及促炎性细胞因子基因和蛋白表达水平的影响。
本发明可成功地再现人工心脏泵速的不同脉动工作模式下在体动脉内皮细胞承受的血压、壁面剪应力和牵张应变信号，并可对上述血流动力学信号组合刺激下培养的动脉内皮细胞功能的差异化影响进行实时监测。
Claims (6)
- 一种用于优化连续流人工心脏脉动工作模式的内皮细胞体外培养系统，其特征在于，所述内皮细胞体外培养系统包括三个基本单元：第一个基本单元为微流控芯片上的细胞培养腔及芯片外多元件主动脉弓后负荷流体力学循环回路；芯片外多元件主动脉弓后负荷流体力学循环回路包括与细胞培养腔串联的流感、阻力阀、第一弹性腔和第二弹性腔；第一弹性腔、第二弹性腔分别设置于细胞培养腔的两侧；第二个基本单元为模拟心血管系统动力源的脉动式流体加载装置及人工心脏装置，流体加载装置利用脉冲式血液泵实现；人工心脏装置并联接入到脉冲式血液泵的两端，再将二者串联接入到芯片外多元件主动脉弓后负荷流体力学循环回路之中；第三个基本单元为外围检测与反馈控制系统，包括倒置的荧光显微镜、CCD高速摄像系统、压力传感器、流量传感器和比例-积分-微分反馈控制系统，压力传感器和流量传感器均设置在细胞培养腔的两侧，荧光显微镜位于细胞培养腔Rc的上方，CCD高速摄像系统与荧光显微镜相连，CCD高速摄像系统、压力传感器、流量传感器均与比例-积分-微分反馈控制系统相连。
- 根据权利要求1所述的一种用于优化连续流人工心脏脉动工作模式的内皮细胞体外培养系统，其特征在于，细胞培养腔是断面为凹字形的腔体，与动脉弹性模量类似的弹性薄膜键合在腔体上，弹性薄膜下表面以下的细胞培养腔充满循环液体；弹性薄膜的上表面两侧空腔中通入空气；弹性薄膜的上表面中部紧贴腔体凹部水平方向的内表面；弹性薄膜的上表面中部的两端为光滑圆弧形。
- 根据权利要求1所述的一种用于优化连续流人工心脏脉动工作模式的内皮细胞体外培养系统，其特征在于，所述的体外循环系统的等效电路模型：内 皮细胞培养腔的流阻等效为电阻、培养腔上薄膜的顺应性等效为电容、主动脉弓下游血管床的顺应性、流阻和流感等效为电容、电阻和电感。
- 根据权利要求1所述的一种用于优化连续流人工心脏脉动工作模式的内皮细胞体外培养系统，其特征在于，芯片外多元件主动脉弓后负荷流体力学循环回路的设计需使细胞培养腔薄膜上培养的内皮细胞承受的压力、壁面剪应力、牵张应变与植入人工心脏后心衰病人相应部位内皮细胞承受的血压、剪应力、牵张应变波形一致：首先，以人体或动物实验检测分析得到的在体动脉内皮细胞局部附近的血压p(t)、壁面剪应力τ ω(t)和牵张应变ε(t)波形为模拟目标，为使细胞培养腔薄膜上培养的内皮细胞承受的血压和剪应力波形与在体动脉内皮微环境中的血压和壁面剪应力相等，血流量q(t)和压力降Δp(t)必须满足：
式中，η为细胞培养液粘度，Hc、Wc和Lc分别为细胞培养腔的高度、宽度和的长度；其次，根据流体力学回路与电路之间的相似关系，将主动脉弓后负荷血流动力学行为等效为电路模型，该电路模型将表征主动脉弓下游血管床血流动力学特性的流感L与流阻R串联，并与第二弹性腔C 2并联，并联后与细胞培养腔的流动阻力Rc串联，最后与培养腔上薄膜的顺应性C 1并联，通过系统辨识方法确定集中参数电路模型中上述元器件的参数值；最后，根据流感L、阻力阀R、第一弹性腔C 1、第二弹性腔C 2数值搭建模拟主动脉弓后负荷血流动力学特性的多元件体外液体模拟循环系统。 - 根据权利要求1所述的一种用于优化连续流人工心脏脉动工作模式的内皮细胞体外培养系统，其特征在于，所述的细胞培养腔是通过单向阀和贮液池相配合来为细胞培养腔中的细胞提供循环液体的。
- 根据权利要求2所述的一种用于优化连续流人工心脏脉动工作模式的内皮细胞体外培养系统，其特征在于，使用脉动式流体加载装置结合PID反馈控制装置可实现正常和心衰生理条件下对在体动脉内皮细胞血流动力学微环境中的血压、壁面剪应力和牵张应变信号的模拟，将人工心脏装置与流体加载装置并联，再将二者串联接入到上述流体力学循环回路之中，并结合PID反馈控制装置可产生人工心脏泵速在不同脉动工作模式下主动脉弓不同部位的血流动力学信号波形；将采集的信号反馈到PID控制装置，可进一步调节流体加载装置及人工心脏，从而定量调控作用于多元件模拟循环系统的压力、流量信号幅度和频率的变化，最终在微流控芯片细胞培养腔产生基于人工心脏泵速不同脉动工作模式下的血压、剪应力和牵张应变的组合作用。
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