Studying the response of aortic endothelial cells under pulsatile flow

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Studying the response of aortic endothelial cells under pulsatile flow using a compact microfluidic system Mokhaled Mohammed, Peter Thurgood, Christopher Gilliam, Ngan Nguyen, Elena Pirogova, Karlheinz Peter, Khashayar Khoshmanesh, and Sara Baratchi Anal. Chem., Just Accepted Manuscript • DOI: 10.1021/acs.analchem.9b03247 • Publication Date (Web): 13 Aug 2019 Downloaded from pubs.acs.org on August 13, 2019

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Analytical Chemistry

Studying the response of aortic endothelial cells under pulsatile flow using a compact microfluidic system

Mokhaled Mohammed 1, Peter Thurgood 1, Christopher Gilliam 1, Ngan Nguyen 1, Elena Pirogova 1, Karlheinz Peter 2, Khashayar Khoshmanesh 1, Sara Baratchi 3, †

1

School of Engineering, RMIT University, Melbourne, VIC 3001, Australia 2

3

Baker Heart and Diabetes Institute, Melbourne VIC 3004, Australia

School of Health and Biomedical Sciences, RMIT University, Bundoora, VIC 3083, Australia

† Corresponding author: [email protected]

Abstract We describe a piezoelectric pumping system for studying the mechanobiology of human aortic endothelial cells (HAECs) under pulsatile flow in microfluidic structures. The system takes advantage of commercially available components, including pumps, flow sensors and microfluidic channels, which can be easily integrated, programmed and operated by cellular biologists. Proof-of-concept experiments were performed to elucidate the complex mechanotransduction processes of endothelial cells to pulsatile flow. In particular, we investigated the effect of atheroprone and atheroprotective pulsatile shear stress on endothelial cytoskeleton remodeling, distribution of β-catenin as well as nuclear shape and size. The system is simple to operate, relatively inexpensive, portable and controllable, providing opportunities for studying the mechanobiology of endothelial cells using microfluidic technologies.

Keywords: mechanobiology, microfluidics, self-sufficient, pulsatile flow, shear stress, endothelial cells

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1. Introduction Physical stress associated with blood flow plays an important role in maintaining vascular function and homeostasis 1,2. Endothelial cells lining the inner surface of blood vessels are in direct contact with blood flow and convert hemodynamic forces into biochemical signalling via certain classes of molecules called mechanoreceptors 3. In particular, shear stress plays a major role in the mechanobiology of aortic endothelial cells

4-8.

Arterial shear stress levels

(>15 dyne/cm2) regulate endothelial cell responses, which suppress inflammation, coagulation as wells as transmigration and adhesion of leukocytes, and thus induce ‘atheroprotective’ effects. In contrast, low shear stress levels (< 4 dyne/cm2) do the opposite and induce ‘atheroprone’ effects 9,10. Shear stress induced at the surface of endothelial cells is proportional to the flow rate of the blood passing through the vessel 3. The blood flow is highly pulsatile with zero or even 11.

reversing flow during systole within the coronary arteries

This suggests that the shear

stress is pulsatile and therefore varies over time. The time varying characteristic of shear stress influences several aspects of endothelial responses

12.

Changes in the pulsation

frequency are reported to affect endothelial permeability and adaptive responses 13 as well as the rate of cell cycle arrest

14.

Pulsatile flow also impacts the endothelial orientation,

morphology and importantly cytoskeleton and cell-cell junction remodeling 15 and expression of cell adhesion proteins such as β-catenin 16. Several innovative methods have been explored to generate pulsatile flow in microfluidic systems. This includes the combination of steady and oscillatory flows driven by two syringe pumps through a Y-shaped channel

17,18.

The oscillatory flow was provided by passing the

liquid through a frequency-controlled solenoid valve to induce pressure pulses inside the elastic inlet tube. A similar strategy was implemented by incorporating a pneumatically actuated collapsible chamber in a recirculating setup driven by a peristaltic pump to induce pressure pulses 19. Advances in microfabrication techniques have facilitated the development of various micropumps, which utilise various mechanical electrowetting

26

20-23,

thermal

24,

electrical

25

and

mechanisms for driving of liquid. Such pumps can be integrated into the

microfluidic system, reducing the size and cost of the overall system, enabling self-sufficient, portable and low-cost setups

27.

Importantly, the small size of miniaturised pumps reduces

inertial effects and facilitates pulsatile flow patterns. This includes pneumatically-actuated 2 ACS Paragon Plus Environment

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micropumps, utilising two layers of microfluidic channels that are separated by a thin membrane 28. Applying pressurised gas to one channel deforms the other channel that is filled with liquid. Customised pulsatile flows can be obtained by modulating the pressure and frequency of the pressurised gas

20,29.

More biomimetic versions of pneumatically-actuated

micropumps involving several pumping unites and check valves have been recently developed

30.

The harmonic movement of piezoelectric pins has also been employed to

deform the elastomeric microfluidic channels, leading to the generation of customised pulsatile flows

31.

Piezoelectrically-actuated micropumps, utilising piezoelectric diaphragms

enable highly controllable pulsatile flows 22,32. We envisage that the simplification of microfluidic technologies (in terms of fabrication, programming and operation) along with minimising their reliance on supporting equipment 27,33-37

can accelerate their adoption in mechanobiology labs.

To achieve this, we describe a piezoelectric pumping system for applying customized dynamic flow profiles through microfluidic systems. We take advantage of commercially available miniaturised pumping units, flow sensors and microfluidic channels. The system can be easily integrated, programmed and operated by cell biologists without knowledge and expertise in microfabrication or extensive training in programming and interfacing of microfluidic components. The system is utilised for analysing the mechanobiology of human aortic endothelial cells (HAECs) under pulsatile flow. In particular, we investigate the formation and orientation of stress fibres, assembly and distribution of β-catenin cell adhesion molecule as well as nuclear shape and circularity under pulsatile flow conditions. The system is portable, simple to operate, relatively low-cost and highly controllable. These features make it highly suitable for in vitro investigation of the effect of pulsatile flow on endothelial function and homeostasis.



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2. Materials and methods 2.1 Piezoelectric pumping set up and characterisation The experimental set up consisted of three piezoelectric pumps mounted on a printed circuit board, a microprocessor, a microfluidic system with three independent channels, interconnecting tubes and reservoirs for recirculation of the cell culture medium through the channels (Figure 1a). The entire experimental setup was placed inside an incubator (37°C, 50% relative humidity) during experiments. Our experiments indicated that operating at physiological temperatures does not lead to generation of bubbles inside the microfluidic channel. Commercially available piezoelectric pumps (Bartels Mikrotechnik GmbH, mp6-pi) were used for driving the cell culture medium through the system. These pumps utilise two piezoelectric membranes arranged in series to increase the operating pressure of the pump and can drive water-based solutions at a maximum flow rate of 7 mL/min. Given that the flow rate of piezoelectric pumps is inversely proportional to back pressure (viscous pressure drop induced along the microfluidic channel, interfaces and connecting tubes), the pumps can be coupled in series or parallel configurations to produce desired flow rates according to the design. The piezoelectric pumps are mechanically robust and have a lifetime of >5000 hrs according to the manufacturer’s specifications. We have used the pumps for >150 hrs during our experiments and have not noticed any significant difference in their performance. The flow rate magnitude and pattern were controlled through the DC voltage supplied to the pumps using an Arduino Mega 2560 microprocessor (Arduino Corporation). The microprocessor could be easily programmed with the Arduino software (IDE). The voltage supplied to the pumps was adjusted using the pulse width modulation (PWM) outputs embedded in the microprocessor. The microprocessor had 14 PWM output and therefore could control up to 14 piezoelectric pumps simultaneously. Commercially

available

microfluidic

channels

(ibidi

GmbH,

µslide

VI

0.1,

W×H×L=1×0.1×17 mm) were used for culturing of cells under different flow conditions. These channels are made of bio-compatible polymers, and are tissue culture treated and sterilized, which facilitates studying adherent cells under flow conditions. Silicon tubes (ibidi GmbH, internal diameter = 0.5 mm) with a total length of ~30 cm along with 15 mL conical tubes (Corning® 15mL Centrifuge Tubes) were used for recirculation of the cell culture medium. Mesh filters such as cell strainers can be incorporated into the conical tubes to avoid debris contamination in recirculating flow. 4 ACS Paragon Plus Environment

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Cellular experiments were performed at two shear stress levels, representing both atheroprone (4 dyne/cm2) and atheroprotective (20 dyne/cm2) conditions. Pulsatile flow patterns were generated at the frequencies of 1 and 2 Hz and were monitored in real-time using a digital flow sensor (Sensirion, SLI-1000) (Figure 1b-c). Under laminar flow regime governing the microfluidic channels, the shear stress induced at the cultured endothelial cells was approximated as 𝜏𝑐𝑒𝑙𝑙 = 6 𝜇 𝑄 𝑊𝐻2, in which 𝜏𝑐𝑒𝑙𝑙 is the shear stress, 𝜇 and 𝑄 are the dynamic viscosity and flow rate of the cell culture medium, respectively, while 𝑊 and 𝐻 are the width and height of the microfluidic channel, respectively. Our numerical simulations indicated that this equation is only valid in the middle regions of the channel, which are away from the sidewalls (Figure S1). As such, when studying the orientation of actin filaments in response to various shear flows, we limited our observation to the cultured cells that were 100 𝜇m away from the sidewalls to ensure a uniform shear stress distribution. Given the dimensions of the channels, atheroprone and atheroprotective shear conditions were mimicked by setting the maximum flow rate of the pumps to 56 and 276 µL/min, respectively (Figure 1b-c).



Figure 1. Principles of system operation: (a) Schematics of the experimental setup comprising of, three independently operating piezoelectric pumps for applying constant and pulsatile flows, a microcontroller, a microfluidic system incorporating independent channels, flow sensors, liquid chambers and tubes for recirculation of flow. (b-c) Pulsatile flow patterns for inducing atheroprotective shear at 1 and 2 Hz generated by the piezoelectric pumps.


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2.2 Cell culture and in-vitro generation of pulsatile flow The primary HAECs were purchased from Lonza (Walkersville, USA) and cultured in EGM2 media supplemented with SingleQuots kit according to the supplier’s instructions. Cells at early passages of 2-6 were used in this study. For shear stress experiments, HAECs were cultured inside the ibidi channels according to the supplier’s instructions for 24 hrs so that the endothelial cells could adhere to the bottom surface of the channels. The cultured HAECs were then exposed to stress levels of 4 and 20 dyne/cm2 for 24 hrs.

2.3 Immunocytochemistry and confocal microscopy Immunostaining and image acquisition were performed, as reported before

3,38.

Briefly, cells

were fixed inside the microfluidic channels using 4% paraformaldehyde and permeabilised with 0.2% Triton X-100 in phosphate-buffered saline (PBS). To prevent the non-specific antibody binding, the sample was blocked with 2% goat serum for 1 hr at 37°C. β-catenin was stained with anti-β-catenin antibody (Cell signalling, 1/200) followed by incubation with Alexa488 conjugated anti-rabbit secondary antibody (Cell signalling, 1/400). F-actin was stained with Atto 565-phalloidin (Sigma-Aldrich, 1/500) and nucleus were stained with DAPI (Sigma-Aldrich). Nikon A1 confocal scanning microscope was used for image acquisition.

2.4 Image processing and analysis of actin filaments Image processing and analysis of actin filaments were performed, as reported elsewhere

39.

Briefly, an automated image processing algorithm written in MATLAB determined the orientation of the actin filaments based on the spatial gradients of the image. Specifically, a pixel-by-pixel vector field was constructed from the intensity gradient in the horizontal and vertical directions of the image. The magnitude and directional information of this vector field were then used to compute a histogram of orientations for each pixel in the image. The histograms evaluate the deviation of the vector field in small sub-regions (20×20 pixels in size) from a set of angles ranging from

89 to 90 degrees relative to the horizontal. Finally,

the local dominant orientations of the actin filaments were determined by choosing the angles that corresponded to the largest values in the histograms, as further detailed in 40. Statistical significance was assessed with a global Watson’s U2 test, and statistics were computed using the circular statistic toolbox 41. 7 ACS Paragon Plus Environment


2.5 Quantification of nuclear shape and size Quantification of nuclear shape and size was performed using NIS element software (Nikon). For statistical analysis, one-way ANOVA was performed using Prism 7.02 (GraphPad software) and P

Studying the response of aortic endothelial cells under pulsatile flow

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