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This review describes the use of ultrasound for inducing and retaining cell-cell contact in multi-well microplates combined with live-cell fluorescence microscopy. The two different designs of ultrasonic actuation system used for the multi-well microplate: The wedge-transducer device [18] (a) and the ring-transducer device [17] (b).
The purpose of the ultrasonic actuation in the multi-well microplate is to create an acoustic resonance in each well, so that suspended particles or cells are aggregated and positioned by the acoustic radiation force described by Equation (1).
Detailed experimental evaluation of the trapping and positioning performance of 10 µm particles in the ring-transducer device. For the practical handling of the device with a cell suspension sample, the method is summarized in Figure 5. One of the standard procedure to pump solutions, emulsions or suspensions, into microfluidic chips is based on the use of syringes, through peristaltic pumps.
Luer taper is a standardized system of small-scale fluid fittings used for making leak-free connections between a male-taper fitting and its mating female part on medical and laboratory instruments, including hypodermic syringe tips and needles or stopcocks and needles.
The lack of an efficient interface, or interconnect, between microfluidic devices and the macroscale world is a major impediment to the broader application of micro total analysis system (µTAS).
This tip presents an alternative approach to making simple, cost-effective universal interconnects. Microfluidic and lab-on-a-chip (LOC) devices are typically connected to external sample reservoirs. The reservoirs are cheap, disposable, easy to assemble, and robust enough to support pressures up to 345 kPa (and even higher pressures with the variant described in the supplementary material) while having almost no dead volume. Figure 6.  A picture of a chip connected to three reservoirs placed on a storage rack, each with an independent input line and multiple outputs. It is important to maintain the needle connected to the pressure source above the liquid surface in order to prevent the formation of bubbles. The presented setup above was not tested at pressures higher than 345 kPa, and at pressures above 140 kPa, there was minimal gas leakage.
We mention here that this setup is permanent and should be robust at high pressure applications. In suspension, cells, beads and particles tend to sediment at a speed given by the balance of viscous drag and gravity force. Another common way to prevent sedimentation is to match the densities of the particles and the carrier fluid, or to increase the viscosity of the carrier fluid to increase sedimentation times. Previously in the Chips and Tips section a system has been described for cell injection [1].
Here we combine both systems [1,2] to inject cells, beads and particles efficiently into microfluidic devices using small amounts of liquids (typically ~1 mL). 3) Fill the syringe that does not contain the magnet with water and hold it vertical (Fig 3). 4) Insert the needles in the tubing, fill a third syringe with water, place the needle in position, flush the tubing with water and connect the syringe with the free needle (Fig 4).
5) Invert the syringe to have the magnet in the upper part and fill this part with the liquid containing the cells, beads or particles (Fig 5) and connect a short needle with tubing (Fig 6). 7) The system is now ready to be connected to a chip on one hand and to a syringe pump on the other (see video below). Syringe pumps are often used in a variety of microfluidic applications because of their portability and the ease with which flow rates can be changed. The method described is a quick assessment of the pulsations generated by syringe pump driven flow. This platform has been used for studying the interaction between natural killer (NK) cells and cancer cells at the level of individual cells.
IntroductionDynamic studies of single cells are important for our understanding of cell function and behavior [1]. If the purpose is to trap the particles in the center of each well, one should use the lowest possible resonance mode. The fluid reservoir within the PDMS frame shown in Figure 5a has the purpose of providing a controlled environment of temperature- and CO2-regulated cell medium.
For simplicity, only three out of a hundred wells are shown (vertical cross-section view), and only the PDMS frame and the multi-well microplate (item #4 and #5 in Figure 2). Quantifying Acoustic Energy Density, Acoustic Pressure Amplitude, Acoustic Radiation Forces and Acoustic StreamingIn order to fully characterize the device, it is important to be able to measure the properties of the acoustic field including the acoustic radiation forces and acoustic streaming acting on the particles, cells and the fluid, respectively.
Syringe pumps are usually preferred over peristaltic ones, for their ease of use, for the accurate and stable control of the flow rate and, finally, for the possibility to employ sterile conditions. The fitting is named after the 19th century German medical instrument maker Hermann Wulfing Luer. Luer components are manufactured either from metal or plastic and are available from many companies worldwide but they are usually sold in few standard dimensions and are relatively expensive. Examples of the use of female Luer X female Luer adapter to connect syringes to a commercial chip. Examples of the use of female Luer X male Luer elbow to connect syringes to a homemade chip. There are a number of current products and solutions available that seek to address this issue. The device was developed to allow for standard compression tubing connectors, such as those available from Upchurch, to be interfaced with microfluidic chips. The first component was a port which allowed a standard ferruled connector to be attached perpendicular to the microfluidic device.  It contained a threaded portion for the fitting and a fluid path with small port on the bottom for the interface (Figure 2). The second component was a breadboard clamp which applied pressure between the microfluidic device and interconnects (Figure 3). Setup used for pressure testing (left) and the close up view of a chip in the device (right).
Das, Interfacing of microfluidic devices, Chips & Tips (Lab on a Chip), 27 February 2009.
A significant issue for LOC devices is interfacing multiple sample reservoirs and connecting them both to external pressure controllers or syringe pumps and to the microfluidic chip [1,2]. The reservoirs are made using off-the-shelf threaded microcentrifuge tubes and screw-on septum caps. In order to insert a glass capillary to serve as an output, use a sharp hypodermic needle with an inner diameter (ID) that is just larger than the capillary’s outer diameter (OD). Insert a needle into the membrane screw cap and pass the glass capillary through the needle. Screw the cap onto a reservoir of the desired size and then remove the needle while holding the capillary in place, as shown in Fig.3.
Insert the input pressure needle to appropriate depth (see next section for details) and connect to the Tygon tubing to the needle using the Luer-lock to tubing adapters, as shown in Fig.4.
Reservoir with a needle connection interfaced to a glass capillary (typically connected to the microfluidic chip) and Tygon tubing with a Luer-lock adaptor (typically connected to the pressure source) which can be used to control the flow from the reservoir to the chip.
However, we have developed another setup that works without any leakage, even at 345 kPa, but requires several minutes to make and a precision micro-drill. At the same time, it requires a precision micro drill, more time to make, and it is less flexible in the sense that it is harder to modify the interface. High pressure off-chip containers with one Luer-Lock input and one glass capillary output at the bottom. In classical biological experiments, the most straight-forward solution to keep cells in suspension is to shake large volumes.
However, these solutions are not always suitable due to the biological or physical constraints of a microfluidic experiment.
With a similar system, we observed that sedimentation along the tubing binding the syringe reservoir to the chip hinders good mixing of the cells. The cells are initially placed in a syringe close to the chip and are automatically shaken inside the syringe using a small motor and a magnet. Make sure no air bubbles are present in the water phase: they will increase the response time of the system to flow rate modifications. More sophisticated systems can probably be adapted to allow a temperature control of the syringe.
After bonding is done, flush the microfluidic system with deionized water for several minutes to dissolve and remove salt particles. The syringe pump characteristics (type, age and wear), compliance in the tubing, a mismatch between the size of syringe used and the flow rate desired can generate pulsations in microfluidlic flows. The device consists of two identical channels connected downstream to form a comparator region.

Because the hydrodynamic resistances of the two channels are equal, if any pulsations are present in the syringe-pump driven flow then the dyed fluid will be displaced above or below the symmetry line (shown in white) in the comparator, as shown in Figure 3b and 3c. For a more in depth study, we recommend using a precision hydrostatic head rather than a ring stand.
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The review includes basic principles of ultrasonic particle manipulation, design criteria when building a multi-well microplate device for this purpose, biocompatibility aspects, and finally, two examples of biological applications: Dynamic imaging of the inhibitory immune synapse, and studies of the heterogeneity in killing dynamics of NK cells interacting with cancer cells. Cells are complex biological systems that respond to different kinds of stimuli over time periods ranging from fractions of a second [2] to several days [3]. From the equation, we conclude that the radiation force drives suspended particles in a direction parallel with the gradient of the acoustic field and has a direction and magnitude defined by the contrast factors f1 and f2. The wells used in our multi-well microplate have square-shaped horizontal cross-sections (cf. The length of the Luer connector can be adjusted depending on the specific needs of the researcher (for instance, the distance between the syringe pump and the chip). An example of a simple approach is the direct integration of tubing or a syringe needle into the microchip inlet using epoxy glue.[1] However, direct connection to the inlet reservoir using an epoxy often leads to clogging of the microchannels. A standard breadboard fixture was used to compress the ports against the chip to apply sealing pressure, preventing leaks at the interface.  This method allowed for connections to be made without glue, epoxy, or any form of bonding, enabling the components to be easily and quickly reused or reconfigured without the need to machine new fixtures or re-bond ports. This breadboard approach was particularly useful for prototyping devices having different dimensions and port locations, as interconnects could easily be relocated. Arrays of off-chip reservoirs can be assembled within racks and many lines can be connected at high density to a microfluidic chip. Insert the needle into the membrane screw cap and then slide the capillary through the needle, as shown in Fig.
Unscrew the cap from the container and fill it with your reagent and screw the cap back on. Make sure to create slightly smaller sized holes for the input connections on the chip to create a tight seal.
Since the system is pressure controlled, a small leakage rate relative to the supply flow rate can be tolerated. This type of system is relevant at high Reynolds number but is not very compatible with the classical systems used to inject small amounts of liquid in confined geometries, such as microfluidic systems. For example, increasing viscosities increases the pressure drop in microchannels, and additives for density matching are not always biocompatible.
This tubing should therefore be as short as possible, which is sometimes incompatible with the large footprint of the syringe pumps or other equipment around the microfluidic chip.
This syringe is then remotely actuated by another syringe mounted on a syringe pump [2] which reduces the footprint of the injection system in the vicinity of the chip. The speed of the motor is simply controlled by the voltage to guarantee gentle agitation of cells. Lee, Preventing suspension settling during injection, Chips & Tips (Lab on a Chip), 21 August 2007. Figure 1 illustrates a sagging problem, which resulted in the top chamber wall being accidentally bonded to the bottom wall.  Here, we describe a tip to prevent sagging by using regular cooking salt.
Try to perform this action using fine tweezers under a stereomicroscope, if the chamber is too small. As figure 4 shows, the bonded chamber does not exhibit adhesion between the top and bottom wall.
These pulsations could be undesirable for lab-on-a-chip applications in which steady fluid flows are needed. Vary the height of the hydrostatic head until the two fluid flows meet at the symmetry line of the comparator (see the white line in Figure 3a).
We built a stand that can hold a syringe on a precision linear translation stage (Edmund Optics, Part # NT56-796).
Even if the stimulus is kept constant in a certain measurement, the cellular response may vary over time and from cell to cell [4]. To this primed volume of cell medium, a small aliquot of cell suspension is pipetted from above (see Figure 5b) before closing the device with the glass lid (cf.
The average number of cells per well is controlled by the cell concentration in the added drop in (b). Note that in the case of H-PTFE tube (B), the cutting is made with an angle of 45° with respect to the tube major axis. Press and insert the H-PTFE tube end (cut at 45°) into the heated FEP tube, cool down the tube connection by tap water until a permanent deformation is reached. Commercially available connections to microfluidic chips and devices, such as the NanoPort from Upchurch Scientific (Oak Harbor, WA, 98277, USA), accommodate these needs by use of a threaded nut and ferrule system. The working pressure of these ports was tested with a syringe pump, and withstood pressures in excess of 1000 psi without any signs of leakage or failure.
A wide variety of bolt locations in the clamping plates allowed for a wide variety of component sizes and configurations. The connectivity between reservoir and pressure source(s) and chip(s) can be adapted to specific needs, for example to form a reservoir with a single input and a single output, or with multiple inputs and outputs. These plug-and-play reservoirs are a versatile solution to the world-to-chip interface and should find widespread application for microfluidic experimentation.
When using multiple reservoirs, they can be placed into microcentrifuge tube racks for easy handling (Microtube Storage Racks), as shown in Fig. Only a limited number of connections, up to 10 in our experience, can be connected to a single cap or else leakage becomes excessive. Using suitable micro-sized drill bits, different sized holes can be made in the micro tube or the cap. Interestingly, in another Chips and Tips section, a system for temperature control of syringe pumps has been described [2].
The tubing length from the syringe containing the cells to the chip can then be as small as 5 cm, which reduces the residence time of cells in the tubing and therefore reduces sedimentation. Nastruzzi, An easy temperature control system for syringe pumps, Chips & Tips (Lab on a Chip), 22 April 2008. The method described here is a quick means to assess the degree of pulsations present in flows driven by syringe pumps.
In standard bulk-based assays, such variations are not resolved since the measured parameter is typically the average signal from many cells [1].
Since steeper field gradients result in stronger forces, standing-wave fields are most often utilized. For a square-shaped cavity (Lx= Ly), the lowest possible mode is either the (1,0)-mode or the (0,1)-mode. Experimental confirmation of the simulations using 5 µm particles at single-frequency actuation (c) and with frequency-modulated actuation (d) using the same frequency intervals as (a) and (b).
These fittings can be integrated onto chip reservoirs by means of adhesive rings or epoxy glue.[2] The drawback to this method is that NanoPorts require the use of an epoxy glue that takes time to cure and is non-removable, or an adhesive ring that is not compatible with many solvents. Kumacheva,  Reusable, robust NanoPort connections to PDMS chips, Chips & Tips (Lab on a Chip), 8 October 2008.
The reservoirs work well up to 345 kPa, and for higher pressure applications solid caps can be used as described in the supplementary material.
After inserting the output capillaries or steel tubing and the input syringe needle, Instant Krazy glue [10] is applied to affix the inputs and outputs at the precise place. The idea there was to use a syringe pump to pump liquid in order to move the piston of another syringe placed in a temperature-controlled bath. The basic principle relies on using a microfluidic comparator [1,2] to detect small pressure fluctuations in fluid flows. Using this hydrostatic head and a video camera, one can precisely determine the number of fluctuations per unit time for a specified pump, flow rate, syringe size and tubing. Thus, there is a need for screening methods where an individual cell’s properties are measured.
In a standing-wave field, the radiation force drives most suspended particles either to the pressure nodes or the pressure anti-nodes, depending on the signs of the contrast factors f1 and f2.
According to Equation (3), these two modes have identical frequencies, and it is therefore not clear how the pressure field can be described at this frequency.
Scale bar is indicated by the wells (300 µm wide squares).The figure is reproduced from Ref.
Harnett, Integrated reservoirs for PDMS microfluidic chips, Chips & Tips (Lab on a Chip), 22 April 2008. This system still showed problems related to the length of the connection tubing to the chip. One such established and very efficient screening method is flow cytometry, or fluorescence-activated cell sorting (FACS) [5]. In principle, particles stiffer than the suspension medium are driven to the pressure nodes (defined by the first term in Equation (1a)), while particles denser than the suspension medium are driven to the velocity antinodes (defined by the second term in Equation (1a)).

The strength of this method, besides simplicity, is the ability to study different numbers of cells per well interacting.
In simple standing-wave fields (such as a one-dimensional field), the pressure nodes and the velocity antinodes are co-located. Thus, the average number of cells per well is controlled by the cell concentration in the added drop (cf.
Thus, knowing these acoustic field properties is important for estimating the trapping efficiency of cells, but also for estimating the risk of having cavitation in the sample (see Section 3). However, flow cytometry in its standard format is not compatible with dynamic monitoring, and therefore, only instantaneous cell properties are measured. Although the standing-wave field in the multi-well plate is of more complex type, experimental observations confirm that the cells used in our work are driven to the positions of the numerically calculated pressure nodes [16].When several particles are driven to a pressure node, they tend to aggregate in tight clusters. In our multi-well microplate, where a complex acoustic interaction between all 100 wells occurs, the result is a pressure field inside each well having a node oriented as shown in the simulated pressure field in Figure 3a (and described more thoroughly in Ref. Figure 5b), and the seeding principle causes a stochastical distribution around this average. It is very difficult to measure the acoustic field properties by direct methods in an acoustofluidic device.
On the contrary, live-cell fluorescence microscopy is a suitable tool for measuring dynamic cell properties [6].
In one-dimensional (1D) standing-wave fields, the clusters typically take the form of flat monolayers in the pressure nodal planes.
Finally, when ultrasound is applied with the frequency-modulation method presented in Figure 3, the cells are aggregated and positioned in the center of each well where they can be monitored over time by high-resolution fluorescence microscopy.
Today, many different fluorescent probes exist for monitoring a variety of cellular functions, processes and status.
The reason for this is the particle-particle interaction force, sometimes called the Bjerknes force [23]. As seen from the simulation, the node of the half-wave resonance in each well is not a pure (1,0)- or (0,1)-mode, but rather something in between. If confocal microscopy is used, it is a benefit to know the exact locations of the 100 cell aggregates. Here, we will present three different methods used in our lab: Light intensity, particle tracking and particle image velocimetry (PIV). Lee, A novel high aspect ratio microfluidic design to provide a stable and uniform microenvironment for cell growth in a high throughput mammalian cell culture array, Lab Chip, 2005, 5, 44-48. However, in 2D or 3D standing-wave fields, the cluster shapes are more complicated to predict or control [24].The theoretical model above (Equation (1)) is valid for spherical particles with well-known material properties (density and compressibility) suspended in an inviscid fluid. If the excitation frequency is slightly changed, the node orientation changes (primarily it rotates). Since confocal microscopy is a relatively slow method, only the small area where the cells are located needs to be scanned, instead of the whole microplate. However, in order to combine both dynamic and single-cell monitoring, a tool for keeping track of each cell over time is needed. For this reason, a simple method for generating point-shaped pressure nodes in the center of each well is to quickly average a set of such single-frequency resonances.
All methods are based on one-dimensional (1D) geometries, but can be used for 2D geometries for order-of-magnitude estimations of the energy, pressure and forces.The first method, light intensity, has been developed in collaboration with Rune Barnkob and Henrik Bruus (Denmark Technical University, Copenhagen, Denmark) [28,32].
We have realized this by cycling linear frequency sweeps around the nominal (1,0)- or (0,1)-resonance frequency. This method is specifically designed for acoustofluidic chips that are optically transparent and compatible with standard bright-field microscopy [6]. In addition, the material properties of cells are also dependent on many external and internal factors. One strength of the light intensity method is that the chips do not need to be compatible with high-resolution microscopy.
However, in this review we focus on the use of microplates for live-cell microscopy imaging of individual cells [9].When studying single cells in multi-well microplates, there are two different strategies that can be employed.
The simulated acoustic resonances in Figure 3a,b are experimentally confirmed using the wedge-transducer device in Figure 3c,d, respectively. One strategy is to dispense one cell per well for keeping track of each individual property of each single cell.
Nevertheless, attempts have been made to estimate the contrast factors and corresponding radiation forces acting on different cell types. Here, the pressure field is visualized by the shapes and positions of 5 µm polyamide particle aggregates driven to the pressure nodes by the acoustic radiation force. In brief, the method is based on measuring the total transmitted light intensity passing through a certain part of a microchannel or a microchamber during the focusing process of suspended particles. Experimentally, frequency modulation actuation is realized by sawtooth-modulation with a center frequency corresponding to the (1,0) or (0,1) resonance (which is around 2.5 MHz for a 300 ? 300 µm2 well) and a typical bandwidth of 100 kHz and a cycling rate of 1 kHz. When particles are focused and trapped in the pressure node, the fluid cavity gradually becomes more transparent for light. An alternative strategy is to dispense several cells per well and then to study individual interactions between two or more cells.
This modulation function is very simple to implement since it is a built-in function in most signal generators. This has been used for studying interaction between natural killer cells and different types and numbers of target cells [9,11,12,13,14], and also for migration studies [15].Although microplates combined with live-cell fluorescence microscopy is a powerful tool for parallel and dynamic single-cell studies, it is still problematic when studying cell-cell interactions. They concluded that the radiation force was about one order of magnitude smaller for these cells, relative polystyrene particles of equal size. The modulation bandwidth is difficult to predict theoretically and needs instead to be optimized experimentally for each microplate device.
The reason is that the time to cell–cell contact may vary depending on cell type, environment and microplate design [11].
In Figure 6, we demonstrate quantification of the acoustic energy density in a microchannel when it is operated at a single (optimal) frequency, and we compare with the corresponding energy density for the same (center) frequency, but with the frequency-modulation function active (100 kHz bandwidth and 1 kHz rate). In addition, there is a stochastic distribution in time to contact between different wells. As seen in the diagram, the average energy density is just slightly lower for frequency-modulation relative single-frequency actuation. Finally, a similar experimental verification for the ring-transducer device is shown in Figure 4. This result is important since it means that there is no compromise between positioning accuracy and radiation forces when using frequency- modulation actuation instead of single-frequency actuation.Another relatively simple and straightforward method for measuring acoustic energy density is particle tracking. Furthermore, it may also be of interest to study the effects of forced interaction and compare with spontaneous interaction.
Here, we used a microplate with well size 350 ? 350 µm2 actuated with center frequency 2.30 MHz and modulation bandwidth of 200 kHz.
This method is based on either manual or automated tracking of the position of individual particles over time.
For these reasons, we have during the last few years developed and implemented ultrasonic particle manipulation technology into a microplate device designed for high-resolution live-cell fluorescence microscopy [16,17,18].
When using a lower concentration of 10 µm particles, it is clear from the experiments that particle trapping and aggregation work for both single-frequency actuation (blue aggregates) and frequency-modulation actuation (red aggregates). In this review, we summarize our work on the use of ultrasound as a tool for inducing and retaining cell–cell contacts in such multi-well microplates.
However, the accurate positioning of aggregates in the center of each well can only be accomplished by frequency-modulation actuation. The review discusses basic principles of ultrasonic manipulation technology and design criteria for such microdevices (Section 2), how to design a biocompatible manipulation system (Section 3), and finally an example of a biological application of the platform where the interaction between natural killer (NK) cells and cancer cells are studied (Section 4).
In summary, we may expect the acoustic radiation force in a given acoustic field to be roughly a few times smaller for cells than for polystyrene particles of similar size, and that the corresponding trapping time is expected to be a few times longer.
In addition, frequency modulation also provides more compact aggregates [17].A limiting factor for the trapping performance in any acoustophoretic device is acoustic streaming [30]. In the multi-well microplate, acoustic streaming causes the trapped particles or cells to be flushed away upwards if very high actuation voltages are used (approx. One reason for this streaming is that it is not an accurate approximation to model the wells in the microplate as 2D cavities (cf. Thus, Equations (2) and (3) are useful for qualitative understanding of how resonances are built up in the system and for predicting trapping positions of cells, but for accurate quantitative modeling including acoustic streaming, a 3D model is needed. Still, it should be mentioned that the frequency-modulation methods tends to suppress acoustic streaming when comparing with single-frequency actuation [30,31]. Thus, for moderate actuation voltages using the frequency modulation method, acoustic streaming is not causing any problem for the trapping efficiency and trapping stability over time.

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