Review Article: Recent advancements in optofluidic flow cytometer

Abstract

There is an increasing need to develop optofluidic flow cytometers. Optofluidics, where optics and microfluidics work together to create novel functionalities on a small chip, holds great promise for lab-on-a-chip flow cytometry. The development of a low-cost, compact, handheld flow cytometer and microfluorescence-activated cell sorter system could have a significant impact on the field of point-of-care diagnostics, improving health care in, for example, underserved areas of Africa and Asia, that struggle with epidemics such as HIV∕AIDS. In this paper, we review recent advancements in microfluidics, on-chip optics, novel detection architectures, and integrated sorting mechanisms.

Figure 1

Schematic of a FACS system that can detect two scattered and two fluorescent light signals. It consists of (1) a fluidic system, (2) an optical system (illumination and detection), (3) a sorting system, and (4) an electronic control system (Ref. 9).

Figure 2

Sample flow from the center inlet is focused by sheath flow to 20 μm width. The flow rate ratio of sample flow to sheath flow is 1:10. (Ref. 9).

Figure 3

The schematic design of (a) a chevron-patterned device (Ref. 26) and (b) the contraction-expansion array (Ref. 27) for three-dimensional flow focusing.

Figure 4

Particles experience two lateral forces, wall effect lift (F_LW) and shear gradient lift force(F_LS), resulting in migration to the lateral equilibrium positions (X_eqaway from the top and bottom surfaces) inside the microfluidic channel (Ref. 29).

Figure 5

Schematic of the inertial focusing process in an asymmetrically curved channel followed by a straight channel. The combination of the curved and the straight channels in series biases the particles in the fluid to one half of the channel. As a result, the particles are focused to a single vertical streamline within the straight channel (Ref. 33).

Figure 6

A variety of methods to incorporate optical elements: (a) integrated waveguides forming a y-splitter in PDMS (Ref. 48), (b) a lensed waveguide facet for altering beam divergence (Ref. 45), (c) a fiber sleeve filled with PDMS prepolymer for waveguiding, and (d) the same sleeve with an optical fiber inserted (Ref. 39).

Figure 7

Freestanding integrated lenses: (a) air-filled lenses created in PDMS (Ref. 51), (b) image of light tracing (via alumina scattering centers) from a fluid-filled lens in PDMS, and (c) the corresponding optical modeling for the lens based on shape and refractive index (Ref. 46).

Figure 8

Schematic of a microfluidic cytometer design based on exclusion optics. Light originating from either of the two interrogation points (circles in fluidic channel) couples through a flat facet into the collection waveguides. Light originating between these points (or outside of them) is incident on an angled facet, experiencing significant reflection losses as well as a path change that will drastically reduce waveguide coupling (Ref. 63).

Figure 9

Schematic illustration of COST coding technology device structure. Each fluorescence color is first temporally coded by the spatial pattern of the first three beam blocking apertures. Next, the color is encoded by the red, green, and blue color filters and the associated spatial filter (apertures). The COST-coded color fingerprint is unique to each color, enabling multicolor registration by a single PMT (Ref. 62).

Figure 10

(a) COST-coded output waveform of dragon green fluorescence. The first three space-time coded signals are followed by three color-coded peaks of different intensities. (b) Histogram of the green filtered COST-coded signal normalized to the red filtered signal of dragon green and envy green fluorophores. (c) Histogram of the blue filtered COST-coded signal normalized to the red filtered signal of dragon green and envy green fluorophores (Ref. 62).

Figure 11

Electro-osmosis-induced particle deflection. After hydrodynamic focusing, the highly concentrated fluorescent beads are deflected to the right or left collection channel, depending on the polarity of the applied voltage (Ref. 70).

Figure 12

Superimposed images of 200 particles passing through the sorting junction. After dielectrophoretic focusing, ∼100particles are deflected to each of the collection channels (Ref. 72).

Figure 13

Principle of an optical cell sorter. After sample focusing, cells are analyzed in the analysis region based on their fluorescence. The detected fluorescence triggers laser-directed cell manipulation, causing targeted cells to be sorted to the collection channel. All the nontargeted cells flow to the waste channel (Ref. 73).

Figure 14

Hydrodynamic flow switching using an external check valve. Upon detection of fluorescence from the sample, the electronic feedback system triggers the opening/closing of the check valve, causing sudden redirection of the fluid flow (Ref. 24).

Figure 15

Sorting mechanism of the piezoelectric (PZT)-actuated cell sorter. The voltage-induced PZT bending causes temporary fluid displacement (to the right or left), leading to deflection of targeted particles down to either side of the collection channels. The polarity and magnitude of the input voltage control the direction and magnitude of the deflection, respectively. Without PZT actuation, nontargeted cells exit directly down the waste channel (Ref. 56).

Figure 16

Schematics showing the external and internal workings of the integrated μFACS system. (a) Schematics of the setup showing on-chip illumination, spatial filter (transparency mask) modulated fluorescent detection, and FPGA-embedded electronic control system. (b) The spatial filter is placed at the image plane of the device, resulting in modulated fluorescence signals detected by the PMT when a fluorescent cell passes the detection zone (producing a 111 signal) and the postsorting region (producing a 1101 or 1011 signal, depending on the collection channel). (c) The FPGA-implemented real-time process control unit enables real-time signal amplification (∼18 dB Signal to Noise Ratio enhancement) of the modulated signals by using a match filter. The control unit outputs a time-delayed waveform (user-defined) to trigger PZT actuation when the detected signals reach the user-defined threshold. (d) An example of space-time coded signal by the spatial filter of (b). The 111 coded signal at the detection zone is followed by the 1011 coded signal, confirming a successful sorting event (Ref. 61).

Figure 17

Purity analysis of mammalian cell sorting. Before sorting (left histogram), the initial mixture ratio of nonfluorescent to fluorescent cells is∼1:150. After sorting by theμFACS, the final mixture ratio is ∼1.86:1 showing an enrichment factor of 230-fold (Ref. 61).