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Article

A Low-Cost Laser-Prototyped Microfluidic Device for Separating Cells and Bacteria

by
Sinan Gucluer
* and
Osman Guler
Department of Mechanical Engineering, Aydin Adnan Menderes University, Aydin 09010, Türkiye
*
Author to whom correspondence should be addressed.
Appl. Sci. 2023, 13(13), 7938; https://doi.org/10.3390/app13137938
Submission received: 1 June 2023 / Revised: 4 July 2023 / Accepted: 5 July 2023 / Published: 6 July 2023
Browse Figures
Figure 1
<p>The microfluidic device is shown with: (<b>a</b>) schematic depiction; (<b>b</b>) top-view of actual picture of the device; (<b>c</b>) top-view dimensions of the device. (<b>d</b>) side-view schematic of the device is shown to illustrate the layers and tube insertion. The radius of each loop of the serpentine is 0.75 mm, radii of the spirals 1, 2, and 3 are 8.7-, 12.7-, and 16.7-mm, respectively. The actual picture of the device is injected with food dye for better visualization. (<b>e</b>) A schematic showing the working principle of the particle separation in spiral channel, F<sub>L</sub> and F<sub>D</sub> are inertial lift force and Dean drag forces, respectively.</p> "> Figure 2
<p>Particle focusing behavior is demonstrated using the combined device. (<b>a</b>) Device schematic outline and image collection regions are shown. Streak images of particle flows are formed from; (<b>b</b>) entrance; (<b>c</b>) exit of the serpentine section and; (<b>d</b>) exit of the spiral section.</p> "> Figure 3
<p>(<b>a</b>) The mixture of 1 and 5 µm polystyrene particles before entering the device. Particles collected from; (<b>b</b>) the outlet A and; (<b>c</b>) the outlet B are shown.</p> "> Figure 4
<p>The purity of the particles collected from the outlets A and B at different flow rates. Lower and higher box bounds are in the 25th and 75th percentiles, respectively; the line within the box is the median; and the lower and upper error lines are in the 5th and 95th percentiles, respectively.</p> "> Figure 5
<p>Separation efficiency for 1 and 5 µm polystyrene particles collected at different flow rates. Lower and higher box bounds are in the 25th and 75th percentiles, respectively; the line within the box is the median; and the lower and upper error lines are in the 5th and 95th percentiles, respectively.</p> "> Figure 6
<p>Separation of yeast cells and 1 µm particles. (<b>a</b>) Cell and particle mixture before separation; (<b>b</b>) Sample collected from the outlet A; (<b>c</b>) sample collected from the outlet B; (<b>d</b>) purity of samples collected from the outlets A and B. Lower and higher box bounds are in the 25th and 75th percentiles, respectively; the line within the box is the median; and the lower and upper error lines are in the 5th and 95th percentiles, respectively.</p> ">
Versions Notes

Abstract

:

Featured Application

This study is a novel, simple, and low-cost device demonstration for cell separation applications.

Abstract

Simple and rapid fabrication of microfluidic devices can enable widespread implementation of lab-on-chip devices in resource-limited environments. However, currently most of the microfluidic devices are fabricated in cleanroom facilities that are well-funded and not accessible to most of the researchers in developing countries. Herein, a simple, low-cost, and reliable method is shown to fabricate microfluidic devices for separating cells and bacteria-size microparticles. For this purpose, serpentine and spiral microfluidic channels are designed and fabricated using rapid laser prototyping. This single inlet microfluidic device is shown to successfully separate yeast cells and smaller microparticles with an efficiency of 85% which is very promising for many lab-on-chip applications including cell-based diagnostics and therapeutics.

1. Introduction

Separating bacteria and other impurities from cells is an essential step in various biomedical applications, including cell culture, drug discovery, and diagnostics [1,2]. Impurities such as bacteria, viruses, and other microorganisms can interfere with the desired biological processes and contaminate the final product, leading to inaccurate results and potential harm to the end user [3]. Therefore, effective separation techniques are crucial for ensuring the purity and safety of cell-based products.
Microfluidics is a rapidly developing field that has revolutionized cell separation techniques [4,5,6,7]. Microfluidic devices use tiny channels and chambers to manipulate fluids and particles at the microscale level, allowing for precise and efficient separation of cells from other impurities [8,9]. Microfluidic-based cell separation techniques offer several advantages over conventional methods, including reduced sample size, higher purity, and faster separation times [10,11]. There are two primary approaches to microfluidic cell separation: passive and active [12,13,14,15]. Passive methods rely on physical differences in cell size, shape, or deformability to separate cells, whereas active methods use external forces to manipulate cells. Some of the mainstream active methods are dielectrophoresis (DEP), electrophoresis, acoustic separation, optical tweezers, and magnetophoresis [16,17,18]. The DEP method applies an electric field to cells, allowing for their separation based on their electrical properties [19]. Cells with different dielectric properties experience different forces and are separated based on their response to the electric field. However, DEP and other electric field-based methods rely on the sample and medium conductivity and dielectric properties, and these methods can be detrimental to delicate biological samples due to Joule heating [20]. In acoustic-based methods, acoustic waves are used to separate cells based on their size, shape, and compressibility [2,21,22]. Cells experience different forces based on their acoustic properties, which allows for their manipulation [23]. While acoustic methods are gentler towards biological samples due to their inherited biocompatibility characteristics, acoustic systems may require bulky and expensive external equipment for functioning. Similarly, magnetic and optical methods can be quite complex and expensive due to the required external peripheral equipment [24,25,26].
Passive methods include inertial focusing, deterministic lateral displacement, hydrodynamic filtration, and viscoelastic separation [27,28,29,30]. Zeming et al. presented the asymmetrical deterministic lateral displacement (DLD) method for enhancing the separation and throughput of red blood cells (RBCs) [31]. The researchers propose a new approach by changing the ratio of lateral-gap to downstream-gap, which allows for efficient separation of non-uniformly shaped and sized RBCs while maintaining a high throughput. The study introduces the concept of the separation index (SI) as a means to compare the separation efficiency across various DLD devices. The SI is shown to have an index of over 95%, indicating the effectiveness of the asymmetrical DLD gap-size method. Passive methods including filtration-based approaches are more prone to problems depending on the cell size variations due to their lacking dynamic adjustability. One of the most promising microfluidic-based passive cell separation techniques is called inertial focusing [32,33,34]. In this method, cells are separated based on their size and shape as they flow through microchannels under the influence of fluid forces [35]. This technique is particularly useful for applications such as cell-based assays and clinical diagnostics, where high purity and efficiency are essential. Inertial cell separation methods are a group of microfluidic techniques that utilize the physical properties of cells, such as their size, shape, and deformability, to sort and separate them from other cells or impurities. Two popular types of inertial cell separation methods are serpentine channels and spiral microfluidics [36,37,38,39]. Both serpentine channels and spiral microfluidics offer several advantages over traditional cell separation methods, such as higher throughput, lower sample volume, and increased purity. These methods also require minimal external equipment and can be easily integrated with other microfluidic devices, making them ideal for a wide range of applications, such as point-of-care diagnostics, drug discovery, and cancer research. Even though the passive microfluidic methods are simpler compared to the active ones, the majority of the passive devices rely on expensive and costly cleanroom microfabrication facilities which are not readily available in developing countries [40,41]. While there are some preliminary reports on 3D-printed microfluidic devices for cell separation, more examples and demonstrations are critical to extend and advancing this field with simpler and better tools [42,43]. Therefore, more efforts are still needed to investigate low-cost fabrication alternatives to conventional microfluidic devices to achieve low-cost and accessible cell separation applications.
Microfluidic-based separation techniques have emerged as a powerful tool for the efficient isolation and sorting of cells and bacteria [44,45]. These techniques utilize the unique properties of microscale fluid flow to manipulate and separate biological entities with high precision and throughput. Numerous studies have demonstrated the effectiveness of microfluidic platforms in various applications, including cell sorting, cell enrichment, and bacterial analysis. For example, in one study by Ai et al., a microfluidic device was developed for the separation of heterogeneous particles or cell mixtures in a continuous flow using acoustophoresis [46]. In this device, a standing surface acoustic wave (SSAW) field was created across a microchannel by propagating two identical surface acoustic waves (SAWs) generated by interdigital transducers (IDTs). The developed device successfully separated two types of fluorescent microspheres with different sizes. Furthermore, Escherichia coli bacteria mixed with peripheral blood mononuclear cells (PBMCs) were efficiently isolated using the SSAW-based separation technique. Flow cytometric analysis of the collected samples indicated a purity of 95.65% for the separated E. coli bacteria.
In another study, a label-free, size-based bacteria separation method from whole blood was achieved using elasto-inertial microfluidics [47]. In elasto-inertial microfluidics, the migration of blood cells into a non-Newtonian solution is facilitated by viscoelastic flow, while smaller bacteria remain in the streamline at the blood sample entrance and can be effectively separated. The flow conditions were optimized using particles, resulting in a continuous separation of 5 μm particles from 2 μm particles with yields of 95% and 93% at respective outlets. Subsequently, bacteria were continuously separated from undiluted whole blood samples with an efficiency of 76%. Even though these examples demonstrate good separation efficiencies, the fabrication of the device requires lengthy and expensive procedures that limit the application of such devices in rapid disease screenings. Such wide-scale applications require low-cost single-use, disposable devices. Thus, to enable disposable and affordable point-of-care microfluidic devices, alternative fabrication approaches should be explored to further prove the feasibility of low-cost manufacturing for lab-on-chip applications.
In this work, we demonstrate a simple and low-cost microfluidic device fabricated by implementing rapid laser prototyping using polymethyl methacrylate (PMMA). The device is designed to have a single inlet, a serpentine section, and a spiral microfluidic channel to achieve sheath-less cell separation. As a proof-of-concept application, yeast cells and 1 µm diameter polystyrene particles are separated in this microfluidic device with 85% separation efficiency. Considering the simplicity and low cost, the performance of the device is promising for many lab-on-chip applications.

2. Materials and Methods

2.1. Device Fabrication and Operation

The microfluidic device is designed to have a single inlet, a serpentine focusing section, and a spiral separation section with two outlets as shown in Figure 1. As for the serpentine section, similar designs in the literature were considered in terms of the number and dimensions of the serpentines that result in satisfactory particle focusing [33]. The spiral section was designed to have three loops with at least 2 mm distance between each loop to provide rigidity which is important during the assembly of the device. The three layers of the device were laser-cut and assembled via double-sided transparent tape (3M, Maplewood, MN, USA) on the bottom and top layers. A standard 100W CO2 laser plotter (LF7010, Lazerfix, Konya, Türkiye) was used for PMMA cutting. This machine has a working area of 70 × 100 cm and features a glass CO2 tube laser. The control unit of the system is a Ruida 6442 module. Motion in the x and y axes is provided by Nema 34 stepper motors. The maximum feed rate of the laser plotter is 500 mm/s, and precision of the system is 0.01 mm/meter. The PMMA sheets in this study were cast acrylic sheets that provide higher optical clarity and better laser cutting compared to extruded acrylic sheets. Thicknesses of the bottom layer, the middle layer, and the top layer were 300 µm, 100 µm, and 3 mm, respectively. The top layer was chosen to be much thicker to provide at least 2 mm depth for the inlet and outlet tubes. Polyethylene tubings with a 1 mm outer diameter and 0.5 mm inner diameter were used for the inlet and outlet of the device. A five-minute epoxy glue (E340, Akfix, Istanbul, Turkey) was used to permanently fix the tubings to provide a better fluid-tight fitting that could resist high-pressure flow.
For testing the device, 5 and 1 µm polystyrene particles (Sigma Aldrich, St. Louis, MO, USA) were used with 107 and 3 × 107 particles/mL, respectively. For simulating cell and bacteria separation, Saccharomyces cerevisiae yeast cells (about 107 cells/mL) and 1 µm polystyrene particles were used. Microparticles with similar sizes to bacteria were used mimic bacteria in the separation experiments. For the characterization of particle and cell counts, a standard hemocytometer was employed. For this, the sample containing the microparticles was diluted five times to achieve a lower concentration for easier characterization. The hemocytometer was placed under a microscope, and the grid lines within the counting chamber were focused upon. The microparticles were observed and counted using both automated and manual approaches for cross-check [48,49]. For automated counting, images were thresholded and processed to better visualize the particles [48].
A home-built syringe pump system was used to inject the cell/particle solution [50]. An inverted microscope (OX.2053-PLPH, Euromex, Arnhem, The Netherlands) and CMOS camera (HD, Euromex, Arnhem, The Netherlands) were used to capture images. For the yeast cells, the yeast growth medium was prepared by dissolving 10 g yeast extract, 20 g peptone, and 20 g dextrose in 1 L of sterile distilled water. The Saccharomyces cerevisiae yeast culture was inoculated into 10 mL of the growth medium and incubated at 30 °C with mild shaking. The yeast cells were then collected by centrifugation at 4000× g for 5 min. The supernatant was discarded, and the yeast cells were resuspended in 10 mL of sterile distilled water.

2.2. Working Mechanism of the Device

The theory of inertial particle separation, also known as the Dean effect or Dean flow, explains how particles of different sizes behave when flowing through a curved channel [51,52]. Dean forces increase with the increasing curvature, flow rate, and channel size [53]. Particles can be moved towards the inner or outer wall of a microchannel depending on their size and the drag force they experience, which is caused by the Dean vortices. Pressure and inertial lift also contribute to the forces experienced by particles in a curvilinear channel. Particles tend to occupy equilibrium positions where the opposite lift forces are balanced, resulting in a net lift force acting on them. These equilibrium positions also form narrow bands. Particles can be focused in serpentine channels by utilizing the inertial forces that act on the particles as they flow through the curved sections of the channel [34,54]. In particular, when a fluid containing particles flows through a curved section of the channel, inertial forces act on the particles causing them to move towards the outer wall based on their sizes. As the fluid flows into a straight section of the channel, the particles do not move laterally due to the balancing hydrodynamic forces. This results in the particles being focused towards the center of the channel. By arranging multiple curved and straight sections in a serpentine pattern, particles can be continuously focused based on their size. In the spiral section of the channel, dean flow, inertia forces, and drag forces are balanced at specific locations in the channel for particles based on their sizes. It is known from the literature and the inertial focusing in spiral channels that larger particles move close to the inner wall and smaller particles are pushed to the outer wall of the microfluidic channel as shown in Figure 1e [14,32,33]. In this work, the serpentine section of the microfluidic channel is used to focus the particles at equilibrium positions before they enter the spiral microfluidic channels. This way, different diameter particles can be separated in the spiral channel with only a single inlet for sample injection without using a sheath flow.

3. Results

A mixture of 1 and 5 µm polystyrene particles was infused into the microfluidic device at varying flow rates. Figure 2 shows the streak images obtained from the entrance, the end of the serpentine section, and the exit of the spiral section of the device for the inlet flow rate of 1000 µL/min. It was observed that the polystyrene particles entered the device with random distribution across the width of the channel at the entrance (Figure 2b). At the end of the serpentine section, polystyrene particles seemed to form a visible focused stream (Figure 2c). This was in accordance with the results reported in the literature for the similar device geometries [33,37,54]. This focusing behavior was intended to be used before the spiral channel to eliminate the need for a hydrodynamic flow focusing which would require multiple inlets and sheath flow. After the serpentine section, focused particles entered the spiral section of the microfluidic device to be focused towards different streams based on their dimensions (Figure 2d). A darker stream appeared close to the inner wall which is formed mostly by 5 µm particles.
For a quantitative analysis of the particle separation performance of the device, different inlet flow rates were applied, and particles collected from outlets A and B were analyzed. It was found that inlet flow rates between 600 and 1000 µL/min resulted in much higher separation purity and efficiencies. An example of particle pictures is shown in Figure 3 for 1000 µL/min inlet flow rate. It is clear that the majority of the 1 µm sized polystyrene particles exited the device from outlet A and 5 µm particles exited from the outlet B. This is an expected outcome based on the predictions from the inertial focusing behavior of larger particles in the spiral microfluidic channels that is also in agreement with the literature.
Next, the collected particle numbers were characterized by standard hemocytometer counting. In Figure 4, the purities of samples collected from the outlet A and B are shown. Here, purity was calculated by analyzing the number of 1 and 5 µm polystyrene particles collected from each outlet and calculating the ratio of the particles. By trying different inlet flow rates, improved purities were observed at 600, 800, and 1000 µL/min. It was found that the best performance in terms of purity of the collected particles was achieved at 800 µL/min flow rate for this device design. For the outlets A and B, the purities of 1 and 5 µm particles were found to be 93% and 89%, respectively. The outlet B was found to be more contaminated by 1 µm particles compared to the contamination of the outlet A by the 5 µm particles. This was due to the higher degree of dispersion of the 1 µm particles compared to the larger 5 µm ones. Larger particles were better focused by the inertial and dean flows. The efficiency of particle separation was also characterized by the collected samples as shown in Figure 5. Efficiency is defined as the ratio of the number of the collected targeted particles divided by the total number of the target particles. It was found that the separation efficiencies of 1 and 5 µm were 89% and 93%, respectively. Due to their larger size and better focusing behavior, most of the 5 µm particles were collected from the intended outlet of the device. Below 600 µL/min flow rate, particle separation efficiency and outlet purities were lower which is likely due to inertial effects that resulted in poor particle focusing. Above 1000 µL/min flow rate, we also observed a monotonic decrease in separation efficiency and eventually device fails above 3000 µL/min due to significant delamination. It is likely that a small amount of leaking of the fluid outside the channel, which might be starting above 1000 µL/min, negatively affects the separation efficiency due to deterioration of inertial focusing.
Finally, yeast cells and 1 µm particles were separated in the device to simulate the separation of bacteria and cells for cell therapy applications. Yeast cells were chosen for this work because they are easy to grow and maintain in an ordinary lab and demonstrate living organism physical properties. 1 µm sized polystyrene particles are close to the size of most of the bacteria so they are used to depict bacteria. Figure 6a shows the mixture of the yeast cells and particles before they are separated. Figure 6b,c shows the samples collected from the outlets A and B, respectively. Separation purities of yeast cells and 1 µm particles are given as 91% and 85% in Figure 6d. Due to their lower homogeneity and variation in sizes, yeast cells yielded a decreased purity compared to the 5 µm particles. The average separation efficiencies of the yeast cells and 1 µm particles were also found as 85% and 88%, respectively.

4. Discussion

The presented device features a single inlet for sample injection and device operation which provides simplicity in terms of required syringe pumps and reduces the footprint of the separation process. In many of the cell/bacteria separation works reported in the literature sheath flows are generally required for sample focusing and separation. In our device, the first module (serpentine) did not provide an observable particle separation, but it served the purpose of focusing the particles before they entered the spiral section of the device. In the spiral section of the device, larger particles stayed mostly close to the inner wall. This is due to the fact that Dean drag acts and pushes particles towards the inner wall and for larger particles a larger force is exerted which pushes the larger particles further than the smaller particles.
Different inlet flow rates were tested to quantitatively analyze the particle separation performance of the device, with particles collected from outlets A and B being analyzed. Inlet flow rates ranging from 600 to 1000 µL/min demonstrated higher separation purity and efficiency. At 1000 µL/min, the majority of 1 µm particles exited from outlet A, while 5 µm particles exited from outlet B, in line with predictions based on inertial focusing behavior. Standard hemocytometer counting was performed to characterize the collected particles. Improved purities were observed at flow rates of 600, 800, and 1000 µL/min, with the best performance at 800 µL/min. Purity percentages were 93% for 1 µm particles and 89% for 5 µm particles from outlets A and B, respectively. Outlet B exhibited a higher level of contamination from 1 µm particles in comparison to the contamination observed in outlet A caused by 5 µm particles. This discrepancy can be attributed to the greater dispersion exhibited by the 1 µm particles in contrast to the larger 5 µm particles. It is worth noting that larger particles tend to be more effectively focused by the inertial and Dean flows. Separation efficiencies were 89% for 1 µm particles and 93% for 5 µm particles. The device was also tested for separating yeast cells and 1 µm particles, achieving purities of 91% for yeast cells and 85% for 1 µm particles, with separation efficiencies of 85% and 88%, respectively. While these are separation efficiencies comparable to the results reported in the literature, the simplicity and lower cost of our device fabrication and operation are significant factors that can greatly contribute to the ecosystem of disposable on-chip applications. The presented device only costs about 2 US dollars including the operational costs. It can also be possible to further reduce this cost by batch fabrication of larger numbers of devices. This is significantly lower than traditional microfluidics devices considering the cost of using clean-room fabrication facilities.
An important limitation of the presented device is delamination at high flow rates due to increased pressure inside the channels. We observed that over about 3000 µL/min delamination started, and the device failed. It is also important to note that over about 2000 µL/min, slight observable leakage of the fluids occurred outside the fluidic channels. As a matter of fact, separation efficiency monotonically decreased above 1000 µL/min, and this is likely due to the emergence of leaking that negatively affected inertial focusing of particles. This limitation can be overcome with better bonding strategies. For example, chemical bonding of PMMA layers can be implemented using suitable setups.
Low-resourced environments greatly benefit from the use of low-cost and disposable microfluidic devices, highlighting their significance. These devices offer numerous advantages that address the unique challenges faced in such settings. Firstly, their affordability ensures accessibility, allowing healthcare providers in low-resourced areas to utilize advanced diagnostic and analytical tools without incurring significant costs. Additionally, their disposable nature eliminates the need for complex and expensive cleaning and sterilization processes, reducing the risk of contamination and the burden on limited resources. Moreover, these devices are often designed to be portable and user-friendly, enabling point-of-care testing and rapid disease screening in remote and underserved regions. By providing cost-effective and user-friendly solutions, low-cost and disposable microfluidic devices play a crucial role in enhancing healthcare delivery and disease management in low-resourced environments.
The separation of bacteria from blood and other cell solutions holds immense significance in biomedical applications and rapid disease screening, particularly in low-resource areas. Efficiently isolating bacteria from these complex samples enables a deeper understanding of their characteristics, behavior, and response to different treatments. By studying bacteria separately, researchers and healthcare professionals can develop targeted therapies, design effective antibiotics, and enhance patient outcomes. Moreover, the ability to rapidly screen and identify bacteria in low-resource settings is crucial for timely disease diagnosis and management. Separating bacteria from blood and cell solutions in these settings allows for cost-effective and accessible screening methods that can aid in the early detection and control of infectious diseases, reducing their spread and mitigating public health risks. This capability is especially crucial in resource-limited areas where access to advanced diagnostic technologies and facilities may be limited, enabling timely interventions and improving overall healthcare outcomes.
Low-cost PMMA (Polymethyl methacrylate) passive microfluidic cell separation devices have the potential for various applications [55]. One prominent application is in the field of biomedical research, where these devices can be utilized for the isolation and separation of specific cell populations from complex mixtures [56]. Such devices can aid in studying cellular interactions, analyzing cellular characteristics, and investigating disease mechanisms. For example, Dean flow-based cancer cell enrichment was demonstrated using microfluidic devices fabricated through traditional methods [57]. These rare cell enrichment studies can be undertaken by low-cost devices. In addition, these low-cost devices can be employed in clinical diagnostics, enabling the efficient and affordable detection and analysis of various diseases and infections. They can facilitate the separation of target cells, such as circulating tumor cells, from biological samples, allowing for early disease diagnosis and monitoring. Furthermore, these microfluidic devices can be utilized in resource-limited settings for point-of-care testing, enabling rapid and cost-effective disease screening and monitoring. The versatility and cost-effectiveness of low-cost PMMA passive microfluidic cell separation devices make them promising tools for various biomedical applications, offering potential advancements in research, diagnostics, and healthcare delivery.

5. Conclusions

In this work, a low-cost and simple method to fabricate a microfluidic device for inertial particle separation was designed, fabricated, and tested. The device consisted of a single inlet serpentine focusing section and a spiral separation section with two outlets. The device was tested with 1 and 5 µm polystyrene particles and Saccharomyces cerevisiae yeast cells. The results showed successful separation of particles with a single inlet for sample injection without using a sheath flow. With the presented device, yeast cells are separated with an 85% efficiency which is very promising considering the simplicity of fabrication and implementation compared to traditional devices that rely on cleanroom fabrication. Further improvements of the device geometry and optimizations of fluidic parameters can be studied to improve the performance of the device further. The fabricated device is compact, easy to assemble, and has great potential for lab-on-a-chip applications.
This work demonstrates that very simple approaches can be implemented to fabricate low-cost and disposable microfluidic devices that can perform cell and bacteria separation with good efficiencies. In the future work, device geometry and fluidic parameters can be better optimized to further improve the separation efficiencies. Moreover, other passive microfluidic cell separation methods such as DLD and viscoelastic focusing can be incorporated together to obtain higher efficiencies.

Author Contributions

Conceptualization, S.G. and O.G.; methodology, S.G. and O.G.; formal analysis, S.G. and O.G.; investigation, S.G. and O.G.; resources, S.G.; data curation, S.G. and O.G.; writing—original draft preparation, S.G. and O.G.; writing—review and editing, S.G. and O.G.; supervision, S.G. All authors have read and agreed to the published version of the manuscript.

Funding

This research received no external funding.

Institutional Review Board Statement

Not applicable.

Informed Consent Statement

Not applicable.

Data Availability Statement

Not applicable.

Acknowledgments

We thank Erturan Yetiskin for his help with device fabrication. We also thank Fatih Akkoyun for his help and guidance with image processing for quantification of particle numbers.

Conflicts of Interest

The authors declare no conflict of interest.

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Figure 1. The microfluidic device is shown with: (a) schematic depiction; (b) top-view of actual picture of the device; (c) top-view dimensions of the device. (d) side-view schematic of the device is shown to illustrate the layers and tube insertion. The radius of each loop of the serpentine is 0.75 mm, radii of the spirals 1, 2, and 3 are 8.7-, 12.7-, and 16.7-mm, respectively. The actual picture of the device is injected with food dye for better visualization. (e) A schematic showing the working principle of the particle separation in spiral channel, FL and FD are inertial lift force and Dean drag forces, respectively.
Figure 1. The microfluidic device is shown with: (a) schematic depiction; (b) top-view of actual picture of the device; (c) top-view dimensions of the device. (d) side-view schematic of the device is shown to illustrate the layers and tube insertion. The radius of each loop of the serpentine is 0.75 mm, radii of the spirals 1, 2, and 3 are 8.7-, 12.7-, and 16.7-mm, respectively. The actual picture of the device is injected with food dye for better visualization. (e) A schematic showing the working principle of the particle separation in spiral channel, FL and FD are inertial lift force and Dean drag forces, respectively.
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Figure 2. Particle focusing behavior is demonstrated using the combined device. (a) Device schematic outline and image collection regions are shown. Streak images of particle flows are formed from; (b) entrance; (c) exit of the serpentine section and; (d) exit of the spiral section.
Figure 2. Particle focusing behavior is demonstrated using the combined device. (a) Device schematic outline and image collection regions are shown. Streak images of particle flows are formed from; (b) entrance; (c) exit of the serpentine section and; (d) exit of the spiral section.
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Figure 3. (a) The mixture of 1 and 5 µm polystyrene particles before entering the device. Particles collected from; (b) the outlet A and; (c) the outlet B are shown.
Figure 3. (a) The mixture of 1 and 5 µm polystyrene particles before entering the device. Particles collected from; (b) the outlet A and; (c) the outlet B are shown.
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Figure 4. The purity of the particles collected from the outlets A and B at different flow rates. Lower and higher box bounds are in the 25th and 75th percentiles, respectively; the line within the box is the median; and the lower and upper error lines are in the 5th and 95th percentiles, respectively.
Figure 4. The purity of the particles collected from the outlets A and B at different flow rates. Lower and higher box bounds are in the 25th and 75th percentiles, respectively; the line within the box is the median; and the lower and upper error lines are in the 5th and 95th percentiles, respectively.
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Figure 5. Separation efficiency for 1 and 5 µm polystyrene particles collected at different flow rates. Lower and higher box bounds are in the 25th and 75th percentiles, respectively; the line within the box is the median; and the lower and upper error lines are in the 5th and 95th percentiles, respectively.
Figure 5. Separation efficiency for 1 and 5 µm polystyrene particles collected at different flow rates. Lower and higher box bounds are in the 25th and 75th percentiles, respectively; the line within the box is the median; and the lower and upper error lines are in the 5th and 95th percentiles, respectively.
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Figure 6. Separation of yeast cells and 1 µm particles. (a) Cell and particle mixture before separation; (b) Sample collected from the outlet A; (c) sample collected from the outlet B; (d) purity of samples collected from the outlets A and B. Lower and higher box bounds are in the 25th and 75th percentiles, respectively; the line within the box is the median; and the lower and upper error lines are in the 5th and 95th percentiles, respectively.
Figure 6. Separation of yeast cells and 1 µm particles. (a) Cell and particle mixture before separation; (b) Sample collected from the outlet A; (c) sample collected from the outlet B; (d) purity of samples collected from the outlets A and B. Lower and higher box bounds are in the 25th and 75th percentiles, respectively; the line within the box is the median; and the lower and upper error lines are in the 5th and 95th percentiles, respectively.
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Gucluer, S.; Guler, O. A Low-Cost Laser-Prototyped Microfluidic Device for Separating Cells and Bacteria. Appl. Sci. 2023, 13, 7938. https://doi.org/10.3390/app13137938

AMA Style

Gucluer S, Guler O. A Low-Cost Laser-Prototyped Microfluidic Device for Separating Cells and Bacteria. Applied Sciences. 2023; 13(13):7938. https://doi.org/10.3390/app13137938

Chicago/Turabian Style

Gucluer, Sinan, and Osman Guler. 2023. "A Low-Cost Laser-Prototyped Microfluidic Device for Separating Cells and Bacteria" Applied Sciences 13, no. 13: 7938. https://doi.org/10.3390/app13137938

APA Style

Gucluer, S., & Guler, O. (2023). A Low-Cost Laser-Prototyped Microfluidic Device for Separating Cells and Bacteria. Applied Sciences, 13(13), 7938. https://doi.org/10.3390/app13137938

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