# Finger-Actuated Micropump of Constant Flow Rate without Backflow

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## Abstract

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## 1. Introduction

- Easy to fabricate by using the Embedded Scaffold Removing Open Technology (ESCARGOT) method.
- Controlling the total deformation of the flexible diaphragm until the bottom chamber of a finger-actuated micropump allows for indirect regulation of the volume of fluid moving.
- A combination of Tesla valves, serpentine microchannels, and hydrogel-assisted reservoirs can increase the reverse flow resistance, in other words, no backflow can be obtained. This can reduce the risk of hygiene and make sensing reading redundant.

#### 1.1. Proposed Concept Design

- Take note that the extraction rate should not go above 1–2 mL/min to prevent discomfort, pain, distortion, and tearing of the tissue [67,68,69]. Hence, the low extraction rate (from 10 µL/min to a few hundred microliters) is painless [69] and suitable for transdermal application [70,71]. 1 μL/min flow rate is appropriate flow rate for maximizing recovery factor (RF) and obtaining high accuracy measurement in detection biosensor [72]. The recovery factor (RF), which is the ratio of the glucose concentration in the extracted ISF to the fresh ISF in the skin, can be calculated using Equation (1). In cases where the flow rate is less than 1 μL/min, although the recovery factor is greater, a longer time is taken to reach the sensor, and a higher sensitivity of the sensor is required [73]. At high flow rates, the working electrodes of the sensor cannot fully interact with the glucose, making the sensor less accurate at detecting glucose [74,75,76].
- The ability to maintain a consistent flow rate enables stable sensing signals and data transmission for real-time analysis, judgement, and treatment decision-making.
- It should be consumed in the smallest possible volume for diagnosis within the desired flow rate range. Continuous monitoring, such as measuring glucose levels in a diabetic patient, is challenging to carry out since it requires a sampling process at a high frequency within a day, and a large sample size is impractical. As long as the sensor is able to analyze the relevant analytes of interest and produce a sensing signal using a minimal sampling size that is applied to the sensing reaction area, it is appropriate for the overall system. Laboratory techniques may require a sample volume of >100 μL for precise analysis and lengthy testing times. However, in practice, most diagnosis devices extract a small volume of 1–10 μL of ISF [17,77] and 0.6–1.1 μL of blood sampling size for biosensors [78], respectively, within a minute [79]. A adequate of ISF sample can produce an accurate and reliable measurement, so a larger volume of sample fluid than is necessary to fully saturate the sensing region’s reaction area is recommended. Incorrect measurements can lead to wrong diagnoses and ineffective treatments.
- It should have robust performance under various conditions and be economical. High manufacturing costs are a barrier to the widespread use of microfluidic devices since they are required to integrate multiple components to achieve desired results. Single-use devices (disposables) have had an overwhelmingly large impact on society and the economy compared to applications that consider continuous long-term monitoring strategies.
- No backflow to avoid hygiene or incorrect analysis due to repeated use of the same sample
- Leakage-free

^{−12}m

^{3}. Based on the given ${\mathrm{Q}}_{(\mathrm{n}=120)}$, ideally, ISF will take 0.8-s (~0.014 min) to fill up the inlet cavity and reach the sensor for analysis. Assuming the space between the MNs array and the inlet is very small, the volume gap is negligible. The sensor region area will connect with the working electrode of reverse iontophoresis to enhance the extraction process. However, due to Fused Deposition Modeling (FDM) limitation, the prototype microfluidic device design is scaled up 10 times higher from original design and the scale up volumetric intended volumetric, ${*\mathrm{Q}}_{(\mathrm{n}=120)}$ is $3.0\times {10}^{-11}{\mathrm{m}}^{3}/\mathrm{s}$ (1.8 μL/min).

#### 1.2. Theoretical Analysis

#### 1.2.1. Consistent Flow Rate Analysis

^{−3}Pa.s is used [36]. ISF has a density of 1000 kg.m

^{−3}[93] and a surface tension of 70 dyn/cm (0.07 N/m) [94].

#### 1.2.2. Backflow Analysis

- ISF used as working fluid is Newtonian fluid
- The fluid flow is laminar (Re < 2300) [89], incompressible fluid flow $(\nabla \stackrel{\mathrm{\u20d1}}{\mathrm{V}}=0;\frac{\partial \mathrm{u}}{\partial \mathrm{x}}+\frac{\partial \mathrm{v}}{\partial \mathrm{y}}+\frac{\partial \mathrm{w}}{\partial \mathrm{z}}=0)$ in which the fluid properties, such as its density and viscosity, are constant; steady flow $(\frac{\partial \mathsf{\rho}}{\partial \mathrm{t}}=0)$ in which density does not change as a function of time
- There will be no leakage or additional fluid in the system; in other words, the volumetric mass flow rate will be the same everywhere in the system ($\mathrm{Qin}=\mathrm{Qout}$)
- Isothermal (Temperature = constant)
- Fully developed $\left(\frac{\mathrm{dw}}{\mathrm{dt}}=0\right)$. A constant velocity profile is maintained by the flow when the pressure gradient and shear forces are equalized. The z-axis pressure gradient remains constant [96].
- Velocity along z-axis is function of x and y. ($\mathrm{u}=0,\mathrm{v}=0$, w $\ne 0)$
- Fluid flow moving horizontally, thus gravity effect is negligible (g = 0)

_{h}) for rectangular microchannel. Re, Dh, and Le described in Equations (8)–(10). The Darcy friction factor (f) for laminar flow [98,99] with a function of pressure drop expressed in Equation (11).

_{L}is the local loss coefficient, which can be obtained according to the local loss coefficient table and V is average velocity.

_{c}and K

_{e}can be obtained from ${\mathrm{K}}_{\mathrm{e}\mathrm{or}\mathrm{c}}={(1-\frac{{\mathrm{A}}_{1}}{{\mathrm{A}}_{2}})}^{2}$. The pressure drop is the difference in total pressure between the two points. Diodicity (Di) is a metric used to assess the performance of a microfluidic device in both forward and reverse liquid flow conditions, which is also called the ratio of flow resistances. Di can be calculated by Equation (24)

## 2. Materials and Methods

#### 2.1. Material

#### 2.2. Simulation and Their Boundary Condition

#### 2.3. Swelling Ratio Profile, Absorption Rate and Water Retention of Hydrogel

#### 2.4. Fabrication of Microfluidic Device

#### 2.5. Experimental Microfluidic System Versatility Confirmation

## 3. Result and Discussion

#### 3.1. Head Loss, Pressure Drop and Diodicity Analysis

#### 3.2. Hydrogel-Assisted Reservoir for Interstitial Fluid Collector and Backflow Reduction

_{(n=1)}= 65 ± 3). The relative SR in a droplet of saline solution for the 4-min experimental value fits the trend of the Voigt model. Based on a single hydrogel bead, it is also important to note that the bead swelled, indicating absorption once it came into contact with the solution. There were three distinct phases to the absorption behavior: in the first phase (1 min), the amount of water absorbed increased rapidly; in the second phase (2–4 min), absorption slowed down to a moderate speed; and in the third phase (4 min), absorption reached an almost steady state with no further increases in water absorbency (Figure 8A). To examine the swelling performance of a larger quantity of hydrogel, the hydrogel sample is immersed in saline solution via a teabag. After being submerged in saline solution until equilibrium swelling is achieved, the weight of the hydrogel sample (0.22 g) increased to 31.95 g, which is equivalent to a SR of 144 g/g. The obtained SR within the other water-absorbent hydrogel from past studies (80–5520 g/g) [131,132,133,134,135,136]. This is due to the fact that the superabsorbent hydrogel possesses multiple hydrophilic macromolecules and is composed of crosslinked networks that acquire coil conformation in their dry state, but they have the capability to absorb water up to 500 times their own weight [137] and significantly expand their structure to a large size once they are exposed to water or an aqueous medium [138]. In a 20 mL saline, the absorption rate is 83.3 μL/min. The obtained flow rate for this work is almost identical to the flow rate obtained in a previous study by Seo et al. [59] (~80 μL/min and an absorption volume of ~20 mL).

- Effect of reservoir entrance quantity on the static area

- 2.
- Effect of diffuser width, b on the static area

#### 3.3. Characterization of Microfluidic Fabrication in the ESCARGOT Process

#### 3.4. Experimental Versatility

_{max}= 47 μL/min|v

_{min}= 2 μL/min), shown the consistency is lower with large standard deviation ($\sigma =13.82)$. This is due to the fact that it is difficult to develop consistent and fully deformation on the diaphragm. Thus, the output pressure became non-uniform under fast actuation [32]. The consistency improved by increase the duty cycle. Under duty cycle of 5 s, the average flow rate is 7.1 μL/min (v

_{max}= 13 μL/min|v

_{min}= 1.3 μL/min|$\sigma =3.92)$. Under duty cycle of 10 s, the average flow rate is 7.6 μL/min (v

_{max}= 12 μL/min|v

_{min}= 4 μL/min|$\sigma =1.84$). Under duty cycle of 20 s, the average flow rate is 2.2 μL/min (v

_{max}= 4.0 μL/min|v

_{min}= 1.4 μL/min|$\sigma =1.84$). Based on this result, duty cycle of 20 s or more is the optimum condition for intended microfluidic system which posse’s lowest standard deviation.

^{−9}m

^{3}) and reach the sensor at the bottom of the inlet. The measured flow rate is 2.2 μL/min which is slightly higher than the scaled-up flow rate under ${*\mathrm{Q}}_{(\mathrm{n}=120)}=3.0\times {10}^{-11}{\mathrm{m}}^{3}/\mathrm{s}$ (1.8 μL/min), with an error gap of about 22%. Nonetheless, the measured flow rate value still meets the transdermal for medical application. In the prototype device, the volume of extracted saline is about 2 μL (thickness layer = 449 μm) available to be used for the sensor at one pressing. A single pressing generates a dimensionless length of 41.6 mm in the rectangular channel. Haldkar et al. [142] reported volumetric flow rate of 3.4 μL/s (240 μL/min) at their chamber and 0.471 μL/s (28.26 μL/min) available at biosensor location., which is higher than our prototype. This is due to the fact that our prototype placed the sensor at the bottom of the inlet, while theirs is placed 1 mm away from the MNs. It will be stored 3.16 mL with flow rate of 2.2 μL/min in the hydrogel reservoir within 24 h.

## 4. Conclusions

## Author Contributions

## Funding

## Acknowledgments

## Conflicts of Interest

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**Figure 1.**Schematic illustration of the microfluidic device: (

**A**) General overview; (

**B**) Inlet connected with sensor and microneedle array; (

**C**) Finger actuated micropump. (Points (1)–(9) indicated the point-to-point components in the microfluidic system, and the red arrow indicated ISF moving flow).

**Figure 6.**(

**A**) Total head losses for each component in microfluidic system. (

**B**) Head losses and head pump required against flow rate.

**(C**) Point-to-point pressure drop and diodicity. (

**D**) Diodicity comparison of Tesla valve, Tesla Serpentine, and Tesla Serpentine-Hydrogel integration in a microfluidic device.

**Figure 7.**Comparison of friction coefficient (

**A**) Experimental, simulation and analytical value in this work (

**B**) Comparison of empirical correlation and experimental value (

**C**) Laminar flow error analysis.

**Figure 8.**(

**A**) Swelling process of one-unit hydrogel after applied single droplet water over time (

**B**) Simulated result of storage chamber design optimization by varying (

**B**) quantity of reservoir entrance port and (

**C**) diffuser width.

**Figure 9.**Schematic of the ABS printed mold-removal fabrication method for manufacturing microfluidic devices showed photo of cure PDMS with ABS printed mold and sacrificial template ABS printed mold removal create cavity microchannel.

**Figure 10.**Height and width comparison in various extrusion temperature, bed temperature, and print speed.

**Figure 11.**Flow rate for the operation of a finger actuated-based microfluidic system in (

**A**) 0.5 s, (

**B**) 5 s, (

**C**) 10 s and (

**D**) 20 s intervals.

Part | Parameter | Design | Prototype |
---|---|---|---|

Inlet | Inlet diameter, D_{i} | 0.25 mm | 2.5 mm |

Inlet height, h_{i} | 0.05 mm | 0.5 mm | |

Microchannel cross-sectional geometry | Width, w | 0.10 mm | 1.0 mm |

Height, h | 0.05 mm | 0.5 mm | |

Microchannel | Length, L(2→3) | 0.84 mm | 8.42 mm |

Length, L(3→4) | 2.8 mm | 28 mm | |

Length, L(4→5) | 0.3 mm | 3 mm | |

Tesla | Length, L(5→6) | 0.38 mm | 3.8 mm |

Pump chamber | Top diameter, D_{p(T)} | 0.5 mm | 5 mm |

Top height, H_{p(T)} | 0.3 mm | 3 mm | |

Bottom diameter, D_{p(B)} | 0.3 mm | 3 mm | |

Bottom height, H_{P(B)} | 0.05 mm | 0.5 mm | |

Diffuser | Width, a | 0.05 mm | 0.5 mm |

Width, b | 0.15 mm | 1.5 mm | |

Length, L(8) | 0.15 mm | 1.5 mm | |

Reservoir | Outer Diameter, D_{R1} | 1.6 mm | 16 mm |

Inner Diameter, D_{R2} | 0.8 mm | 8 mm | |

Air vent | Diameter, D_{o} | 0.1 mm | 1.0 mm |

First Author | Friction Coefficient Correlation | |
---|---|---|

Shah [125] | $\mathrm{f}=\frac{96}{\mathrm{Re}}\left(1-1.3553\mathsf{\beta}+1.9467{\mathsf{\beta}}^{2}-1.7012{\mathsf{\beta}}^{3}+{0.9545\mathsf{\beta}}^{4}-{0.2537\mathsf{\beta}}^{5}\right)$ | (30) |

Troniewski [124] | $\mathrm{f}=\frac{64}{{\mathrm{Re}}^{*}}$ for laminar $\mathrm{f}=\frac{0.3164}{{\mathrm{Re}}^{*0.25}}$ $\mathrm{w}\mathrm{h}\mathrm{e}\mathrm{r}\mathrm{e}{\mathrm{Re}}^{*}={\mathrm{Re}\left(2\mathsf{\beta}\right)}^{0.16}$ | (31) |

Kakac [126] | $\mathrm{f}=\frac{24}{\mathrm{Re}}\left(1-1.3553\mathsf{\beta}+1.9467{\mathsf{\beta}}^{2}-1.7012{\mathsf{\beta}}^{3}+{0.9545\mathsf{\beta}}^{4}-{0.2537\mathsf{\beta}}^{5}\right)$ For Smooth channel, aspect ratio $\mathsf{\beta}$ < 1 | (32) |

Spiga [127] | $\mathrm{f}=\frac{96}{\mathrm{Re}}\left(1-1.20233\mathsf{\beta}+{0.88119\mathsf{\beta}}^{2}+{0.88819\mathsf{\beta}}^{3}\phantom{\rule{0ex}{0ex}}-{1.69812\mathsf{\beta}}^{4}+{0.72366\mathsf{\beta}}^{5}\right)$ | (33) |

Hsieh [128] | $\mathrm{f}=\frac{48.1}{{\mathrm{Re}}^{0.94}}(\mathrm{F}\mathrm{o}\mathrm{r}\mathrm{R}\mathrm{e}240)$ | (34) |

Kandlikar [129] | $\mathrm{f}=\left(0.0929+\frac{1.01612}{\mathrm{L}/\mathrm{De}}\right){\mathrm{Re}}^{*\left(-0.268-\frac{0.3293}{\mathrm{L}/\mathrm{De}}\right)}$ ${\mathrm{where}\mathrm{Re}}^{*}=\mathrm{Re}\left(\frac{2}{3}+\frac{11}{24}\mathsf{\beta}(2-\mathsf{\beta}\right)$ | (35) |

Zhai [130] | $\mathrm{f}=\frac{24}{\mathrm{Re}}\left(1-1.3553\mathsf{\beta}+1.9467{\mathsf{\beta}}^{2}-1.7012{\mathsf{\beta}}^{3}+0.9564{\mathsf{\beta}}^{4}-0.2537{\mathsf{\beta}}^{5}\right)$ | (36) |

Song [96] | $\mathrm{f}=\frac{0.8227}{{\mathrm{Re}}^{0.42}};\mathrm{Re}\le 2700$ $\mathrm{f}=\frac{0.0352}{{\mathrm{Re}}^{0.0211}};2700\le \mathrm{Re}\le 4000$ $\mathrm{f}=\frac{0.07019}{{\mathrm{Re}}^{0.10498}};\mathrm{Re}\ge 4000$ | (37) |

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## Share and Cite

**MDPI and ACS Style**

Ahmad, N.N.; Ghazali, N.N.N.; Abdul Rani, A.T.; Othman, M.H.; Kee, C.C.; Jiwanti, P.K.; Rodríguez-Gómez, A.; Wong, Y.H.
Finger-Actuated Micropump of Constant Flow Rate without Backflow. *Micromachines* **2023**, *14*, 881.
https://doi.org/10.3390/mi14040881

**AMA Style**

Ahmad NN, Ghazali NNN, Abdul Rani AT, Othman MH, Kee CC, Jiwanti PK, Rodríguez-Gómez A, Wong YH.
Finger-Actuated Micropump of Constant Flow Rate without Backflow. *Micromachines*. 2023; 14(4):881.
https://doi.org/10.3390/mi14040881

**Chicago/Turabian Style**

Ahmad, NurFarrahain Nadia, Nik Nazri Nik Ghazali, Ahmad Taufiq Abdul Rani, Mohammad Hafiz Othman, Chia Ching Kee, Prastika Krisma Jiwanti, Arturo Rodríguez-Gómez, and Yew Hoong Wong.
2023. "Finger-Actuated Micropump of Constant Flow Rate without Backflow" *Micromachines* 14, no. 4: 881.
https://doi.org/10.3390/mi14040881