## 1 Introduction  
As red blood cells (RBC) are the most common cells in blood, they have a strong influence on the mechanical properties of blood samples. Various mechanical properties of blood samples (i.e., blood viscoelasticity, RBC deformability, and RBC aggregation) have been used to monitor physical variations in blood samples. Blood viscosity is determined primarily by RBC aggregation at lower shear rates and by deformability at higher shear rates. Plasma proteins contribute to enhancing RBC aggregation. Owing to their unique features, such as high surface-volume (S-V) ratios, low cytoplasmic viscosities, and soft plasma membranes, RBCs can travel into micro-sized vessels that are smaller than their diameters. Hemoglobin, hydration, and ion pumps contribute to varying RBC flexibility and blood viscosity [1]. Altered membranes and cell functions in inherited RBC disorders exhibit decreases in RBC deformability. RBC deformability decreases due to altered RBC geometry, increased cytoplasmic viscosity, and altered membrane viscoelasticity [1]. Altered RBC mechanical properties are strongly correlated to various diseases, such as hereditary disorders, metabolic disorders, adenosine-triphosphate-induced membrane changes, oxidative stress, and paroxysmal nocturnal hemoglobinuria [2]. For these reasons, hemorheological properties have been stressed as potential indicators of specific pathological disorders. They are considered as label-free biomarkers for monitoring cell status and clinical relevance [2]. As blood samples flow dynamically through blood vessels (i.e., arteries, arterials, and capillaries), it is necessary to measure static property (i.e., viscosity or pressure) obtained under constant flow conditions and dynamic property (i.e., elasticity or compliance) obtained under transient flow conditions. Using blood delivery tools (i.e., syringe pumps, pressure sources), blood samples can be supplied into a microfluidic device at a constant flow rate. Blood viscosity can then be obtained by means of several techniques (i.e., coflowing streams [3; 4], reversal flow switching [5], advancing meniscus [6; 7; 8; 9; 10], electric impedance [11; 12], droplet length [13], and digital flow compartment [14]). Among these methods, the parallel flow method has been used widely due to simplicity and continuous monitoring. It does not require a calibration procedure with standard fluid. However, the method requires two syringe pump to supply the blood sample and reference fluid. As syringe pumps are bulky and very expensive, it is necessary to reduce the number of syringe pumps. Furthermore, the previous method does not provide information on blood elasticity. More recently, the author has suggested the viscoelasticity measurement technique in dynamic varying flow conditions. First, by controlling the blood flow rate in the oscillatory pattern, the viscosity and elasticity can be obtained periodically by analyzing the interface in coflowing streams (i.e., blood stream, 1\(\times\) phosphate-buffered saline (PBS) stream) [15]. Second, viscosity and elasticity can be obtained sequentially by monitoring reversal flow switching in a Wheatstone bridge circuit under a steady flow and transient flow. As a limitation, the previous methods require two pumps to supply two fluids. Instead of blood viscosity, blood pressure is suggested to monitor the mechanical properties of blood samples, where blood samples are delivered into a microfluidic device with a single syringe pump. [PERSON] et al. measured the liquid pressure inside the microfluidic channel by monitoring the air-liquid interface in the glass substrate [16]. To avoid air leakage throughout the polydimethylsiloxane (PDMS) polymer, the microfluidic device was fabricated with a glass substrate. [PERSON] et al. measured liquid pressure with the same principle suggested by [PERSON] et al. [17]. They used a glass capillary to monitor the liquid-air interface. [PERSON] et al.