The mechanical behavior of the SiO2 microbridge and the electrical response of the piezoresistors are analyzed using the Finite Element Method (FEM). Subsequently, glucose solution is used to prove the functionality of the device and it is tested for the different glucose concentration levels ranging from 50 mg/dL to 450 mg/dL.
Diabetes and Glucose Sensing Devices
The sensitivity of the sensor is greatly affected if there is any change in the diffusion layer during the chemical process. The movement in the membrane provides the necessary mechanical stroke to the microsyringe, which delivers the required amount of insulin.
Motivation and Problem Statement
The geometric parameters of the system play an important role and determine the response time, sensitivity and linearity of the devices. The operation of the packaged device has been successfully demonstrated for various glucose concentrations ranging from 50 to 450 mg/dL relative to the reference solution.
Summary of Contributions and Results
- Design of a micro-bridge for Glucose Sensing Application
- Design of an Osmotic Pressure Sensor
- Fabrication and Packaging of an Osmotic Pressure Sensor
- Experimental Setup and Testing of an Osmotic Pressure Sen- sorsor
- Design of a Controlled Drug Delivery Pump
The next step is the boron diffusion to the front of the substrate for the piezoresistors. The sensitivity (output voltage per [mg/dL] per voltage per time) for the 25 µm device is about 0.5 µV/(V·mg/dL·min) and it is low compared to 10 µm device which is 2µV/(V ) ·mg/dL·min), without any amplification of the output voltage.
Organization of the Thesis
Introduction
The deflection in the microbridge can be measured by different methods such as optical, electrostatic, piezoelectric and piezoresistive. The micro-bridge used in this work differs from the existing literature in the following ways. The two piezoresistors are placed at the ends of the micro-bridge instead of the single piezoresistors along the micro-bridge [24].
Design of SiO 2 -based Piezoresistive micro-bridge
In addition, two more piezoresistors are placed at the anchoring area of the micro-bridge, where the voltage is equal to zero, as shown in the figure. Substituting these five constants in equation (2.5), we get the deflection of the micro-bridge. The relationship between the stress and strain of the microbridge is obtained by calculating the bending moment at different values along the x direction.
Simulation Results
The deformation of the microbridge increases from 0 to 0.787 µm when the force increases from 0 to 2 µN, as shown in Fig. The two piezoresistors are placed near the edges of the microbridge to measure the deformation as shown in Fig. 2.8(a) shows that the deformation increases as the thickness decreases from 2µm to 0.2µm when a force of 2µN is applied to the surface of the microbridge.
Summary
For the different values of the input voltages, the change in output voltage of the SiO2 microbridge is shown in figure. The change in output voltage of the SiO2 microbridge is from 0 to 1.472 mV, when a force value is varied from 0 to 2 µN. The output voltage of the SiO2 microbridge was from 0 to 1.472 mV, when a force value was varied from 0 to 2 µN, without any amplification of the output voltage.
Introduction
In this chapter, we have focused on the design and simulation aspects of an osmotic pressure sensor. The osmotic pressure sensor consists of a square cavity on the bottom side to fill the osmotic active substance. The simulation results of an osmotic pressure sensor are discussed in section 3.4 and finally the chapter is summarized in section.
Osmosis Principle
The basics of the osmosis principle and the design details of the osmotic pressure sensor are explained in Sections 3.2 and 3.3, respectively. Due to the difference in concentration, osmotic pressure develops and water molecules move from the region of lower concentration to the region of higher concentration as shown by the arrows. An ideal semipermeable membrane allows only water to pass through, but rejects large molecules (ions) of the solute.
Design of an Osmotic Pressure Sensor
Different Types of Sensing Techniques
Three different sensing techniques such as capacitive, piezoelectric and piezoresistive are studied to measure the displacement in the Si membrane of the osmotic pressure sensor. The top electrode is fixed and attached to the side walls of the device with studs. The change in the capacitance due to the displacement in the membrane is given by equation Here d0 is the initial gap between the plates and dz is the gap between the plates due to displacement.
Simulation Results
The Si membrane at the top of the device is solid and the square cavity at the bottom is liquid. The output voltage versus pressure for the different sizes of the Si membrane is shown in figure. The thickness of the Si membrane and the surface area of the semi-permeable membrane also affect the performance of the device.
Summary
Introduction
Finally, the device must be packaged to provide mechanical protection with the possibility of an interface between the MEMS device and the environment to produce the desired output. The device is packaged using the Polycarbonate (PC) material using a simpler technique instead of following the standard packaging procedure. The device performance is discussed in Section 4.4 and finally the chapter is summarized.
Fabrication of a Piezoresistive Pressure Sensor
S1813, positive photoresist (PPR) is coated on the front side of the wafer over SiO2 using a spin coater. During BSG etching, the thermally grown oxide on the backside was thinned and a cross-section is shown in the figure. After device fabrication, I/V characterization is performed using a DC probe (Agilent 4155C) to check the metal contacts and resistance values of the piezo resistors as shown in Fig.
Packaging of a Piezoresistive Pressure Sensor
Fabrication of a Testing Chamber
The test chamber is designed and fabricated according to the dimensions of the packaged device. The packaged device can be precisely fixed in the test chamber so that it can interact with the fluids. The packaged device is attached to the fluid test chamber to test the device for fluid applications as shown in Fig.
Performance Testing
A small hole is made in the square acrylic sheet, and it exactly matches the cavity of the device on the bottom side. A plastic tube is fixed in the hole with the same adhesive material as the passage for external pressure. The device is tested by applying external pressure and its value is monitored using a commercial pressure gauge which is already calibrated.
Summary
The temperature used to place the device in the cavity of the PC material was low there by the thermally induced stress was reduced. Additionally, PC was a soft and flexible material and was an excellent choice for packaging actuators, pressure sensors, and fluid interaction devices. PC appears to be a promising material for packaging prototype devices and does not require any specific equipment.
Introduction
The hydrogel is enclosed in a pressure sensor to facilitate measurement of changes in glucose concentration levels. In this chapter, we demonstrated an osmotic pressure sensor that uses the principle of osmosis to measure glucose concentration levels. The osmotic pressure sensor has been successfully demonstrated to measure different levels of glucose concentration ranging from 50 mg/dL to 450 mg/dL.
Experimental Setup for an Osmotic Pressure Sensor
Experimental Setup
To begin testing, the square well is filled with a standard glucose concentration of 100 mg/dL. To characterize the device's response time, the glucose concentration in the square cavity must be reset to the reference value of 100 mg/dL, that is, the output voltage must be brought back to the reference value close to zero. In a separate measurement, a glucose concentration of 100 mg/dL is placed in the test chamber for 720 minutes, and the output voltage obtained was constant for the entire duration.
Results and Discussion
The response time of an osmotic pressure sensor is characterized as the time required to reach 95% of the stable value of the output voltage for the change in glucose concentration in the test chamber (e.g. 150 mg/dl) [10] . The variation in the output voltage appears to always be proportional to the change in glucose concentration. The fabricated glucose sensor can measure a small variation in glucose concentration levels (results for 50 mg/dL are plotted, but can be as low as 20 mg/dL).
Summary
The micropump used in this work differs from others in the literature in the following ways. Moreover, the electrodes in the micropump are driven by a low amplitude AC voltage of low frequency f0 and thus suitable for low power portable devices. An additional advantage of using low voltage is the absence of electrolysis in the system.
Electro-osmosis Theory
The fluid velocity will be zero at both low and high frequencies [70], because most of the applied voltage falls across the double-layer and bulk electrolyte, respectively. The fluid velocity at the surface of the electrodes is formulated by Helmholtz-Smoluchowski. The relationship between the electroosmotic velocity and the tangential component of the applied electric field is given by the equation.
Design of an Electro-osmotic Pump
The flow rate of the micropump depends on the frequency and amplitude of the applied signal and on the conductivity of the electrolyte. The typical transition frequency, f0, is much smaller than the charge relaxation frequency of the bulk electrolyte. In general, the output voltage of the glucose sensor increases corresponding to the change of glucose concentration levels in the human body from a normal level.
Simulation Results
In an AC electroosmotic pump, the geometry of the device affects the overall pump performance in more ways than one. If the outlet is smaller than the micropump, its hydraulic resistance will be greater and part of the liquid will circulate back into the upper part of the micropump, where it encounters a smaller resistance. In addition, the electrodes lie on the same plane, and the shape of the applied electric field is circular [71] and the movement of the liquid directly above the surface of the electrode will follow such a shape.
Summary
The geometric effect is that the outlet height can induce flow recirculation in the pump channel when the outlet height is smaller than the micropump [69]. The fluid velocity moves in the opposite direction when the height of the micropump is above 140 µm. This electroosmotic pump has the potential to deliver insulin according to glucose concentration levels if integrated with a continuous glucose monitoring system.
Conclusion
The packaging cost was low due to the PC material, and it does not need steps such as laser cutting, lithography, etching or anodic bonding. A sensitivity of 1 µV to 2 µV/(V·mg/dL·min) was achieved for a 10 µm membrane thickness device compared to 0.3 µV to 0.5 µV/(V·mg/dL·min) for a 25 µm thick membrane, without any amplification of the output voltage. The response time of the designed devices was found to be less and the output voltage corresponding to the variation in glucose concentration was found to be proportional.
Future Work
The flow rate at the outlet of the electroosmotic pump increased from 32 to 107 µm/s when the input voltage was switched from set1 to set4. Two different devices can be fabricated on the same substrate, one being an osmotic pressure sensor and the other being a micro-actuator. Displacements generated in both devices are used for controlled drug delivery when a microfluidic pump is attached to the sensor's movable membrane.
List of Publications International Journals
Lin, “A mems affinity glucose sensor using a biocompatible glucose-responsive polymer,” Sensors and Actuators B: Chemical , vol. Ji, “Fabrication and characterization of SiO2 microcantilevers for microsensor application,” Sensors and Actuators B: Chemical , vol. Chang, “A piezoresistive bridge-microcantilever biosensor by {CMOS} process for measuring surface stress,” Sensors and Actuators B: Chemical , vol.
