ECE 5210

Design of “Smart” Bandage Using Paper-Based MEMS: Final ReportLisa Anders ECE, Vivek Jayabalan ME, Sai Ma SBES 5/11/2013

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IntroductionObjective

The goal is to create a "smart" bandage that would incorporate an inexpensive modular sensing platform for monitoring healing including temperature, pressure, attachment, and bandage viability using an active electronics design. The signal can then be sampled and processed by a microcontroller and wirelessly transmitted to the doctor's office or the hospital to provide real-time monitoring of wound recovery. Paper-microelectromechanical systems (MEMS) technology is used to produce an inexpensive end product. The bandage and sensor components are disposable while the remaining electronics are reusable.

Figure 1: Schematic showing the current and proposed aims of the project.

Team Qualifications/Research Team (0.5-1.0)Timeline and task assignment

Our team is composed of Lisa Anders, a first year EE graduate student, Sai Ma, a second year BMES graduate student, and Vivek Jayabalan, a first year ME graduate student.

Lisa Anders’ research focuses on Lab-on-a-Chip rare cell isolation. In particular her current focus is on the manipulation and characterization of tumor initiating cancer cells using dielectrophoresis. She is very interested in using sensing technology to improve medical conditions on a patient level. For her undergraduate senior design project she built an open-source electroencephalography (EEG) machine.

Sai Ma’s research focuses on applications of microfluidic chips for biological assays, such as on chip PCR, electroporation for curing infection, and epigenetic analysis. He is interested in developing new tools which benefits biological and medical research. He also has an analytical chemistry background and studied paper-based nanoparticle labeling detection technology for protein and bacteria for his undergraduate thesis.

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Vivek’s research focuses on investigating the thermal and electrical conductivities of cancerous and noncancerous tissues. Being a mechanical engineer he wishes to explore the relationship between bandage pressure and the wound healing process. For his undergraduate thesis, Vivek worked on developing meso-scale piezo-actuated pumps.

Table 1: Diagram showing responsibilities and timeline.

Sensor Responsibility Timeline (weeks)

Modeling Fabrication Integration

Temperature Sai 1-2 3-8

Pressure Vivek 1-2 3-8

Moisture Lisa 1-2 3-8

Attachment Lisa 1-2 3-8

Signal Processing Lisa 4-10

Presentation N/A Lisa, Sai, Vivek 10-11

Final Report N/A Lisa, Sai, Vivek 10-11

Significance and backgroundThere were 4.5 healthcare associated infections for every 100 hospital in-patient admissions in 2007, a total of 1,737,125 cases of infection. 290,485 of those cases were due to surgical site infections. Using the consumer price index it has been shown that these infections cost the US medical system $35.7 to 45.0 billion annually. $25.0 to 31.5 billion would be saved if 70% of those infections could be prevented. In addition these infections cost individuals personal time, work time, increase in short term and long term morbidity, and psychological costs. (CDC stats)

Focusing on inexpensive prevention and early detection of infections using better monitoring systems would lead to a more scalable and affordable solution to healthcare. (Milenkovic - wireless sensor networks for personal health monitoring)

The smart bandage could be used either to monitor a post- operative care patient or a patient with a deep tissue injury. Such a modular sensing platform would allow for real-time monitoring and prevention of infections or complications during the healing process. It could also notify when the bandage needs to be changed, allowing for more efficient use of hospital resources.

A paper is an attractive substrate choice for MEMS devices for three main reasons: 1. low cost and wide availability, 2. the ability to “wick” fluids allowing for passive transport of fluids, and 3.knowledge from analytical chemistry’s use of paper as a platform for decades. In addition to those reasons, paper can be acquired in a range of thicknesses, is easy to stack, store, and transport, being made from cellulose it is compatible with biological samples, it is flexible allowing for diverse geometries, and it is flammable allowing for safe and easy disposal without specialized equipment. Early paper-based MEMS include such popular diagnostic tools as the

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litmus test and urine analysis dipsticks. (Martinez, A.W., et al., Diagnostics for the developing world: microfluidic paper-based analytical devices. Anal Chem, 2010. 82(1): p. 3-10.)

The combination of multiple sensors on a single paper-MEMS platform that we have proposed has not previously been applied to bandage technology. The signal-processing and power electronics, and the smart bandage will be built into two separate modules. The bandage/paper module will be disposable while the remaining electronics will be reusable. This will allow for the bandage to be removed every 2-3 days as necessary at minimal cost while retaining the more expensive electronics.

Similar to our project is the field of body sensing. The use of personal sensing systems such as stress monitoring (Jovanov, stress monitoring using a distributed wireless intelligent sensor system), electrocardiography (ECG) and foot switch data acquisition (Milenkovic - wireless sensor networks for personal health monitoring), and even a sensor to detect sleep apnea using a blood oximeter (wireless body sensor networks for health-monitoring applications, Hao) have been demonstrated. Other smart sensing systems include everything from portable ECG’s and electroencephalography (EEG’s) (http://www.gla.ac.uk/media/media_136620_en.pdf) to fitness trackers like fitbit (http://www.fitbit.com/).

Basis behind each sensorTemperature It is known that local temperature increase is one of the main symptoms of skin infection. Temperature difference between periwound skin and an equivalent contralateral control site was found to be less than 2°C without an infection presentation. If an infection is present the difference has been shown to be greater than 2°C. For example, on average, skin temperature at the hottest spot on an infected limb was 34.4 degrees C, compared with 30.9 on an uninfected limb. [http://www.ncbi.nlm.nih.gov/pubmed/20631603] A temperature sensor would help improve the treatment of skin and soft tissue infections.

PressureTactile sensing is a useful healthcare tool while handling sensitive wounds, especially amongst geriatrics. A simple way of realizing tactile sensors is by fabricating paper-based diaphragm pressure sensors. The pressure sensor in the bandage could subsequently be employed in negative pressure therapy, for severe burn wounds, monitoring patient compliance for physical rehabilitation, and in compression therapy, for venous leg ulcers. Although, there has been sufficient work on incorporating pressure sensors for compression therapy [1], little work has been done in improving these sensors for tactile purposes on bandages. In particular, sensitive pressure sensors could become very useful for wound recovery for diabetic patient [2].[1] Casey, Vincent, et al. "Wearable sub-bandage pressure measurement system." Sensors Applications Symposium (SAS), 2010 IEEE. IEEE, 2010.[2] http://www.webmd.com/a-to-z-guides/wound-care-10/diabetic-wounds?page=2

AttachmentThe skin impedance depends on the skin type and thickness, skin hydration, and electrode geometry. For frequencies below 10kHz the impedance is dominated by the stratum corneum while above 10kHz it is dominated by viable skin. This means that skin resistance is nonlinear and even varies day to day in the same patient. Even similar skin locations’ resistances can vary greatly from 250 kohm square on the forearm-ventral-distal to 840 kohm square on the forearm-ventral-middle and 560 kohm square on the forearm-ventral-proximal. (XBioimpedance and Bioelectricity)

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Figure 2: shows expected conductivities at 10kHz from the literature at different skin sites. (XBioimpedance and Bioelectricity).

This nonlinearity implies that for measuring attachment a switch approach would be more effective than an analog signal.

Moisture SensorPaper, as a permeable substrate, will form an electrical connection when doped with a conductive fluid. Using known values for the conductivity of expected fluids we can predict the resistance between the two electrodes.

Table 2: Resistances of different liquid is predicted for 9cm gap. DI water, Tap water and PBS are tested. Data is shown in the following section.

Material Conductivity Resistance Predicted resistance for 9cm gap

DI water 0.1 100,000 ohm/m

70,000 ohm

Tap water 500-800 20 ohm/m 1.4 ohm

PBS 1.4 0.7142 ohm/m 0.05 ohm

Blood 0.7 1.428 ohm/m 0.49ohm

Preliminary Results

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Individual Sensor Calibration/Generation 1

Generation 1 of the device was split into 4 sensors; temperature, attachment, moisture, and pressure. The preliminary design for each sensor is demonstrated and discussed below.

Temperature Sensor

The temperature sensor is a T type thermocouple which consists of copper and constantan wires. The thermocouple has an ideal measurement range from -200 to 300°C and a sensitivity of roughly 43µV/°C.

Figure 3; a) Generation 1 temperature sensor design. A T type fine gauge thermocouple is embedded in a pre cut filter paper. A small piece of tape is used to attach of the tip of thermocouple to the filter paper. b) Experimental setup. A plate thermal cycler is used to the calibrate temperature measurement. The voltage change due to the temperature change of thermocouple is measured by a multimeter.

To calibrate the thermocouple it was tested under physiologically relevant conditions. It had a very linear measurement range as shown in figure 3.

Figure 4: a) The linear response of the voltage signal produced by thermocouple to temperature is measured by multimeter. b) Specific body temperature range (35 to 40°C) is measured.

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PressureMost pressure sensors that measure greater-than-atmospheric pressure utilize deformable diaphragms [2] due to straightforward fabrication and a simple working mechanism. Thus, a diaphragm based pressure sensor using paper, aluminum foil, and PDMS was developed. A relative pressure differential causes a deflection in the diaphragm. That change in geometry changes the effective capacitance of the sensor as shown in figure 4.

Figure 5: Schematic demonstrating the working principle of the diaphragm based pressure sensor.

A simple mathematical model can be constructed to understand the change in capacitance when a differential pressure is applied on the diaphragm. The model considers the circular diaphragm to behave as an ideal Kirchoff-Love plate [1], with clamped boundary conditions.

w z (r )= P a4

64D [1−( ra )2]2

D= Eh3

12 (1−ν2)where, w z (r ) is the deflection at a radius r of the plate along the ‘z’ direction, while assuming the plate to sit on the ‘x-y’ plane. P is the applied pressure, a is the radius of the plate, and D is the flexural rigidity of the plate as demonstrated in the subsequent equation.

The effective capacitance of this bent plate pressure sensor can be modeled as

c= ϵAd−w z

'

where w z' is the average displacement of the diaphragm. The net change in the capacitance is

determined by the following non-linear equation

Δ c (P )=ϵA 815

Pa4

64 D

d− 815Pa4

64D

By making the assumption that the net deflection of the plate is much less than the gap between the plates the equation above can be linearized to the following

Δ c (P )≅ϵA 815Pa4

64Dd

=C1 P

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where the constant C1 is a function of the geometric and material properties of the designed pressure sensor.

Generation 1 of the diaphragm-based pressure sensor was achieved by sandwiching conductive layers between patterned layers of the thick papers. While this method is useful when making capacitors with small openings where the conductive layers can’t come in contact but are difficult to calibrate. Making sensors with larger openings risks the two conductive layers deflecting enough to short. Th is was improved by incorporating a thin insulating layer of PDMS to preventing shorting. This layer also acts as an additional nonlinear spring in the sensor.

The following schematics show the fabrication of the capacitive pressure sensor.

(a) (d)

(b) (e)

(c)

Figure 6: Fabrication of the paper-based diaphragm pressure sensor. 1a) Thick paper is (b) patterned by either cutting or spot heating to burn the region. Patterned paper alternates a (c) solid conductive layer such as aluminum foil. (d) A thin layer of PDMS is deposited on the composite. (e) Another aluminum-paper composite shown sandwiches (c) the previously fabricated structure to make the pressure sensor.

The devices were fabricated as demonstrated in figure 6 and were calibrated. The fabricated pressure sensor had an initial capacitance of 5.8pF. When the structure was bent, which applied a force on the diaphragm, the capacitance changed to 8.5 pF and when this force was released, the capacitance returned to around 6pF. The following section discusses the calibration of a pressure sensor with a 4cm square opening.

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Figure 7: Changes in capacitance of the pressure sensor due to a change in geometry. The capacitance changes from 5.8pF in the right to 8.5 pF in the left on application of external pressure, by bending the device.

The aforementioned capacitive pressure sensors were calibrated using the TENMA 72-1025 Impedance Meter. A pressure sensor, with an opening of roughly 4cm square was used in calibration, and the sensor had a PDMS insulating layer between the conductive aluminum pads. Figure 8 demonstrates the experimental setup, where a 10mL beaker with varying amounts of bolts acted as the weight load.

Figure 8: Image of the experimental setup used in measuring the capacitance of the pressure sensors. The beaker with bolts acts as the weight load on the sensor.

Several experiments were conducted that showed variability in experimental results. That variability is due to the plasticity [3] of the thin PDMS layer.

sensor hysteresis. This non-linearity can be attributed to large deflections relative to the designed gap and the hysteresis occurs due to plasticity [3] of the thin PDMS layer. The change in the capacitance with the weight is projected in the figure ??

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Figure 9: Change in capacitance with increasing weight.

Attachment Sensor

While initially the assumption was that short distances between the attachment sensors would allow for minimal resistance measurements the preliminary results show this to not be the case. The relationship between distance and skin resistance is not correlated, suggesting that resistance depends more on skin region and skin factors rather than distance between electrodes. Therefore a spacing of 9cm was chosen in order to minimize the total number of electrodes needed to measure the peripheral edges of the bandage.

0.5 1 1.5 2 2.5 3 3.505

101520253035404550

Across pinkie finger MohmAcross pointer finger MohmAcross arm Mohm

Distance between electrodes (cm)

Resis

tanc

e (M

ohm

s)

Figure 10: Preliminary attachment sensor results from different skin regions by varying the distance between the electrodes. The resistance is more closely related to skin region than electrode distance.

Moisture Sensor

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Figure 11: (a) Picture of the Generation 1 moisture sensors. (b) Moisture sensor preliminary results from doping with different conductive fluids.

By testing several commonly found conductive fluids the operation of the moisture sensor was calibrated. The found resistance is about three orders of magnitude higher than the predicted resistance values from table 2. This is due to the non-conductive paper substrate which decreased the overall conductivity.

Incorporation/Generation 2

After each sensor had been individually calibrated, the temperature and attachment sensors were incorporated as Generation 2. The initial designs for the developed pressure sensor required stiffer papers and therefore was separately calibrated.

Figure 12: Incorporation of the temperature and attachment sensors resulted in Generation 2 of the smart bandage.

Incorporated Temperature Measurement

To avoid the direct contact of the sensor metals to the skin a single layer of gauze was positioned between temperature sensor and skin. In figure 6 the measured temperature with a gauze layer showed slightly lower values than the sensor without the gauze layer which results from the low heat conductivity of the gauze layer.

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Lower arm

Middle arm

Upper arm

2526272829303132333435

Tem

pera

ture

/ °C

Lower arm

Middle arm

Upper arm

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28

30

32

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LisaSaiVivek

Tem

pera

ture

/ °C

Figure 13: a) Temperature measurement from the thermocouple in direct contact with the skin. b) Temperature measurement with a thin gauze layer separating the thermocouple from direct skin contact. The temperature values measured with the inserted gauze layer were a bit lower than the sensor without the gauze layer. Calibration is needed for precise measurement, however this shows proof of concept that the sensor is still sensitive enough to measure the skin temperature even through a gauze layer.

Incorporated Attachment Measurement

Skin impedance was found to vary greatly both by day, position, and individual. This suggests that instead of treating the skin impedance as an analog sensor it would be better to treat it as a digital sensor with thresholds for unattached, attached, and shorted. The range of our experimental results suggests thresholds set at 0.07Mohms and 90Mohms.

Figure 14: Attachment sensor results from 2 days, 3 individuals, and 6 skin locations. High variability in individual measurements from day to day and location to location suggests the use of a threshold approach.

Proposed Research

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The scope of our project thus far was to develop a proof of concept prototype of the smart bandage sensors. Our further work involves moving toward a commercial end product. Specifically we propose to

Aim 1: Establish a bulk fabrication protocol and incorporate additional sensors.

The preliminary objective of this project is to build the entire smart bandage economically. Moving towards a bulk fabrication technique is therefore the easiest way to reduce overall cost.

We propose to investigate two approaches to fabrication. First incorporate, during the fabrication process, all four sensors on the same paper substrate. The second approach would be to incorporate the attachment, moisture, and temperature sensors during fabrication whereas the pressure sensor would be incorporated afterwards as an additional module. The pressure sensor requires a substrate with different and specific material properties. It may be easier and more cost effective to fabricate separately.

Fabricating the sensors on a paper substrate can be done by employing the screen printing techniques proposed by the Dr. Whitesides group at Harvard in [Liu, Xinyu, et al. "Paper-based piezoresistive MEMS sensors." Lab on a Chip11.13 (2011): 2189-2196.]. The metals can be deposited using evaporation, spray deposition, or sputtering where a stencil is used for patterned deposition.

(a)

(b)

(c)

(d)

(e)

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(f) (g)Figure 15: Schematic for the first proposed bulk manufacturing process.

Aim 2: Establish feasibility of wireless integration.

There are several approaches to integrating a microcontroller with the bandage. Our approach is to use a microcontroller for sampling, signal processing, and encoding with an attached wireless module for transmission. Another approach that was briefly investigated would be to have short-range data transmission to a body-worn hub capable of longer-range transmission to reduce the size of the transmission device at the bandage site.

The MSP430 Launchpad (TI, Texas), is a suitable microcontroller choice due to its low weight, customizability, and ultra-low power modes. A low power is the most important characteristic as the bulk weight of the system typically is dominated by the size and weight of the batteries (Milenkovic - wireless sensor networks for personal health monitoring). As an ultra-low power microcontroller the MSP430 would allow the entire system to be run off of a single power source, for example a commercially available 150-300mAh lithium coin battery with a weight of less than 5 grams. In addition it is known for its integrated emulator interface allowing for real-time in-system programming and debugging while connected to a PC. It is also RoHS compliant meaning that it has been tested for toxic compounds, an important certification for biomedical devices.

Figure 16: Block diagram of the MSP430 architecture from from http://lp-hp.com/files/2009/09/cheryl2.jpg.

The ANT wireless chip and protocol (TI, Texas), has been selected for use with the MSP430 Launchpad microcontroller due to ease of integration in addition to its minimal size and weight.

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There are several programming routines that could be implemented. The microcontroller could be programmed to signal regularly for continual monitoring or as an alert system where a signal would be transmitted only if the sensors seem to indicate an abnormal condition. Both routines could have advantages depending upon the type of monitoring and availability of resources.

The largest foreseen challenge is security and privacy requirements. To prevent eavesdropping, identity spoofing, or redirection of private data to unauthorized individuals all wireless medical sensors are required by law to meet certain privacy requirements. Encoding and keyshaking

Aim 3: Move into long term monitoring and establish an animal model.

The final aim of the integrated smart bandage is to monitor the wound healing process within several days. Patients should be little or no feeling of the exist of the sensors which requires a robust and low weight package of the entire system. Even though the bandage part of our design is replaceable, each bandage with incorporated sensors still need to be on skin for about three days to reduce labor of medical care. Before long time monitoring experiments of human subjects, comfort tests and animal tests need to be done.

Nude mice is a widely study animal model. The exposed skin is an ideal experimental subject for temperature and attachment sensing test. The experimental group with infected wound can be compared with control group which is treated by antibiotics to avoid unwanted infection.

The major challenge would be telling the difference between infected wound and uninfected wound. Temperature sensor can be placed on periwound skin to monitor infection process. The sensitivity and threshold of detection need to be decided depending on experimental data. Once the threshold of whether the wound is infected or not is determined, the sensor would be ready for judging real wound healing process.

The place to position attachment sensor also needs to be investigated since the mice are moving all the time. The position should has no influence on the movement of mice and also need to be tight enough to judge the attachment of the bandage. The position information gathered on the mice would be the best guide for human smart bandage design

Broader Impacts (0.5)

Our knowledge of smart sensing systems and our success with the smart bandage indicates that truly integrated smart bandage technology is feasible.

Since the sensing platform should be considered disposable along with the bandage, paper-based MEMS would be incorporated to build a cost effective system. Using paper as a substrate and established paper-MEMS fabrication techniques will allow for ease of construction, low cost, and bandage flexibility. Future implementations to reduce cost would be to screen print the electrical contact pads and the thermocouple wires. Due to its low power capabilities the signal processing microcontroller will be the TI MSP430 ‘Launchpad’ microcontroller.

Impact on Medical Community/SocietyAffordable healthcare is imperative in the coming years. Health care expenditures reached $1.8 trillion in 2004 and is predicted to reach almost 20% of the US GDP by 2016. Inexpensive, non-invasive, continuous, real-time health monitoring promise to revolutionize health care. The key

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is managing and monitoring health problems before they become a problem. (Milenkovic - wireless sensor networks for personal health monitoring)

Educational ImpactMEMS is a rapidly growing field, with a 37% market expansion rate in 2011 to $2.25 billion just in consumer electronics and cell phones (http://www.eetimes.com/electronics-news/4219599/IHS-projects-record-growth-for-consumer-MEMS), and developing engineers for careers in MEMS fields needs to a priority. Despite this growth the cost of building MEMS devices is prohibitive to the undergraduate classroom and few engineers are exposed to MEMS technology before graduate education. Paper-MEMS projects such as the one developed in this paper are inexpensive and relatively simple and could be used in an undergraduate or even high school lab or classroom. This could provide the bridge for young engineers looking to MEMS for future careers.

References

[1] Young, Warren C., and Richard G. Budynas. Roark's formulas for stress and strain. Vol. 6. New York: McGraw-Hill, 2002.[2] Eaton, William P., and James H. Smith. "Micromachined pressure sensors: review and recent developments." Smart Materials and Structures 6.5 (1997): 530.[3] Sawano, Satoshi, et al. "Sealing method of PDMS as elastic material for MEMS." Micro Electro Mechanical Systems, 2008. MEMS 2008. IEEE 21st International Conference on. IEEE, 2008.

Let’s use endnote!

Outline: (NSF)● 1.0-1.5 pages Introduction/Objective

● 0.5-1.0 pages Team Qualifications/Research Team

● 3-4 pages Significance and Background (why critical what is the point, summary of other peoples work, can grab other peoples figures)

● 3 pages Preliminary Results

● 5-6 pages Proposed Research

● 0.5 pages Broader Impacts

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