Epidermal Differential Impedance Sensor for Conformal Skin Hydration Monitoring
© The Author(s) 2012
Received: 22 June 2012
Accepted: 2 August 2012
Published: 23 August 2012
We present the design and use of an ultrathin, stretchable sensor system capable of conformal lamination onto the skin, for precision measurement and spatial mapping of levels of hydration. This device, which we refer to as a class of ‘epidermal electronics’ due to its ‘skin-like’ construction and mode of intimate integration with the body, contains miniaturized arrays of impedance-measurement electrodes arranged in a differential configuration to compensate for common-mode disturbances. Experimental results obtained with different frequencies and sensor geometries demonstrate excellent precision and accuracy, as benchmarked against conventional, commercial devices. The reversible, non-invasive soft contact of this device with the skin makes its operation appealing for applications ranging from skin care, to athletic monitoring to health/wellness assessment.
Skin hydration monitoring is a well-established and important technique in dermatology, for analyzing various diseases and determining the effectiveness of medical therapies [1–4]. Hydration measurement is also useful in cosmetology, for assessing the function of anti-aging and moisturization treatments [5–8]. Skin hydration levels are typically characterized by measurements of skin electrical impedance [9–12] or thermal conductivity , or by optical spectroscopic techniques [9, 14, 15] including reflectivity . Indirect methods include evaluation of mechanical properties of the skin [17–19] or its surface geometry [20, 21]. Among these methods, electrical impedance provides the most reliable and established assessment, due to its instrumental simplicity and minimized cost. Several commercial systems that perform this measurement are available [22–25]. Such technologies generally rely on physical contact between the soft, curved surface of the skin, and rigid, planar electrodes . Here, the accuracy and repeatability both depend critically on the contact force between the electrodes and the skin. As a result, most devices incorporate pressure-sensing components that help the user to apply, with precision, the appropriate force during measurement. Such devices are not well suited for spatial mapping, for evaluating sensitive areas of the skin, or for continuous monitoring, thereby greatly limiting their utility .
Here we present a different approach, based on extensions of previously reported concepts in ‘epidermal’ electronics, in which semiconductor and related devices mount on ultrathin, elastomeric sheets in open mesh geometries, and integrate with the skin in a soft, van der Waals process, without any external application of pressure. This process achieves intimate biotic/abiotic contact in a way that does not mechanically load or constrain the natural motions of the skin . When incorporating devices with multiple functions, these epidermal systems enable non-invasive detection of physiological parameters such as electrical signals associated with activity in the brain, heart and skeletal muscles. In the following, we describe ideas and demonstration examples for an epidermal class of differential hydration sensor based on electrical impedance detection. The sensors effectively eliminate requirements on precise, external control of contact force, in a mode of integration that is mechanically ‘invisible’ to the wearer. The measurement outcomes can be interpreted using conventional approaches for impedance evaluation [28, 29]. The experimental results demonstrate, in fact, precision and accuracy that compare favorably to commercial moisture meters (CMM). Differential sensing schemes can be incorporated naturally, to compensate for variations in temperature, stresses/strains associated with human activity, and other disturbances that may lead to unexpected impedance changes. This capability is particularly valuable for applications in continuous monitoring. Arrays of miniaturized sensors can be achieved, for spatial mapping and, potentially, depth profiling of hydration. These collective attributes suggest that this type of technology, particularly when combined with other classes of sensors in a single system, could be valuable for wide-ranging uses in human healthcare and wellness evaluation.
2 Principle and Sensor Design
The epidermal hydration measurements reported here use impedance measurements performed directly on the skin, to exploit known correlations between electrical parameters of biological tissues and their water content . In particular, both the conductivity and permittivity of human skin change with the skin hydration levels [25, 30]. As a result, the hydration levels can be determined through impedance measurements, through appropriate calibration.
2.1 Epidermal Skin Hydration Sensor with Differential Detection
The fabrication of this device involves first spin-coating a sacrificial layer of polymethylmethacrylate (PMMA; 500 nm thick) and a layer of polymide (PI; 1 μm thick) on a silicon (Si) wafer substrate. Photolithographically patterned layers of Cr (5 nm) and Au (400 nm) form the electrodes as well as the serpentine interconnections, both of which include a top coating of PI (1 μm). Reactive ion etching (RIE) through selected regions of the PI defines vias to allow subsequent metallization to contact the serpentine interconnect traces at well-defined points. An additional patterned bilayer of Cr/Au (5 nm/200 nm) forms the floating ground plates. A final, uniform coating of PI (1 μm) encapsulates the system. Patterned RIE through the entire multilayer stack creates openings to allow contact with the skin and the open mesh structures needed to achieve desired levels of mechanical deformability. Removing the underlying PMMA by immersion in acetone for 5 min at 100 °C, allows the mesh to be removed from the Si wafer by use of water-soluble poly(vinylalcohol) (PVA) tape (Wave Solder Tape 5414, 3M Co.). Layers of Ti/SiO2 (5/40 nm) evaporated onto the backside of the released structure facilitate bonding to a thin, low-modulus silicone substrate (Solaris, Smooth-On, Inc.), first activated by treatment with ultraviolet induced ozone (10 min) to create reactive hydroxyl functionality on the surface. Dissolving the PVA tape in water for 20 min completes the fabrication, resulting in a mechanical construction that affords excellent stretchability (Fig. 1c) and flexibility (Fig. 1d), with a low effective modulus, which are all requirements for epidermal integration.
2.2 Conformal Skin Contact of the Hydration Sensor
3 Experimental Setup and Methods
3.1 Epidermal Differential Hydration Sensor
The sensors exploit three different designs. The circular electrodes involve a concentric design (Fig. 1b), consisting of inner disks (150 or 125 μm in radius) surrounding by open-ended rings (inner radius: 200 μm, outer radius: 300 μm). The interdigitated electrodes incorporate 8 (upper digit) or 7 (bottom digit) fingers each with dimensions of 20 × 50 μm2 and spacings of 20 μm. The meander electrodes involve two sets of four semicircular lines (radii: 30, 90, 150, and 210 μm, width: 20 μm) and two 60° and 150° arc lines (radii: 310 and 270 μm, width: 20 μm) interconnected into spiral shapes. In each of these three cases, the electrodes connect to corresponding bonding pads (width: 500 μm, spacing: 500 μm) and a common ground pad through serpentine traces (width 50 μm; ~225 μm radii of curvature). The floating ground plates consists of 600 × 600 μm2 pads situated above the reference electrodes and connected with serpentine traces through vias in the PI (230 × 230 μm2 in dimension). The use of different electrode designs in a single device allows us to assess various hydration sensing methods and sensor parameters under the same experimental conditions. The meander electrodes perform as resistors, which can be used for assessment of skin hydration through purely resistivity measurement. The other electrodes measure both resistance and capacitance changes. The circular electrodes offer symmetries that facilitate theoretical analysis; these geometries are also found in macroscale electrodes used in CMM [22, 23], to enable direct comparison.
3.2 Releasable, Stretchable Connector
3.3 Multiplexing Impedance Analyzer for Hydration Measurement
An impedance analyzer circuit equipped with multiplexing measurement capability serves as a system for quantifying the response of the sensors (Fig. 3e). Here, an impedance analyzer chip (AD5933, Analog Devices) with a 12-bit resolution provides an alternating current (AC) voltage (Vin) (2 V peak to peak) at frequencies between 1 and 100 kHz to each channel of the hydration sensor through a multiplexer (ADG 708, Analog Devices). Changing the combination of voltages supplied to channel selecting pins in the multiplexer through an I/O controller (USB-8451, National Instruments) and a computer allows each individual sensor channel to be probed at a given value of Vin. The amplitude and phase of the output voltage (Vout) measured from the common ground changes with the electrical properties of the skin. The values of Vout from all sensor channels can be obtained in a time-sequence, and converted back to impedances within the impedance analyzer.
3.4 Experimental Methods and Materials
Experimental evaluations involve attaching (and removing) epidermal hydration sensor systems to the skin of the ventral forearm to reveal frequency dependent changes in impedance associated with application of moisturizing lotions (Intensive Rescue Moisture Body Lotion, Vaseline Inc.). Hydration measurements performed at frequencies where changes in impedance are maximized provide optimal sensitivity. A CMM (MoistureMeterSC Compact, Delfin Inc)  that is based on impedance measurement enables calibrated conversion of impedance measurements to hydration levels.
4 Results and Discussion
4.1 Frequency-Dependent Impedance Changes with Hydration Levels
Figure 4c, d show changes in impedance amplitude and phase with skin hydration levels at selected frequencies. As the hydration increases, the skin conductivity and permittivity both increase, leading to a decrease in the real part (resistive part) and the imaginary part (capacitive part) of the complex impedance. The magnitudes of these changes diminish with increasing frequencies: for the range studied, the impedance amplitude and phase change by 89 and 73 %, respectively, at 15 kHz and by 34 and 38 %, respectively, at 95 kHz. The highest sensitivity is, therefore, obtained at low frequencies in this RF range.
Similar impedance responses to hydration changes (as Fig. 4a, b) appear in other measuring channels (channel 1–3). As the hydration changes from 40 to 110, the impedances (at 15 kHz) of the meander electrode in channel 1 and the interdigitated electrode in channel 2 change from 255 to 24 Ω and from 2.54 to 2.52 MΩ, respectively. The interdigitated electrodes in channel 2 provides a useful design for electrodes [23, 24, 37], but the complexity in the associated distribution of electrical field lines leads to complications in theoretical analysis . The symmetric sensor geometry of channel 3 and 4 enables direct comparison to outputs from the CMM, which has similar electrode construction, and the symmetry facilitates analytical modeling. For these reasons, the experimental results presented in the following focus on circular electrodes. In all cases, the impedance responses of the reference electrodes (channel 5–8) smaller than those of the measuring electrodes. In particular, the impedance changes (at 15 kHz) of the reference channels (from channel 5 to 8) are only 65, 31, 62, and 75 % of the changes in the corresponding measuring channels, and are likely due to electrical parasitic effects of the serpentine lines and passivated electrodes.
4.2 Comparison Between Hydration Measured by the CMM and Impedance Measured by the Epidermal Hydration Sensor
4.3 Repeatability in Measurements of Hydration Using Epidermal Devices and CMM
4.4 Spatial Mapping of Hydration Levels
The results reported here illustrate a soft, ‘skin-like’ technology, capable of intimately and non-invasively integrating with the skin for the purpose of quantitative assessment of hydration levels. The accuracy and repeatability of the measurement are both comparable to conventional devices that rely on rigid electrodes and controlled pressures during contact with the skin. The ability to perform internally referenced detection, in a differential mode, and the straightforward scaling to arrays of sensors for spatial mapping applications represent two additional, valuable features. Development of silicone substrates with enhanced ability to accommodate skin transpiration, and incorporation of wireless data transmission capabilities represent important directions for future work.
This work was carried out at the Fredrick Seitz Materials Research Laboratory Central Faciliteis, University of Illinois, which are partially supported by the US Department of Energy under grant DE-FG02-07ER46453 ad DE-FG02-07ER46471.
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