SAME COPYRIGHT NOTICE

[0001] A portion of the disclosure of this patent document contains material, which is subject to copyright protection. The copyright owner has no objection to the facsimile reproduction by anyone of the patent document or the patent disclosure, as it appears in the Patent and Trademark Office patent files or records, but otherwise reserves all copyright rights whatsoever,

GOVERNMENT SUPPORT

[0002] None.

RELATED APPLICATIONS

[0003] The present application relates to and claims priority to U.S. Provisional Patent

Application Serial No. 61/680500 entitled "A POLYMER-BASED MICROFLUIDIC

RESISTIVE SENSOR FOR DETECTING DISTR IBUTED LOADS, METHODS, AND PROCESSES FOR FABRICATING THE SAME" having a priority date of August 7, 2012.

FIELD OF THE INVENTION

[0004] The present invention relates generally to the field of small scale detection of distributed loads, both static and dynamic.

BACKGROUND

[0005] Detecting distributed loads at the micron-millimeter scale is needed for studying biological materials in biomedical industry, examining viscoelastic materials in various manufacturing industries, and feeling the texture of an object in robotics industry. However, there are currently no reliable, affordable commercial miniature products to meet this need. The most commonly adopted technical approach to detect distributed loads at the micron-millimeter scale is to develop sensor arrays, where each sensor has its own mechanical element to respond to a point load and its own transducer to convert the point load to an electrical/optical signal, and all the sensors are arranged into an array configuration to detect distributed loads across the array. The approach of sensor arrays renders the developed products to be unreliable, too costly, and take a long time to develop.

[00Θ6] With significant advances in biomedieal/surgical, manufacturing, and robotics technologies, there is a need to incorporate the function of detecting spatially distributed static and dynamic loads into their associated systems. Conventional distributed-load sensing products are commercially available, however these lack in size, spatial resolution, and force range/resolution for systems needed for detection of spatially distributed static and dynamic bads in, for example, minimally invasive surgeries. For instance, Teksean (South Boston, MA) offers FlexiForce® Matrix -sensors based on pressure-sensitive ink or rubber, and

Pressure Profile Systems (Los Angeles, CA) sells TaciArray™ sensors based on eapacitive transduction, and Novel (Munich, Germany) produces Pliance® sensor array, in fact, some of these distributed-load sensing products have been tried in the da Vinci® Surgical System (intuitive Surgical, Inc., Sunnyvale, CA) for providing haptic feedback to improve surgical outcomes, With the proved improvement on surgical outcomes comes the authentic

performance requirement for distributed-load sensing, I has been shown that the conventional distributed-load sensing products lack in size, spatial resolution, and force range/resolution for minimally invasive surgeries, Conversely, the critical factor of moving conventional

distributed-load sensors into the market is the low cost resulting from their relatively mature manufacturing technology as compared with the MEMS-based solutions under intensive development in the past two decades.

[0007] Microfluidic devices have been widely explored for various biological and chemical applications [1]. Generally speaking, these micro fluidic devices contain raicrochanneis or microchambers where fluids and/or particles are manipulated and analyzed. Owing to its low cost and biocompatibilitv, polydimethylsiloxane (PDMS) has become one of the most commonly used building material for microfluidic devices [2]. Standard fabrication technologies, including forming a patterned PDMS structure and bonding a PDMS structure to a glass substrate, has been well established for fabricating PDMS-based microfluidic devices. Disclosed herein is a sensor using PDMS rectangular microstructure with an embedded electrolyte-filled microchannei for detecting distributed loads, which are commonly encountered in biomedical [3], robotics [4], food processing and manufacturing applications [5],

[0008] In the meantime, miniature tactile sensors, employing different transduction mechanisms and fabricated using various fabrication technologies, have been developed for biomedical and robotics applications [6, 7], such as minimally invasive surgeries, tissue health diagnostics, and robotic fingertips. A general trend in the tactile sensing technology has been shifting the structural material from silicon to various polymers [8-13], due to their low elastic strength, which is compatible with a wide range of soft biological and viscoelastic materials,

biocompatibility and removal of the need for protective packaging. Among these polymer-based tactile sensors, a few of them are essentially microfluidic devices in the sense that these devices contain a polymer microstructure filled with electrolyte. For instance, Gutierrez et al [1 1] developed a parylene force sensor containing a parylene microchamber filled with an electrolyte. Tseng et al [12] demonstrated a PDMS/ olyimide tactile sensor including a hemispheric microchamber filled with an electrolyte and an initially empty microchannei. Park et al [13] developed a PDMS tactile sensor encompassing a PDMS microchannei filled with eGain.

[0009] The core of the above mentioned microfluidic tactile sensors is a polymer microstructure filled with electrolyte. In response to an external load, the microstructure generates a deflection, and, as impedance transducer, electrolyte in the microstructure converts the deflection to an impedance change. In terms of long-term stability, repeatability, and response time, the feasibility of employing electrolyte-filled polymer microsiructures as tactile/force sensors has been demonstrated by these microfluidic sensors. However, the parylene-based force sensor [1 1] involves great fabrication complexity, and its electrolyte filling method renders it unsuitable for operation in dry environment. Although the fabrication of the PDMS/polyiroide microfluidic device [12] is relatively simple, electrolyte filling and device operation is complex. In contrast, the PDMS tactile sensor [13] demonstrates great fabrication simplicity and ease of electrolyte filling. However, since the electrolyte in the whole PDMS microchannei is utilized for impedance transduction with the ends of the microchannei for electrical connection, this sensor is not capable of measuring distributed loads along the microchannei length.

[0010] In many practical applications, it is necessary to detect non-uniform distributed loads in aqueous and dry environments, which are induced by heterogeneous biological or viscoelastic materials or by texture of an object, For instance, Ahrnadi et al [3] demonstrated an optical fiber tactile sensor for detecting heterogeneous tissues in minimally invasive robotic surgery. In this tactile sensor, several fibers are embedded underneath a beam in order to convert the deflections at different, locations along the beam into optical signals. Since this sensor is manually assembled together, it has the drawback of large dimension and high cost. In order to detect the texture of an object, a sensor array with high spatial resolution (Imrn) is needed for a robotic fingertip [3, 6]. SUMMARY OF THE INVENTION

[0011 J Microfluidic devices are devices thai contain raicrochannels or micro chambers where fluids and/or particles are manipulated and analyzed. Disclosed is a po!ymer-based microfluidic sensing platform that can be configured to detect distributed loads at th micron-millimeter scale. In embodiments, the sensing platform comprises a unitary polymer microstructure integrated with electrolyte-enabled distributed transducers, The polymer microstructure is utilized as a sensing element to generate different deflections in response to distributed loads along the microstructure length, while electro I yte-enabled distributed transd ucers underneath the microstructure convert the deflections along the microstructure length to different resistance changes. This sensing platform can be fabricated using standard MEMS/microfiuidie fabrication technology as known to ordinarily skilled artisans.

[0012] In embodiments described herein, using a custom-built electronic circuit and a custom Lab VIEW program, the resistance changes from the distributed transducers can be measured and consequently the distributed loads can be extracted from the detected resistance changes: the sensing platform has been analyzed and verified using an embodiment of a sensing platform as described herein.

[0013] Disclosed are embodiments of devices, methods, and processes for fabricating a polymer-based microfluidic sensing platform for detecting distributed static and dynamic loads with a micron-millimeter (um-mm) spatial resolution in aqueous and dry environments. The sensing platform comprises a unitary polymer microstructure; and an electrolyte-enabled distributed transducer. The electrolyte-enabled distributed transducer comprises: a

microchannei formed in the unitary polymer microstructure configured to hold an electrolyte; and an electrode underneath the microchannei. The microfluidic sensing platform is configured ίο detect distributed loads at the rnieron-niillinieter scale. Detecting static and dynamic distributed loads with n μπϊ-rnm-rnm spatial resolution is needed various biomedical/surgical, robotics and manufacturing applications. Despite these needs, there are currently no reliable, affordable commercial products made using MEMS/raicrofluidic fabrication technology, and most such prototype devices still remain in the laboratory. By incorporating electrolyte-enabled distributed transducers in the design and employing well-developed standard fabrication technology for polymer-based icrofliridie devices, the disclosed sensing platform offers great simplicity in its design and fabrication, thus promising high reliability, low cost, and disposability, as compared with the commonly adopted sensor-array approach currently under intensive development. Demonstration of the feasibility of the proposed sensing platform for detecting static and dynamic distributed loads with n μη -ram spatial resolution, exploration of the performance space of the proposed sensing platform with the design limits imposed by the fabrication process, and the feasibility of the proposed sensing platform to meet the market needs at acceptable cost are contemplated.

[0014] A new generation of sensing products that will satisfy the current and future needs of customers in biomedical/surgical, robotics, and manufacturing industries. By providing the functionality of detecting distributed loads at an affordable cost, the proposed sensing platform will help the development of intelligent instruments, robotics and manufacturing equipment that rely on such sensing platforms to examine anatomical structures of tissues, provide haptic feedback to surgeons in minimally invasive surgeries, investigate the details of viscoe!astic materials for developing new manufacturing processes, as well as determine the texture of an object in robotics. Deploying this sensing platform in the associated systems is expected to deliver new value and add vital functionalities to these systems. Combination of the reliable performance and low fabrication cost positions the proposed solution for rapid

commercialization. Societal impacts in biomedical applications include improved surgical outcomes and affordable medical care. Utilization of these sensing products in manufacturing will increase quality, and in robotics will impact robot-environment interactions. The project will, also enhance the scientific and technological understanding in MEMS and microfluidics by developing a new-generation sensing platform that supports a wide range of designs and enables new products.

[0O1SJ The design, fabrication, and performance characterization of a PDMS-based microfluidic resistive sensor for detecting distributed loads is disclosed. This sensor is comprised of a PDMS rectangular microstructure with an embedded electrolyte-filled microchannel and an array of electrodes allocated along the microstructure length. Such sensors may be fabricated using a CNC machine. Electrolyte is filled into the sensor through the reservoirs at its ends. With a custom built electronic circuit and a custom LabVIEW program, the static and dynamic performance of the fabricated sensor is characterized, verifying the design concept of this sensor, As compared with the above mentioned polymer-based microfluidic tactile sensors, the disclosed sensor not only provides the capability of detecting distributed loads, but also offers quite a few advantages, including great fabrication simplicity, ease of electrolyte filling, and operation in both aqueous and dry environments.

0016] The feasibility of producing accurate, reliable, disposable polymer-based microfluidic sensing platforms for detecting static and dynamic distributed loads with a spatial resolution of ΙΟΟμηι-ln m is contemplated. The static and dynamic performance of a sensing platform can be correlated to its design parameters; and the performance space of the sensing platform can be explored with the design limits imposed by the fabrication process employed. [0017] Design, fabrication, and systematic experimental investigation of prototype polymer- based microti ui die sensing platforms for specific targeted applications, for example

minimally invasive surgeries, are contemplated. A sensing platform can be expanded into a 2D array by manufacturing a number of polymer raicrostructures in parallel using the same mask, which is not expected to lower the yield rate, Sensing platforms for other applications is contemplated, such as for example robotics and manufacturing processes of viscoelastic materials and biomateriaJs. It is further contemplated design sensing platforms for achieving more functionality, including measuring the softness of a tissue.

[0018] As will be appreciated, sensing platforms in accord with the embodiments set out herein will be more cost-effective and cost-efficient than conventional distributed-load sensing products, as a small amount of raw materials is needed; a high yield rate can be expected from the standard polymer-based micro-fabrication process; and the same mold can be repeatedly used for producing large quantities of the same disposable sensing platforms. Moreover, disposability of the disclosed sensing platform(s) is more suitable for minimally invasive surgeries, since no sterilization is needed after use. Further, the disclosed sensing platform offers the potential of developing into many different models of products for different markets and being brought into the market sooner due to its fast turn-around, low-cost production, as compared to the sensor- array solutions based on MEMS fabrication technology,

BRIEF DESCRIPTION OF DRAWINGS

[0019] To the accomplishment of the foregoing and related ends, certain illustrative

embodiments of the invention are described herein in connection with the following description and the annexed drawings. [0020] FIG, 1. (a) Schematic and (b) picture of the proposed polymer-based microfluidic sensing platform.

[0021] FIG, 2, Equivalent electrical circuit of each distributed transducer of the polymer-based microfluidic sensing platform,

[0022] FIG, 3, (a) Side view and (b) top view of the polymer-based microfluidic sensing platform (drawn not to scale for better illustration).

[0023] FIGo 4, Fabrication process of the polymer-based microfluidic resistive sensor (a) deposit and pattern Au/Cr electrode on a Pyrex slide (b) fabricate a SU8 mold for molding the polymer micro structure (c) form the polymer raicrostructure using the mold and (d) bond the poly m er micro structure to the patterned Pyrex slide.

[0024] FIG, 5, Pictures of the fabricated sensing platforms (a) filled with colored liquid to show electrolyte-enabled transducers and (b) bonded with Al/Si wires for electrical connections.

[0025] FIG. 5(c). Picture of the custom electronic circuit on breadboard.

[0026] FIG, 6, Measured different resistances at the I ^sS transducer due to loading at A and B of a sensing platform with electrolyte transducers of 12mm x 1mm x 200pm.

[0027] FIG. 7, Measured resistance of the 2 ^nd transducer in response to a dynamic load at location A of a sensing platform with electrolyte transducers of 12mm x 1mm x 70μηι.

[0028] FIG, 8. Schematic of a custom electronic circuit connected to one transducer of the fabricated device for measuring its resistance.

[0029] FIG, (a) Schematic of the experimental setup for testing and (b) a picture of the custom circular probe of 4mm in diameter above the sensing platform.

[003Θ] FIG. 10, Pre-defined deflection pattern of the circular probe. [0031] FIG, 11, Measured results from the 1 ^st pair of electrodes, when the custom probe is located at A and B, respectively (vp _P ^:::: l V and ω= 200kHz) (a) locations of the probe (b) DC voltage output, and (c) resistance calculated form the DC voltage output,

[0032] FIG, 12. Measured results from the 3 ^:<! pair of electrodes, when the custom probe is located at A and B respectively (vpp~3V and ω= ^: 200kHz) (a) locations of the probe (b) DC voltage output, and (e) calculated resistance from the DC voltage output

[0033] FIG, 13, A picture of the fabricated sensor showing electrolysis on the 3 ^!d electrode pair an air bubble across the 4 ⁱ electrode pair.

[0034] FIG. 14. Measured resistance values when the electrodes are connected together.

[0035] FIG. IS, Measured results from all the five pairs of electrodes, when the custom probe is located at A (a.) DC voltage output versus the deflection at A (b) the slope of the calculated resistance versus the deflection at A for the 3 ^rd electrode pair, and (c) the slopes of the calculated resistance versus the deflection ai A for the other four electrode pairs.

[0036| FIG, 16. Simulated deformation of the PDMS rectangular micro structure in response to a deflection of 1 ΟΟμηι of a 4mm-in-diameter circular probe located at the top of the 3 ^r " electrode pair (a) top view and (b) bottom view.

[0037] FIG. 17. Simulated deflection profile of the ceniral life along the raicroeharmel length, in response to different deflection levels of the circular probe located at the middle of the device.

DETAILED DESCRIP TION OF THE INVENTION

[0038] These embodiments are indicative, however, of but a few of the various ways in which the principles of the invention may be employed and the present invention is intended to include all such aspects and their equivalents. Other advantages, embodiments and novel features of the invention may become apparent from the following description of the invention when considered in conjunction with the drawings. The following description is given by way of example, but not intended to limit the invention solely to the specific embodiments described, which can be understood in conjunction with the materials that follow,

[0039] Distributed load sensing products are useful and advantageous in a number of fields, For example, capabilities to examine anatomical structures of tissues in tissue health analysis, provide indispensable haptic feedback to surgeons during tissue manipulation/palpation in minimally invasive robotic surgeries (MIRS), investigate the details of viseoelastie materials and bioniaterials for developing new manufacturing processes, and determine the texture of an object in contact with a robot hand, are needed in biomedical, manufacturing and robotics applications. From the mechanical perspective, these capabilities translate to detecting spatially distributed static and dynamic loads with a micron-to-millimeter (μιη-mm) spatial resolution. Therefore, accurate, reliable, and affordable distrihuted-load sensing products (or distributed tactile sensors) are needed to provide such functionality.

[Θ04Θ] The majority of distributed-load sensing devices with μηι-mm spatial resolution currently under development are the sensor-arrays, where multiple individual sensors are batch- produced using MEMS fabrication technology. Various transduction mechanisms

(piezoelectric, capacitive, optical, piezoresistive, and conductive-polymer resistive) and various structural materials (different polymers and silicon) have all been explored for developing these sensor-array solutions. However, these sensor arrays are costly and offer unreliable

performance, solely because multiple individual sensors lead to costly production and a large performance variation across individual sensors from fabrication variations, Therefore, a new approach is necessary in the design and fabrication of sensing products for real-time detecting distributed static and dynamic loads with μίτι-mm spatial resolution,

Π |0Θ4Ι ] Accordingly, disclosed are embodiments configurable and adaptable for producing accurate, reliable, and disposable polymer-based raicrofluidic sensing platforms for detecting distributed static and dynamic loads with μη -ηπη spatial resolution in aqueous and dry environments, In an embodiment, disclosed is: a single deformabie polymer micro structure integrated with electrolyte-enabled distributed transducers and a fabrication process resulting from this core design, one advantage of which being its low cost. The polymer microstructure is utilized to translate distributed loads to deflections, which are converted to electrical resistance changes by the electrolyte-enabled distributed transducers underneath the polymer microstructure. The analysis of the sensing platform described herein demonstrates its capability of detecting static and dynamic distributed loads with a 1.5mm spatial resolution with a time constant less than 100ms,

[CI042] It will be appreciated that: in comparison with commonly adopted sensor-array approach, the disclosed embodiments can provide a better solution to the needs of detecting distributed loads in biomedical, robotics, and manufacturing applications due to its simplicity in design and fabrication and the ability to function in aqueous and dry environments,

[0043] TECHNICAL ADVANTAGES

[0044] SENSOR-ARRAY SOLUTIONS BASED ON MEMS FABRICATION TECHNOLOGY [0045] The small-size and batch-fabrication nature of the MEMS fabrication technology intuitively leads to its application for producing next-generation distributed-load sensing products to meet the requirement on size, tight spatial resolution, and force range/resolution in, for example, minimally invasive surgeries. These MEMS-based distributed-load sensing solutions are sensor arrays, where multiple individual force sensors are arranged into a 2D array with a spatial resolution (e.g., <lmm for MIRS) in order to obtain the distributed loads across a small area (robotics) or anatomical structures across a small piece of tissue (MIRS). Various transduction mechanisms have been explored and various polymers and silicon have been adopted as the structural material in developing MEMS-based distributed-load sensing solutions, with polymers being identified as the suitable structural material The two main bottlenecks preventing the MEMS-based sensor-array solutions from commercialization have remained: 1) high-cost from their costly fabrication process involving multiple masks and multiple depositions/etchings and 2) a large performance variation across individual sensors due to high fabrication tolerances, regardless of transduction mechanisms employed and structural materials used.

[0046] As compared with MEMS-based aeceierometers, gyroscopes and microphones (singular devices), bringing MEMS-based distributed-load sensing devices into the market has proven to be extremely challenging, simply due to the array nature of the distributed-loaded sensing devices. For example, it took Texas instruments (Tl) more than 20 years (1980's to 2000's) to move their array-based (more than 1 million individual mirrors in a 2D mirror array) projection display technology (DLP) into the market, It would be even more challenging to move array- based distributed-load sensing devices to the market, since a direct contact between such devices to their working environment is required. Nevertheless, various sensor arrays based on MEMS fabrication technology have been developed, yet they still remain in the laboratory, [0047] FIG. 1 shows (a) Schematic and (b) picture of the disclosed polymer-based

microfluidic sensing platform. The polymer-based microfiuidic sensing platform promises a new generation of distributed-load sensing devices with numerous advantages, including high sensitivity/accuracy, reliability, disposahility, and at the same time, capability of detecting static and dynamic loads and being operated in aqueous and dry environments. In the embodiment, the sensing platform design incorporates one single deformable polymer microstructure as the sole sensing element to convert distributed loads with μπι-mm spatial resolution to deflections and electrolyte-enabled distributed transducers confined within the microchannei to convert the deflections in the microstructure to electrical resistance changes, which are then recorded by the electrode pairs underneath the microchannei. Two reservoirs at the ends of the microstructure are configured to fill the microchannei with an electrolyte and provide a conduit for electrol te in the microchannei to flow in out during the device

operation.

[0048] This sensor is comprised of a polymer rectangular structure with an embedded electrolyte-filled microchannei and five pairs of metal electrodes. Two reservoirs at the ends of the microstructure are utilized to fill the microchannei with an electrolyte and provide a conduit for electrolyte in the microchannei to flow in/out during the sensor operation. Each electrode pair has opposing electrodes along the channel width, and they record a change of electrical resistance across them. Five pairs of electrodes are allocated along the microchannei length in order to detect the distributed loads acting on the microstructure surface. A rectangular microstructure design is chosen here for simplicity. Certainly, other structural geometries can be easily realized given that the sensor design is compatible with the fabrication process as described later on. Selection of PDMS as the building material for the rectangular microstructure offers great fabrication simplicity and biocompatibility.

[ΘΘ49] Together with an electrode pair, the portion of electrolyte between them forms a transducer for converting the deflection above the electrolyte to a resistance change.

[0050] FIG. 2 illustrates an equivalent electrical circuit of each distributed transducer of the polymer-based microfmidic sensing platform. Electrolyte across the electrodes can be simplified as a resistor, R _s , and a capacitor, C\ in parallel. Owing to the electrical double layer formed at the interface between an electrode and electrolyte, each electrolyte-electrode interface is treated as a double layer capacitor, C<JI, and a charg transfer resistor, in series. Therefore, the impedance across the two electrodes is written as: where ZDL denotes the impedance of each electrolyte-electrolyte interface, and & denotes the frequency of the ac voltage signal, V _ac ,{&), which is applied to one electrode during operation, By choosing an appropriate electrolyte and the operation frequency of the ac voltage signal, the impedance of the electrolyte-electrode interfaces and the capacitance across the electrode pair can be neglected. Then, the device can be treated solely as a resistor. Rather than capaeitive sensing, resistive sensing is chosen to eliminate cross-talk or interference from its working environment.

[0051] FIG. 3 illustrates the operation principle of the sensing platform, together with the key design parameters of a PDMS-based sensor, The key design parameters and their values of the device are summarized in Table 1.

[0Θ52] Table 1 Key design parameters and their values of the PDMS-based micro fluidic resistive sensor

0053] FIG, 3(a) shows a side view and FIG. 3(b) a plan view of the polymer-based microiluidie sensing platform (drawn not to scale for better illustration). Distributed loads acting on the device surface deflect the compliant polymer microstructure at different levels along the microstructure length and compress partial electrolyte into the reservoirs.

Consequently, the cross section of the resistor formed by electrolyte across each electrode pair is reduced and this reduced cross section registers as a resistance change. The resistance change across each transducer is related to its cross section by:

R, ^■ ■ ^■ ■ ^■ ■ ^· · . (2) where ^P _E s the electrical conductivity of electrolyte; Z½ is the distance across an electrode pair;

AE s the microehannel cross-section, which is a function of the deflection, z, of the top of the microstructure; and Subscript i denotes the parameters associated with the it transducer. By- measuring the resistance change at the transducers distributed along the microstructure length, distributed loads acting on the device surface can be obtained,

[0054] Subjected to distributed loads, electrolyte in the microehannel flows into the reservoirs, since electrolyte in the sensor is incompressible. Thus, adding the reservoirs not only completely confines electrolyte within the sensor, hut also allows electrolyte to freely flow during operation. This device can be operated in either aqueous or air environment, since electrolyte is confined within the device and the building materials of the device are nonconductive, except that the portion of the electrodes outside the Polymer microstructure needs to be covered with an insulating layer for operation in aqueous environment, [0055] An embodiment of a process for fabricating the device is described. FIG, 4 shows a fabrication process of the polymer-based mierofluidic resistive sensor (a) deposit and pattern Au/ ^' Cr electrode on a Pyrex slide (b) fabricate a SU8 mold for molding the polymer

microstructure (c) form the polymer microstructure using the mold and (d) bond the polymer microstructure to the patterned Pyrex slide. The polymer-based standard fabrication process, which has been well established in the microfluidics field, can be directly utilized to cost- effectively fabricate the sensing platform using two masks. A mold may be made from polycarbonate using a CNC machine. The height of the mold determines the total height of the polymer microstructure. The inverse of the microchanne! and the reservoirs is patterned at the bottom of the mold. First, 1 OOiim/lOnrn-thick Au Cr electrodes are deposited and patterned on one side of a Imm-thiek Pyrex slide using a lift-off process. Then, SU8 is deposited and patterned on another Pyrex substrate. The thickness of the patterned SU8 defines the height of the microcharmel of the device, A 10:1 ratio of PDMS elastomer to curing agent (Sylgard 184kit, Dow Corning Corp.) is poured over the mold and cured to form the PDMS microstructure. The thickness of the PDMS microstructure is controlled by the amount of mixture used. The cured PDMS structure is peeled off from the SU8 mold. The Pyrex slide with patterned electrode and the PDMS microstructure are treated with oxygen plasma and bonded together. Finally, a hole is punched into each reservoir and the microcharmel is filled with an electrolyte solution using a syringe. Holes in the reservoirs can be further sealed or connected to a tube to avoid leakage.

[0056] FIG, 5 shows pictures of the fabricated device, Pictures of the fabricated sensing platforms at FIG 5(a) are shown filled with colored liquid to show electrolyte-enabled transducers and are shown at FIG 5(b) bonded with Al/Si wires for electrical connections. The electrolyte such as KCI is confined within a microchannei of 12mm x 1mm x 200μ ^η ι, The spatial resolution of the transducers is 1.5mm. The distance between the reservoir centers is 15mm and the width of the polymer microstruciure is 5mm. As compared to more than four masks and multiple deposition/etching steps needed for fabricating multiple individual sensors in a MEMS- based sensor array, the fabrication process for the proposed sensing platform is much simpler, solely due to the employment of electrolyte-enabled distributed transducers.

[0057] Experimental data validating the capability of employing electrolyte-enabled

distributed transducers to detect distributed static and dynamic loads for miniature sensing solutions is given here,

[0058] Note that due to the lack of a load cell for recording a load, the load applied to the device needs be obtained from the numerical simulation of the polymer microstruciure with a given deflection condition. Also, data are recorded for one transducer at a given time. FIG, 6 shows measured different resistances at the 1 st transducer due to loading at A and B of a sensing platform with electrolyte transducers of 12mm x 1mm x 200μηι [32]. FIG 6{b) shows the measured resistance from the 1st transducer, when a custom circular probe is located at A and B, respectively, deflecting the device according to a pre-defined patter shown in FIG, 6(a). The difference in measured resistance demonstrates the capability of electrolyte-enabled transducers to detect distributed loads with a spatial resolution of 1 ,5mm.

[0059] FIG. 7 illustrates how measured resistance varies with a pre-defined deflection at location A of another sensing platform, showing that the proposed sensing platform can monitor dynamic loads without any time delay. As shown in FIG. 7; measured resistance of the 2nd transducer in response to a dynamic load at location A of a sensing platform with electrolyte transducers of 12mm x lmm x 70μπι. PERFORMANCE CHARACTERIZATION

0Θ6Θ] Electronic circuit

[0061] To test the fabricated device, a custom electronic circuit is designed and implemented on breadboard for measuring the resistance across a pair of electrodes, when the device is subjected to external loading, As shown in FIG 8, the circuit for detecting die resistance across a pair of electrodes contains a transimpedance amplifier and a demodulation stage. An ae voltage signal, Vac(®) ₅ s applied to one electrode, while the current, /(e>), coming out from the other electrode, feeds in the inverting terminal of the OP-AMP (OPA656U), The sensing electrode is maintained at virtual ground by connecting the non-inverting terminal to the ground in order to minimize the effect of parasitics on this current signal. For simplicity, the output, v _\ , of the transimpedance amplifier serves as two identical inputs for the multiplier (AD835) to avoid phase difference between the two inputs, The DC component, V _out , of the voltage output, v ₂ , of the multiplier, passes the following third order low-pass filter (LPF) and gives rise to the measurement of the resistance across the electrode pair. As shown in FIG, 5(c), this circuit is implemented on breadboard and can monitor the resistance across a pair of electrodes at a given time. To monitor all the five pairs of electrodes simultaneously, five similar electronic circuits and five ac signals are needed. The time constant of this circuit is simulated to be 11 ,6ms, as shown in the graph in FIG 5(d).

[0062] Now. we relate the DC voltage output of the electronic circuit to the resistance across a pair of electrodes of the sensor. Assume that the ac voltage signal is expressed as:

where vpp and ω denote the peak-peak (p~p) val ue and the frequency of the ac voltage, Then, the current output from the device is given by; (4)

The ac voltage output of the transimpedance amplifier is:

The DC voltage output of the demodulation stage is:

[0063] By using a few resistors with known values., the relation of the DC voltage output to the resistance from the electronic circuit is verified. Consequently, the resistance across a pair of electrodes of the sensor can be extracted from the measured DC voltage using the following relation:

The overal l sensitivity of the sensor to the resistance becomes:

[0064] Hence, a large p-p value of the ac voltage and a large feedback resistance of the transimpedance amplifier contribute to a higher sensitivity of the sensor, Based on the initial resistance and the chosen operation frequency for the ac voltage signal, the resistance of the feedback resistor is fixed at i¾ ^:::: 970O. The p-p value of the ac voltage is varied to keep the DC voltage output around 0,12V when the sensor is not subjected to external loading, while keeping the OP-AMP and the multiplier working in their operation range.

[0065] Experimental setup and method

[0066] FIG 9(a) shows the experimental setup for characterizing the performance of the PDMS- based mierofluidic resistive sensor. A fabricated sensor is mounted on a PCB, where wires are bonded between electrodes of the sensor and copper electrodes on a PCB, as shown in FIG 5.. Then, the PCB is mounted on a custom fixture, which is further fixed on an optical table. A custom circular probe of 4mm in diameter is mounted on a micropositioner, as shown in FIG 9(b), A controller associated with the micropositioner can precisely move the circular probe along the z-axis with a resolution of 0.2μηι. Before each measurement, the circular probe is brought in contact with the PDMS rectangular m cro structure surface, without deflecting the sensor. This is achieved by monitoring the change in resistance.

[0067] To measure the resistance across a pair of electrode, a function generator (HP33220A) is connected to one electrode for providing an ac voltage signal, while the other electrode is connected to the circuit on breadboard. The DC voltage output signal from the circuit is connected to PCI-6133 DAQ board, which feeds in a custom Lab VIEW program for recording data every 0.1s for approximately 70sec, The recorded DC voltage output is then converted to resistance, according to Eq. (7). An important operation parameter is the frequency of the ac voltage. Care must be taken to ensure that this sensor is predominantly resistive so that any change in the measured DC voltage is solely caused by the resistance change across a pair of electrodes. Here, the frequency of the ac voltage signal is chosen to be 200kHz.

[0068] DYNAMIC PERFORMANCE

[0069] Since detecting spatially distributed dynamic loads is indispensable, the time constant of the proposed sensing platform can be analyzed. It should be pointed out that due to the complexity of the stracture-fl id-electrochemistry involved in the device, as is best

understood, no analytical study has been conducted to obtain the time constant of such a device, though experimental studies have provides some values. The viscosity of the

electrolyte can be expected to play a role in determining the time constant of the device.

Therefore, NaCl solutions, among other solutions, with varying concentrations may be used. A numerical analysis on the FEA model that includes the microstriieture and the electrolyte (treated as viscous fluid) underneath is to be conducted. The time constant of the proposed device operating in aqueous and dry conditions were both analyzed to determine the

detectable frequency range of the dynamic loads,

[0070] Based on the results above, the design parameters affecting the performance of the microstructure and the performance of electrolyte-enabled distributed transducers have been systematically varied and incorporated into the two masks for experimental implementation in order to validate the predicted static performance and dynamic performance of the device versus the key design parameters. Mask layouts were made in AUTOCAD and the devices using the two masks fabricated in a cleanroom.

[0071] As mentioned above, the standard polymer-based fabrication processes are utilized to fabricate the device. Although this fabrication process has been used to produce microfluidic devices, it has no previous to this disclosure been fully characterized for the embodiments described herein. To ensure the high yield rate while without challenging the limits of this standard polymer-based fabrication process, the following parameters were varied: the dimensions of the microchannel, the dimensions of the polymer microstructure, as well the dimensions of the electrode pairs so as to obtain the comfortable dimension ranges of these design parameters from the fabrication process so that the design of a device can be optimized, based on the design limits identified from the characterization of the fabrication process.

[0072] DYNAMIC PERFORMANCE - EXPERIMENTAL RESULTS

£0073] Using the controller associated with the microposiioner, a pre-defined deflection pattern, as shown in FIG. 10, is applied to deflect the sensor through the circular probe, The circular probe generates a deflection of 300μίη in the sensor at a constant speed of 300μηι/3 ₅ remains at this deflection for 3s, and then goes back to the un-defleeted position at the same constant speed. This deflection cycle repeats itself for roughly 70sec. The probe is aligned at different locations along the PDMS rectangular micro structure for deflecting the sensor with this pre-defined deflection pattern, and the DC voltage output is recorded across different pairs of electrodes, respectively. Due to the lack of enough equipment, the data is recorded for one pair of electrodes at a given time,

[0074] FIG ϋ shows the measured data from the 1 ^st pair of electrodes, when the circular probe is located at A and B, respectively, deflecting the sensor according to the pattern shown in FIG 10, The ae voltage is kept at v _PP =iV and e> ^~ 200 kHz in this measurement. When this sensor is free of external loading, the resistance across this electrode pair is about 1440Ω, based on the DC output of 0,1 14V. When this sensor is deflected by 300μηα at A, the DC output becomes 0.092V and thus the resistance increases to 1610Ω. The increase in resistance, is caused by the cross section reduction of the electrolyte across this electrode pair. Similarly, when the sensor is deflected by 300μηι at B, the DC output decreases to 0.88V and the resistance increases to 1640Ω. Since deflecting B causes more reduction in the cross section across this electrode pair than deflecting A, the resistance from deflecting B is higher than that from deflecting A.

[0075] FIG 12 shows the measured data from the 3 ^,d pair of electrodes. The probe is utilized to deflect the sensor at A and B, respectively. The DC voltage is read out from, the 3 ^rd electrode pair. The ac voltage is maintained at vp =3V and ω= 200 kHz in this measurement. Note that a high ac voltage is chosen in order to keep the original DC output around 0.12V.

[0076] When this sensor is free of external loading, the resistance across this electrode pair is about 4300Ω from the DC output of 0.17V. When this sensor is deflected by 300μηι at A, the DC output becomes 0,052V and thus the resistance increases to 6400Ω. The increase in resistance is caused by the cross section reduction of the electrolyte across this electrode pair. Similarly, when the sensor is deflected by 300μηι at B, the DC output decreases to 0,082V and the resistance increases 5100Ω. Since deflecting A causes more reduction in the cross section across the electrode pair than deflecting B, the calculated resistance from deflecting A is higher than that from deflecting B.

[0077] The measurement from the 1 ^st and 3 ^rd pairs of electrode validates that this sensor is capable of detecting distributed loads. Ideally, the values of the resistors across all the pairs of electrodes are expected to be very similar, if not exactly the same. The di fference in resistance between the l ^si and 3 ^Til electrode pairs is believed to be caused by electrolysis on the 3 ^]d electrode pair as a result of high ac voltage (See FIG 13),

[0078] As shown in FIG, 14. when the sensor is free of external loading, the resistance is about 430Ω (one input) when one electrode from the 3 ^rd electrode pair serves as the ac voltage input while all the five electrodes on the other side of the micro structure are connected together and serve as the output, in contrast, when the sensor is free of external loading, the measured resistance is about 1030Ω when ail the five electrodes on one side of the mierostrucittre are connected together and serve as the ac voltage input while the other five electrodes are connected together and serve as the output. Note that these resistance values are consistent with those obtained in measuring the 1 ^st and 3 ^rd electrode pairs, individually.

[0079] Finally, it should be pointed out that, as compared with the pre-defined deflection pattern, the DC output does not exhibit any time delay, proving that the time constant of the sensor itself is below 100ms, since the data is recorded at a time interval of 100ms, while the time constant of the electronic circuit, 12ms, is well below this time interval .

[§080] STATIC PERFORMANCE

[0081] Polymer mierostructure (Solid Mechanics): distributed loads→ deflections. [0082] Distributed static loads acting on a polymer microstmeture translate to deflections along the r crostructure length. Since electrolyte underneath the microsiructure can freely move to the reservoirs, the relation of the deflections to distributed ioads can be obtained by conducting static analysis of the microsiructure in finite element analysis (FEA) software, COSMOL. Since it is the top of the polymer mierostructure (or a rectangular thin plate), that responds to distributed loads, the key design parameters are the dimension of the rectangular plate: 1¾ ₂ x W _m s x L _ra (see FIG, 3), Together with the key design parameters from electrolyte-enabled distributed transducers, these three parameters are to be varied for obtaining their effect on the static performance of the device, including pressure resolution, sensitivity and range, as well spatial resolution.

[0083] Electrolyte- enabled distributed transducers (electrochemistry): deflection→

resistance changes.

[0084] Through electrolyte transducers across their electrode pairs, deflections along the microsiructure length are then converted to electrical resistance changes. NaCl solution may serve as the electrolyte, due to its working range and high stability, The concentration of NaCl solution will vary depending on the device dimension. The key design parameters of these transducers are the height and width of the microchannel (:1½2 x Wms), the distance across an electrode pair {¾) is the distance across an electrode pair; the distance between two neighboring transducers, spatial resolution (¾). Due to the geometrical irregularity of electrolyte-enabled transducers, the resistances of the transducers, which are deflected by the distributed loads, can be estimated based on the obtained values,

[0085] STATIC PERFORMANCE - EXPERIMENTAL RESULTS [0086] To obtain the relation of the resistance to the deflection of this sensor, the circular probe is located at A to deflect the sensor at different deflection levels, and the corresponding DC voltage output is recorded at each pair of electrodes, respectively, Note that the DC output is recorded from eac electrode pair separately, while the probe is kept at A, deflecting the sensor in the same way. FIG. 15(a) shows the recorded DC voltage versus the deflection at A.

According to the measured relation of the DC voltage versus the deflection at A, the resistance of each electrode pair is plotted against the deflection at A in FIGS. 15(b) and 15(c).

[0087] Since the probe is located right above the top of the 3 ^rd electrode pair, the resistance across this electrode pair changes the most, as evidenced by its slope, 11 ,541Ω μηι. The 2 ^nd and 4* pairs are close to the probe and demonstrate a slope of 1.1501Ω/μΐη and 0.3966 Ω/μηΐ, respectively. The 1 ^st and 5 ^ta pairs are further away from the probe, and thus demonstrate a low slope of 0,3062 Ω/μχη and 0.2108 Ω/μι», respectively. The relatively high resistance of the 5 ^th electrode pair is due to the existence of an air bubble during the measurement. It should be noted that after all the measurements were finished, the air bubble has moved to the location across the 4 ^lh electrode pair, as shown in FIG, 13, It is believed that misalignment of the probe on the sensor is one of the causes for the large discrepancy on the measured slope of resistance to deflection, because the measurement was conducted on one electrode pair at a time, due to the lack of enough equipment. Another cause for the asymmetry in the measure slopes of the electrode pairs is the existence of the air bubble.

[0088] The maximum deflection of the probe used in the measurement is 300μηι, while the height of the microchannel is 2ϋϋμπι. From the 3 ^rd electrode pair in FIG 15(b). at the beginning, the probe is not in full contact with the sensor surface, since the resistance remains flat for the deflection of a few microns. The resistance becomes flat again, indicating that the top of the PDMS microstructure has touched the bottom of the mieroehannel shortly after the deflection passes 250μτη. Therefore, the rest of the 300μηι deflection is caused by the deformation of the top of PDMS micro structure.

[0089] O wing to the lack of a. force/pressure sensor, no force is directly recorded in all the measurements. Therefore, the force experienced by the sensor when the probe deflects the sensor is obtained by resorting to fini te element modeling. To predi ct the relation of the deflection of the PDMS rectangular microstructure to distributed loads, a nonlinear finite element model is created in COMSOL software. Since a circular probe of 4mm in diameter is utilized to deflect the sensor at different locations to mimic distributed loads. Therefore, the loading condition applied to the sensor in the model is different deflection levels ax different locations along the microstructure length. The simulation results yield the force acting on the device from the deflection of the circular probe. Electrolyte in the microchannel is not. expected to affect the deflection-force relation of the device. Thus, only the PDMS rectangular microstructure is modeled. The material properties of PDMS used in the model are listed in Table 2,

[0090] Table 2 Key design parameters of the PDMS-based microfluidic resistive tactile sensors

[009.1 ] FIG 16 shows the deformation in the PDMS microstructure under a deflection of ΙΟΟμπι of the circular probe located at the top of the 3rd pair of electrodes. The simulation gives rise to a force of LOIN,

[C 092] FIG 17 shows the simulated deilection profile of the bottom central line along the microstructure length, in response to different deflection levels of the circular probe located at A. The deflection profile of the top central line along the microstructure length under a deflection of Ι ΟΟμιτι is also shown in the figure, From this figure, it is clear thai a force acting on the center of the device has a much larger effect on the deflection across the neighboring pairs of electrodes than the pairs at the far ends,

[0093] While the invention has been described and illustrated with reference to certain preferred embodiments herein, other embodiments are possible. Additionally, as such, the foregoing illustrative embodiments, examples, features, advantages, and attendant advantages are not meant to be limiting of the present invention, as the invention may be practiced according to various alternative embodiments, as well as without necessarily providing, for example, one or more of the features, advantages, and attendant advantages that may be provided by the foregoing illustrative embodiments,

[0094] Systems and modules described herein may comprise or interact with software, firmware, hardware, or any combination(s) of software, firmware, or hardware suitable for the purposes described herein. Software and other modules may reside on servers, workstations, personal computers, computerized tablets. PDAs, and other devices suitable for the purposes described herein. Software and other modules may be accessible via local memory, via a network, via a browser or other application in an ASP context, or via other means suitable for the purposes described herein. Data structures described herein may comprise computer files, variables, programming arrays, programming structures, or any electronic information storage schemes or methods, or any combinations thereof, suitable for the puiposes described herein. User interface elements described herein may comprise elements from graphical user interfaces, command line interfaces, and other interfaces suitable for the purposes described herein , Except to the extent necessary or inherent in the processes themselves, no particular order to steps or stages of methods or processes described in this disclosure, including the Figures, is implied. In many cases the order of process steps may be varied, and various illustrative steps may be combined, altered, or omitted, without changing the purpose, effect or import of the methods described.

[0095] Accordingly, while the invention has been described and illustrated in connection with preferred embodiments, many variations and modifications as will be evident to those skilled in this art may he made without departing from the scope of the invention, and the invention is thus not to be limited to the precise details of methodology or construction set forth above, as such variations and modification are intended to be included within the scope of the invention. Therefore, the scope of the appended claims should not be limited to the description and illustrations of the embodiments contained herein.

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