The authors would like to acknowledge the BPI PIAVE AUTONOTEX grant for the financial support.

1. Patterning Conducting Polymers on Textile
<ol>
	<li>Fix a 10 cm x 10 cm textile sheet on a planar surface for easy handling during the process. For the textile, use a 100% interlock knit polyester fabric with a thickness of 300 &#181;m and a knit direction stretch capability up to 50%.</li>
	<li>To make a mask containing the patterning design, use a 125 &#181;m-thick polyimide film; an example of the pattern is illustrated in Figure 1.
	<ol>
		<li>Use a laser cutter (e.g., Protolaser S, LPKF) to pattern the polyimide mask10; the pattern design of an electrode is illustrated in Figure 1.</li>
		<li>Coat the PDMS formulation (10:1 base to curing agent ratio) on top of the mask (polyimide film) using an automatic tape casting tool (K control print-coater, doctor blade) with a wet film thickness of 200 &#181;m and at a 6 m/min coating speed. Use about 0.5 mL for a mask of 3 cm x 5 cm. Perform this process under the fume hood.</li>
	</ol>
	</li>
	<li>Gently transfer the fabric to the PDMS-coated mask. Leave for 10 min, after which the PDMS should be fully absorbed in the textile structure.</li>
	<li>Cure the sample in an air-oven at 100 &#176;C for 10 min.</li>
	<li>Prepare the conducting polymer:PEDOT:PSS dispersion (80 mL), ethylene glycol (20 mL), 4-dodecylbenzenesulfonic acid (40 &#181;L), and 3-methacryloxypropyltrimethoxysilane (1 mL) in the fume hood.</li>
	<li>Brush-coat the PEDOT:PSS solution on the PDMS-free area of the textile until a homogenous penetration of the solution is obtained. Repeat this step to achieve a uniform pattern color. Apply about 1 mL/cm2.</li>
	<li>Cure the fabric at 110 &#176;C for 1 h to dry the PEDOT:PSS solution. Reduce the temperature to 60 &#176;C for textiles that are sensitive to high-temperature treatment, like nylon.</li>
</ol>
2. Organic Device Fabrication
NOTE: The protocol in Section 1 describes the selective deposition of conducting materials on textiles. The following sections will describe the additional steps needed to fabricate organic devices, like stretch sensors, OECT transistors, cutaneous electrodes, and capacitive sensors.
<ol>
	<li>To fabricate stretch sensors, shown in Figure 3a, pattern the electrode lines on the textile, as described in Section 1, steps 1.1-1.5. 
	NOTE: An example of the pattern design is shown in Figure 3a. The fabrication of such sensors does not require any additional steps.</li>
	<li>To fabricate the transistor design shown in Figure 3b, pattern the transistor arrays on a nylon woven ribbon following the steps described in Section 1. Slightly modify the PDMS annealing and PEDOT:PSS curing steps to avoid the thermal degradation of nylon by curing at 60 &#176;C for a longer time.</li>
	<li>For the fabrication of cutaneous electrodes, shown in Figure 3c, deposit an ionic gel on the patterned PEDOT:PSS textiles.
	<ol>
		<li>Prepare an ionic liquid gel mixture containing the ionic liquid, 1-ethyl-3-methylimidazolium-ethyl sulfate; the cross-linking agent, poly(ethylene glycol)diacrylate; and the photoinitiator, 2-hydroxy-2-methylpropiophenone at a (v/v) ratio of 0.6/0.35/0.05, respectively.</li>
		<li>Coat the PEDOT:PSS electrode with ionic liquid (20 &#181;L/cm2) and add the ionic liquid gel mixture from step 2.3.1 (25 &#181;L/cm2) by drop casting.</li>
		<li>Expose to UV light (365 nm) to initiate a crosslinking reaction for 10-15 min, until the gel solidifies. Perform this step in the fume hood. Use a UV-protective cage during UV exposure.</li>
	</ol>
	</li>
	<li>For capacitive sensor fabrication, use PEDOT:PSS textile electrodes insulated with an insulating material (Figure 3d).
	<ol>
		<li>Insulate the keyboard-like PEDOT:PSS electrodes using the PDMS; the keyboard design can be seen in Figure 2b. Dispense the PDMS formulation on top of the fabric and remove the excess with a squeegee.</li>
		<li>Place the fabric in an oven at 100 &#176;C for 10 min. Perform this step in the fume hood.</li>
	</ol>
	</li>
</ol>

<ol>
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The authors have nothing to disclose.

The patterning of conducting materials is one of the first steps in the fabrication of functional electronic devices. This can become challenging, as the fabrication process needs to take into account the chemical and physical properties of such materials, and the process flow needs to consider the material cross-compatibility between the fabrication steps. In the microfabrication of organic electronic devices, these two aspects are even more significant due to the highly reactive nature of organics. Today, however, organic materials are highly attractive to wearable and flexible electronics for their electro-elastic properties14,15. The transfer of such technologies to textiles to obtain fully integrated electronic wearables is limited by their three-dimensional structures. Conventional techniques used in microfabrication are restricted to inkjet or screen printing of conducting inks only on thin substrates and textiles16,17,18. Traditional embroidery technique, where a single fiber is sewn into the textile, still lacks industrial production scalability.
The most critical aspect in the patterning of organic materials is deposition without disturbing the electrical properties. The patterning technique described in Figure 1 relies on the direct deposition of organics, with no need to meet specifications of the deposition technique or the tool. The organic materials are formulated to the best of their performances and can then be directly deposited on the chosen fabric structure. The utility of PDMS is key to pattern materials from solution onto textiles. The application of conducting materials from the low-viscosity solution, rather than using paste-like inks, enables a conformal and deep coating in textile structure. However, it limits the selective deposition and leads to the loss of the patterning resolution. We have overcome this limitation by creating a negative pattern from PDMS to restrain the non-controlled conducting solution penetration into the textile. The strategic choice of the PDMS is based on its visco-elastic properties, which maintain the textile stretchability and flexibility. The PDMS is also hydrophobic and allows the control of the spreading of the PEDOT:PSS water-based solution during the patterning. We observed that the conductive patterns fabricated using this protocol demonstrated good electrical conductivity and stability during mechanical deformations. This method allows the future customization of existing garments with smart components that have electronic capabilities. However, one of the critical and, in some cases, limiting points of the proposed approach is still organic material durability in wearable conditions. Some aspects, such as mechanical stress resistance and behavior after the washing and drying of organic conducting textiles, are still unknown.
The large majority of wearable electronics rely on stretchable devices, where spring-like structures are created to maintain the electrical connection during device deformation. Depending on the textile type, the fibers in knit fabrics are assembled in a horseshoe design, providing mechanical stretchability of the structure. Coating these textiles with conducting materials permits individual fibers to act as strain and motion sensors in smart clothing, as shown in Figure 3a. Moreover, more complex device geometries can easily be patterned, not only on knit, but also on woven fabrics. In Figure 3b, we present an array of OECTs with variable geometries. In conventional photolithography, the simultaneous fabrication of big and small features is almost impossible to achieve without requiring multiple steps. We demonstrate that our patterning technique is able to produce patterns with a resolution that varies from 0.5 mm to about a hundred times greater. Such transistors can be directly used in wearable sweat sensing with an adjustable time response and detection resolution19.
We have demonstrated that PDMS also enables the consecutive deposition of additional functional layers in a selective manner, as shown in Figure 2. Devices can then be integrated in textiles and become fully integrated on wearable systems. The process in Figure 2a shows the fabrication of a cutaneous textile electrode, where the contact between the electrode and the skin is enhanced with an ionic liquid gel. Wearable electrodes in cutaneous electrophysiology suffer from motion artifacts caused by the electrical contact degradation between the wearer and electrodes during recordings. The possibility to integrate ionic gels on textile electrodes opens an efficient communication channel with the human body, which is desired in wearable healthcare devices. An example of such a device can be seen in Figure 3c.
The consecutive deposition of other active materials can result in devices using a stack geometry, such as organic batteries, capacitors, solar cells, transistors, or sensors. Figure 2b shows the deposition route of insulating or dielectric materials. A wearable organic keyboard (Figure 3d) can be fabricated using this process, where the PDMS is used to create a dielectric layer on top of the electrode. Such a device is capable of capacitive variation sensing between the electrode and a finger, which can have potentially interesting applications in wearable computing and human-machine interfacing.

The field of wearable electronics is a fast-growing market expected to be worth 50 billion euros in 2025, over three times the current market. The main challenge facing current wearable devices is that intrusive solid electronic attachments limit the usage of established devices in wearable systems. Using textiles that are already present in everyday life is a very attractive and straightforward approach to avoid this limitation. Due to its elastic capability, some parts of the clothing that we wear are naturally in tight contact with the skin. Many examples of smart clothes available on the market today are based on thin, plastic displays, keyboards, and light source devices embedded in textiles, linking electronics with humans in a fashionable way1. In sport practice, health monitoring relies on textile electrodes, which offer comfortable alternatives to commonly used adhesive electrodes and metal wristbands. Here, conductive fibers are directly integrated with stretchy fabrics to prevent skin irritation and other discomforts during extended wear. Additionally, textiles offer a number of opportunities to integrate curvature sensors to capture motion2, to integrate shear sensors for the development of functional robotic actuators3, and certainly to integrate biosensors through the detection of an analyte in sweat4.
Modern wearable technology relies on carbon-based semiconductor materials that deliver electronic devices with unique properties. The &#34;soft&#34; nature of organics offers better mechanical properties for interfacing with the human body compared to traditional solid-state electronics. This mechanical compatibility, paired with mechanically flexible substrates, enables the use of non-planar form factors in devices such as textiles. The use of organics is also relevant in life sciences due to their mixed electronic and ionic conductivity5. Besides, organic semiconducting and optoelectronic materials empower a large variety of functional devices with display, transistor, logic, and power capabilities6,7,8,9. The main difficulty in the fabrication of such organic devices is the controlled deposition of functional materials on the non-planar surfaces of textiles. Conventional microfabrication techniques are primarily limited by the incompatibility of the deposition process with the structural dimensionality of textile substrates.
Here, we describe a simple and scalable fabrication protocol that allows for the selective deposition of conducting polymers on structured textiles. The presented process enables the fabrication of wearable and conformal electronic devices. The approach is based on the patterning of the commercially available conducting polymer poly(3,4-ethylenedioxythiophene):poly(styrene sulfonate) (PEDOT:PSS) and an elastomeric stencil material polydimethylsiloxane (PDMS) on textile. This combination allows for the efficient confinement of the aqueous PEDOT:PSS solution, as well as for the retention of the soft and stretchable properties of textiles. This simple and reliable fabrication method paves the way for the fabrication of a variety of electronic devices directly on textiles in a cost-efficient and industrially scalable manner.

<table border="0" cellpadding="0" cellspacing="0" width="748">
	<colgroup>
		<col />
		<col />
		<col />
		<col />
	</colgroup>
	<tbody>
		<tr height="42">
			<td height="42" style="height:42px;width:304px;">SYLGARD 184, Silicone elastomer kit (Base and Curing agent)</td>
			<td style="width:159px;">Dow Corning</td>
			<td style="width:119px;"></td>
			<td style="width:167px;">PDMS elastomer</td>
		</tr>
		<tr height="21">
			<td height="21" style="height:21px;width:304px;"></td>
			<td style="width:159px;"></td>
			<td style="width:119px;"></td>
			<td></td>
		</tr>
		<tr height="21">
			<td colspan="4" height="21" style="height:21px;width:748px;">The conducting polymer formulation</td>
		</tr>
		<tr height="21">
			<td height="21" style="height:21px;width:304px;">CleviosTM PH 1000 PEDOT:PSS</td>
			<td style="width:159px;">Heraeus</td>
			<td style="width:119px;"></td>
			<td>Conductive polymer</td>
		</tr>
		<tr height="42">
			<td height="42" style="height:42px;width:304px;">Ethylene glycol</td>
			<td style="width:159px;">Sigma-Aldrich</td>
			<td style="width:119px;">03750-250ML</td>
			<td>Solvent (EG), CAS: 107-21-1</td>
		</tr>
		<tr height="42">
			<td height="42" style="height:42px;width:304px;">3-methacryloxypropyltrimethoxysilane</td>
			<td style="width:159px;">Sigma-Aldrich</td>
			<td>M6514</td>
			<td>Cros linker (GOPs), CAS: 2530-85-0</td>
		</tr>
		<tr height="42">
			<td height="42" style="height:42px;width:304px;">4-dodecylbenzenesulfonic acid</td>
			<td style="width:159px;">Sigma-Aldrich</td>
			<td style="width:119px;">44198</td>
			<td>(DBSA), CAS: 121-65-3</td>
		</tr>
		<tr height="21">
			<td height="21" style="height:21px;width:304px;"></td>
			<td style="width:159px;"></td>
			<td style="width:119px;"></td>
			<td></td>
		</tr>
		<tr height="21">
			<td colspan="4" height="21" style="height:21px;width:748px;">The ionic liquid gel</td>
		</tr>
		<tr height="42">
			<td height="42" style="height:42px;width:304px;">UV lamp DFE 2340</td>
			<td style="width:159px;">C.I.F/ ATHELEC</td>
			<td style="width:119px;">DP134</td>
			<td>UV-365nm</td>
		</tr>
		<tr height="42">
			<td height="42" style="height:42px;width:304px;">1-Ethyl-3-methylimidazolium ethyl sulfate</td>
			<td style="width:159px;">Sigma-Aldrich</td>
			<td style="width:119px;">51682-100G-F</td>
			<td>Ionic Liquid (IL), CAS: 342573-75-5</td>
		</tr>
		<tr height="42">
			<td height="42" style="height:42px;width:304px;">Poly(ethylene glycol) diacrylate</td>
			<td style="width:159px;">Sigma-Aldrich</td>
			<td style="width:119px;">455008-100ML</td>
			<td>Mn 700, CAS: 26570-48-9</td>
		</tr>
		<tr height="42">
			<td height="42" style="height:42px;width:304px;">2-Hydroxy-2-methylpropiophenon</td>
			<td style="width:159px;">Sigma-Aldrich</td>
			<td style="width:119px;">405655-50ML</td>
			<td>Phot Initiator (PI), CAS: 7473-98-5</td>
		</tr>
		<tr height="21">
			<td height="21" style="height:21px;width:304px;"></td>
			<td style="width:159px;"></td>
			<td style="width:119px;"></td>
			<td></td>
		</tr>
		<tr height="21">
			<td height="21" style="height:21px;width:304px;"></td>
			<td style="width:159px;"></td>
			<td style="width:119px;"></td>
			<td></td>
		</tr>
		<tr height="42">
			<td height="42" style="height:42px;width:304px;">The textile fabric</td>
			<td style="width:159px;">VWR</td>
			<td style="width:119px;">Spec-Wipe 7 Wipers</td>
			<td>100% interlock knit polyester fabric</td>
		</tr>
		<tr height="21">
			<td height="21" style="height:21px;width:304px;">The polyimide film</td>
			<td style="width:159px;">DuPont</td>
			<td style="width:119px;">HN100</td>
			<td>Polyimide film with 125 &#181;m thickness</td>
		</tr>
	</tbody>
</table>

a simple and scalable fabrication method for organic electronic devices on textiles

Today, wearable electronics devices combine a large variety of functional, stretchable, and flexible technologies. However, in many cases, these devices cannot be worn under everyday conditions. Therefore, textiles are commonly considered the best substrate to accommodate electronic devices in wearable use. In this paper, we describe how to selectively pattern organic electroactive materials on textiles from a solution in an easy and scalable manner. This versatile deposition technique enables the fabrication of wearable organic electronic devices on clothes.

Traditional methods for applying colors or patterns to textiles rely on removable masking layers to allow the selective deposition of dyes. In Figure 1, we show the adaptation of such an approach to the patterning of PEDOT:PSS electrodes on textiles. As a masking layer, we used hydrophobic polydimethylsiloxane, which can restrain the non-controllable diffusion of the aqueous PEDOT:PSS solution. Moreover, the softness and stretchability of knitted and woven textiles can be preserved thanks to the elastic and mechanical properties of the PDMS.
In Figure 1, the process starts with the preparation of the patterning master from polyimide film (step 1). The design of the pattern outline is carved onto the film by a laser. Using a tape casting tool, the PDMS is applied on this master (step 2), and the textile is placed on top of it (step 3). The PDMS is then progressively diffused into the textile (step 4). To stop this transfer, a short thermal annealing process is required to cure the PDMS. The viscosity and thickness of the PDMS can be adjusted by using different amounts of the curing agent and coating parameters, respectively, to control the diffusion and to insure the faultless replication of the master design. Finally, the conducting solution is brush-painted on the unprotected textile and baked to dry (step 5). The polyimide master is then delaminated from the textile surface. The results of the fabrication flow are illustrated in Figure 1, on the right. In this case, successful patterning was placed on knitted polyester. The patterning resolution on such a textile is greater than 1 mm. However, lower resolution can also be obtained on tightly knit or woven textiles. Using this deposition technique, the estimated sheet resistance of the conducting textile is close to 230 &#937;/sq.
Examples of functional electronic devices on knit and woven textiles are shown in Figure 3a and b, including successfully fabricated PEDOT:PSS electrodes on knit textiles. The natural horseshoe arrangement of the fibers in knit textiles provides adjustable stretchability to fabrics. This spring-like capability of knit structures can result in highly sensitive strain sensors11. A simple deformation in the textile structure is reflected by a change in the electrical resistivity due to the twisting of conductive fibers in the threads. Additionally, by taking advantage of the hygroscopic capacity of the textiles, the array of electrodes in Figure 3b was patterned on textiles to make planar transistors with rectangular channels and different gate widths, which can be used in wearable sweat sensing. Such a geometrical configuration is used in organic electrochemical transistors (OECT) for sensing which channel and gate are linked by a sample of an analyte12.
The presented patterning technique can be extended to fabricate complex organic electronic devices on textiles. As the PDMS stencil remains in the textile after the patterning process, additional layers can be patterned on PEDOT:PSS-coated conducting textiles. In Figure 2, we present the process in which an ionic liquid gel solution (Figure 2a) and the PDMS formulation (Figure 2b) were applied to functionalize or isolate the surface of a PEDOT:PSS electrode, respectively. Ionic gels are largely used in cutaneous electrodes. The incorporation of an ionic gel in conducting textiles was used to fabricate wearable textile electrodes for electrophysiological monitoring10 and is illustrated in Figure 3c. Capacitive sensors were made by insulating the textile electrode surface with PDMS. A change in the capacitance was detected when the electrode was touched. Such a touch-sensitive device was used to fabricate an organic electronic textile keyboard13, as shown in Figure 3d.
<img alt="Figure 1" src="/files/ftp_upload/55439/55439fig1.jpg" /> 
Figure 1. Process flow illustrating the patterning of conducting polymers onto textiles. Process flow illustrating the patterning of conducting polymers onto textiles. Step 1: mask preparation; step 2: PDMS deposition on the polyimide patterning mask defining the outline of the desired design; step 3: transfer of the masking layer by the placement of the textile on the PDMS-coated mask; step 4: transfer of the PDMS into the bulk of the textile, step 5: deposition of conducting polymer solution onto unprotected textile. The pictures on the right show the results of the key steps of the process flow. <a href="http://cloudflare.jove.com/files/ftp_upload/55439/55439fig1large.jpg" target="_blank">Please click here to view a larger version of this figure.</a>
<img alt="Figure 2" src="/files/ftp_upload/55439/55439fig2.jpg" /> 
Figure 2. Two examples of fabrication organic devices. Two examples of fabrication organic devices. a) Ionic liquid gel coating on the PEDOT:PSS textile electrode for cutaneous sensing. b) Insulation layer deposition on the PEDOT:PSS textile electrode for touch sensors. <a href="http://cloudflare.jove.com/files/ftp_upload/55439/55439fig2large.jpg" target="_blank">Please click here to view a larger version of this figure.</a>
<img alt="Figure 3" src="/files/ftp_upload/55439/55439fig3.jpg" /> 
Figure 3. Photographs of organic electronic textile devices. a) PEDOT:PSS electrodes for stretch sensing. b) Array of OECT transistors for wearable biosensing. c) Circular PEDOT:PSS electrode coated with an ionic liquid gel for cutaneous electrophysiology. d) Organic touch sensors for a wearable keyboard. <a href="http://cloudflare.jove.com/files/ftp_upload/55439/55439fig3large.jpg" target="_blank">Please click here to view a larger version of this figure.</a>

Watch this Scientific Journal Video about A Simple and Scalable Fabrication Method for Organic Electronic Devices on Textiles at JoVE.com

A Simple and Scalable Fabrication Method for Organic Electronic Devices on Textiles

In this paper, we present a protocol to selectively deposit organic materials on textiles, which allows for the direct integration of organic electronic devices with wearables. The fabricated devices can be fully integrated in textiles, respecting their mechanical appearance and enabling sensing capabilities.

Today, wearable electronics devices combine a large variety of functional, stretchable, and flexible technologies. However, in many cases, these devices ...

a-simple-scalable-fabrication-method-for-organic-electronic-devices

Universidad San Sebastian

Research

JoVE Journal

Bioengineering

10.3K Views.  Ecole Nationale Supérieure des Mines de Saint-Étienne, CMP-EMSE, MOC. In this paper, we present a protocol to selectively deposit organic materials on textiles, which allows for the direct integration of organic electronic devices with wearables. The fabricated devices can be fully integrated in textiles, respecting their mechanical appearance and enabling sensing capabilities.

Video: A Simple and Scalable Fabrication Method for Organic Electronic Devices on Textiles

Simple Microfluidic Devices for in vivo Imaging of C. elegans, Drosophila and Zebrafish

Micro fabricated fluidic devices provide an accessible micro-environment for in vivo studies on small organisms. Simple fabrication processes are available for microfluidic devices using soft lithography techniques 1-3. Microfluidic devices have been used for sub-cellular imaging 4,5, in vivo laser microsurgery 2,6 and cellular imaging 4,7. In vivo imaging requires immobilization of organisms. This has been achieved using suction 5,8, tapered channels 6,7,9, deformable membranes 2-4,10, suction with additional cooling 5, anesthetic gas 11, temperature sensitive gels 12, cyanoacrylate glue 13 and anesthetics such as levamisole 14,15. Commonly used anesthetics influence synaptic transmission 16,17 and are known to have detrimental effects on sub-cellular neuronal transport 4. In this study we demonstrate a membrane based poly-dimethyl-siloxane (PDMS) device that allows anesthetic free immobilization of intact genetic model organisms such as Caenorhabditis elegans (C. elegans), Drosophila larvae and zebrafish larvae. These model organisms are suitable for in vivo studies in microfluidic devices because of their small diameters and optically transparent or translucent bodies. Body diameters range from ~10 &mu;m to ~800 &mu;m for early larval stages of C. elegans and zebrafish larvae and require microfluidic devices of different sizes to achieve complete immobilization for high resolution time-lapse imaging. These organisms are immobilized using pressure applied by compressed nitrogen gas through a liquid column and imaged using an inverted microscope. Animals released from the trap return to normal locomotion within 10 min.
We demonstrate four applications of time-lapse imaging in C. elegans namely, imaging mitochondrial transport in neurons, pre-synaptic vesicle transport in a transport-defective mutant, glutamate receptor transport and Q neuroblast cell division. Data obtained from such movies show that microfluidic immobilization is a useful and accurate means of acquiring in vivo data of cellular and sub-cellular events when compared to anesthetized animals (Figure 1J and 3C-F 4).
Device dimensions were altered to allow time-lapse imaging of different stages of C. elegans, first instar Drosophila larvae and zebrafish larvae. Transport of vesicles marked with synaptotagmin tagged with GFP (syt.eGFP) in sensory neurons shows directed motion of synaptic vesicle markers expressed in cholinergic sensory neurons in intact first instar Drosophila larvae. A similar device has been used to carry out time-lapse imaging of heartbeat in ~30 hr post fertilization (hpf) zebrafish larvae. These data show that the simple devices we have developed can be applied to a variety of model systems to study several cell biological and developmental phenomena in vivo.

Simple Microfluidic Devices for in vivo Imaging of C. elegans, Drosophila and Zebrafish

A simple microfluidic device has been developed to perform anesthetic free in vivo imaging of C. elegans, intact Drosophila larvae and zebrafish larvae. The device utilizes a deformable PDMS membrane to immobilize these model organisms in order to perform time lapse imaging of numerous processes such as heart beat, cell division and sub-cellular neuronal transport. We demonstrate the use of this device and show examples of different types of data collected from different model systems.

Micro fabricated fluidic devices provide an accessible micro-environment for in vivo studies on small organisms. Simple fabrication processes are ...

Microfluidic Chip Fabrication and Method to Detect Influenza

Fast and effective diagnostics play an important role in controlling infectious disease by enabling effective patient management and treatment. Here, we present an integrated microfluidic thermoplastic chip with the ability to amplify influenza A virus in patient nasopharyngeal (NP) swabs and aspirates. Upon loading the patient sample, the microfluidic device sequentially carries out on-chip cell lysis, RNA purification and concentration steps within the solid phase extraction (SPE), reverse transcription (RT) and polymerase chain reaction (PCR) in RT-PCR chambers, respectively. End-point detection is performed using an off-chip Bioanalyzer (Agilent Technologies, Santa Clara, CA). For peripherals, we used a single syringe pump to drive reagent and samples, while two thin film heaters were used as the heat sources for the RT and PCR chambers. The chip is designed to be single layer and suitable for high throughput manufacturing to reduce the fabrication time and cost. The microfluidic chip provides a platform to analyze a wide variety of virus and bacteria, limited only by changes in reagent design needed to detect new pathogens of interest.

An integrated microfluidic thermoplastic chip has been developed for use as a molecular diagnostic. The chip performs nucleic acid extraction, reverse transcriptase, and PCR. Methods for fabricating and running the chip are described.

Fast  and effective diagnostics play an important role in controlling infectious  disease by enabling effective patient management and treatment. Here, we ...

Fabrication of a Functionalized Magnetic Bacterial Nanocellulose with Iron Oxide Nanoparticles

In this study, bacterial nanocellulose (BNC) produced by the bacteria Gluconacetobacter xylinus is synthesized and impregnated in situ with iron oxide nanoparticles (IONP) (Fe3O4) to yield a magnetic bacterial nanocellulose (MBNC). The synthesis of MBNC is a precise and specifically designed multi-step process. Briefly, bacterial nanocellulose (BNC) pellicles are formed from preserved G. xylinus strain according to our experimental requirements of size and morphology. A solution of iron(III) chloride hexahydrate (FeCl3&#183;6H2O) and iron(II) chloride tetrahydrate (FeCl2&#183;4H2O) with a 2:1 molar ratio is prepared and diluted in deoxygenated high purity water. A BNC pellicle is then introduced in the vessel with the reactants. This mixture is stirred and heated at 80 &#176;C in a silicon oil bath and ammonium hydroxide (14%) is then added by dropping to precipitate the ferrous ions into the BNC mesh. This last step allows forming in situ magnetite nanoparticles (Fe3O4) inside the bacterial nanocellulose mesh to confer magnetic properties to BNC pellicle. A toxicological assay was used to evaluate the biocompatibility of the BNC-IONP pellicle. Polyethylene glycol (PEG) was used to cover the IONPs in order to improve their biocompatibility. Scanning electron microscopy (SEM) images showed that the IONP were located preferentially in the fibril interlacing spaces of the BNC matrix, but some of them were also found along the BNC ribbons. Magnetic force microscope measurements performed on the MBNC detected the presence magnetic domains with high and weak intensity magnetic field, confirming the magnetic nature of the MBNC pellicle. Young&#39;s modulus values obtained in this work are also in a reasonable agreement with those reported for several blood vessels in previous studies.

Here, we present a protocol to make a bacterial nanocellulose (BNC) magnetic for applications in damaged blood vessel reconstruction. The BNC was synthesized by G. xylinus strain. On the other hand, magnetization of the BNC was realized through in situ precipitation of Fe2+ and Fe3+ ferrous ions inside the BNC mesh.

In this study, bacterial nanocellulose (BNC) produced by the bacteria Gluconacetobacter xylinus is synthesized and impregnated in situ with iron oxide ...

Fabrication and Characterization of a Conformal Skin-like Electronic System for Quantitative, Cutaneous Wound Management

Recent advances in the development of electronic technologies and biomedical devices offer opportunities for non-invasive, quantitative assessment of cutaneous wound healing on the skin. Existing methods, however, still rely on visual inspections through various microscopic tools and devices that normally include high-cost, sophisticated systems and require well trained personnel for operation and data analysis. Here, we describe methods and protocols to fabricate a conformal, skin-like electronics system that enables conformal lamination to the skin surface near the wound tissues, which provides recording of high fidelity electrical signals such as skin temperature and thermal conductivity. The methods of device fabrication provide details of step-by-step preparation of the microelectronic system that is completely enclosed with elastomeric silicone materials to offer electrical isolation. The experimental study presents multifunctional, biocompatible, waterproof, reusable, and flexible/stretchable characteristics of the device for clinical applications. Protocols of clinical testing provide an overview and sequential process of cleaning, testing setup, system operation, and data acquisition with the skin-like electronics, gently mounted on hypersensitive, cutaneous wound and contralateral tissues on patients.

This article presents methods to fabricate and characterize a conformal, skin-like electronic system and protocols for the use in clinical applications, particularly on cutaneous wound management.

Recent advances in the development of electronic technologies and biomedical devices offer opportunities for non-invasive, quantitative assessment of ...

Biomembrane Fabrication by the Solvent-assisted Lipid Bilayer (SALB) Method

In order to mimic cell membranes, the supported lipid bilayer (SLB) is an attractive platform which enables in vitro investigation of membrane-related processes while conferring biocompatibility and biofunctionality to solid substrates. The spontaneous adsorption and rupture of phospholipid vesicles is the most commonly used method to form SLBs. However, under physiological conditions, vesicle fusion (VF) is limited to only a subset of lipid compositions and solid supports. Here, we describe a one-step general procedure called the solvent-assisted lipid bilayer (SALB) formation method in order to form SLBs which does not require vesicles. The SALB method involves the deposition of lipid molecules onto a solid surface in the presence of water-miscible organic solvents (e.g., isopropanol) and subsequent solvent-exchange with aqueous buffer solution in order to trigger SLB formation. The continuous solvent exchange step enables application of the method in a flow-through configuration suitable for monitoring bilayer formation and subsequent alterations using a wide range of surface-sensitive biosensors. The SALB method can be used to fabricate SLBs on a wide range of hydrophilic solid surfaces, including those which are intractable to vesicle fusion. In addition, it enables fabrication of SLBs composed of lipid compositions which cannot be prepared using the vesicle fusion method. Herein, we compare results obtained with the SALB and conventional vesicle fusion methods on two illustrative hydrophilic surfaces, silicon dioxide and gold. To optimize the experimental conditions for preparation of high quality bilayers prepared via the SALB method, the effect of various parameters, including the type of organic solvent in the deposition step, the rate of solvent exchange, and the lipid concentration is discussed along with troubleshooting tips. Formation of supported membranes containing high fractions of cholesterol is also demonstrated with the SALB method, highlighting the technical capabilities of the SALB technique for a wide range of membrane configurations.

Biomembrane Fabrication by the Solvent-assisted Lipid Bilayer SALB Method

We present an experimental protocol to form a supported lipid bilayer on solid substrates without using lipid vesicles. We demonstrate a one-step method to form a lipid bilayer on silicon dioxide and gold as well as supported membranes with cholesterol-enriched domain for various biological applications.

In order to mimic cell membranes, the supported lipid bilayer (SLB) is an attractive platform which enables in vitro investigation of membrane-related ...

Rod-based Fabrication of Customizable Soft Robotic Pneumatic Gripper Devices for Delicate Tissue Manipulation

Soft compliant gripping is essential in delicate surgical manipulation for minimizing the risk of tissue grip damage caused by high stress concentrations at the point of contact. It can be achieved by complementing traditional rigid grippers with soft robotic pneumatic gripper devices. This manuscript describes a rod-based approach that combined both 3D-printing and a modified soft lithography technique to fabricate the soft pneumatic gripper. In brief, the pneumatic featureless mold with chamber component is 3D-printed and the rods were used to create the pneumatic channels that connect to the chamber. This protocol eliminates the risk of channels occluding during the sealing process and the need for external air source or related control circuit. The soft gripper consists of a chamber filled with air, and one or more gripper arms with a pneumatic channel in each arm connected to the chamber. The pneumatic channel is positioned close to the outer wall to create different stiffness in the gripper arm. Upon compression of the chamber which generates pressure on the pneumatic channel, the gripper arm will bend inward to form a close grip posture because the outer wall area is more compliant. The soft gripper can be inserted into a 3D-printed handling tool with two different control modes for chamber compression: manual gripper mode with a movable piston, and robotic gripper mode with a linear actuator. The double-arm gripper with two actuatable arms was able to pick up objects of sizes up to 2 mm and yet generate lower compressive forces as compared to elastomer-coated and non-coated rigid grippers. The feasibility of having other designs, such as single-arm or hook gripper, was also demonstrated, which further highlighted the customizability of the soft gripper device, and it's potential to be used in delicate surgical manipulation to reduce the risk of tissue grip damage.

This protocol describes a rod-based approach, combining 3D-printing and soft lithography techniques for fabricating the soft gripper devices. This approach eliminates the need for an external air source by incorporating a chamber component and reduces the chance of occlusion during the sealing process, particularly for miniaturized pneumatic channels.

Soft compliant gripping is essential in delicate surgical manipulation for minimizing the risk of tissue grip damage caused by high stress concentrations ...

Wet Chemistry and Peptide Immobilization on Polytetrafluoroethylene for Improved Cell-adhesion

Endowing materials surface with cell-adhesive properties is a common strategy in biomaterial research and tissue engineering. This is particularly interesting for already approved polymers that have a long standing use in medicine because these materials are well characterized and legal issues associated with the introduction of newly synthesized polymers may be avoided. Polytetrafluoroethylene (PTFE) is one of the most frequently employed materials for the manufacturing of vascular grafts but the polymer lacks cell adhesion promoting features. Endothelialization, i.e., complete coverage of the grafts inner surface with a confluent layer of endothelial cells is regarded key to optimal performance, mainly by reducing thrombogenicity of the artificial interface.
This study investigates the growth of endothelial cells on peptide-modified PTFE and compares these results to those obtained on unmodified substrate. Coupling with the endothelial cell adhesive peptide Arg-Glu-Asp-Val (REDV) is performed via activation of the fluorin-containing polymer using the reagent sodium naphthalenide, followed by subsequent conjugation steps. Cell culture is accomplished using Human Umbilical Vein Endothelial Cells (HUVECs) and excellent cellular growth on peptide-immobilized material is demonstrated over a two-week period.

Cell-adhesiveness is key to many approaches in biomaterial research and tissue engineering. A step-by-step technique is presented using wet-chemistry for the surface modification of the important polymer PTFE with peptides.

Endowing materials surface with cell-adhesive properties is a common strategy in biomaterial research and tissue engineering. This is particularly ...

Fabrication of Three-dimensional Paper-based Microfluidic Devices for Immunoassays

Paper wicks fluids autonomously due to capillary action. By patterning paper with hydrophobic barriers, the transport of fluids can be controlled and directed within a layer of paper. Moreover, stacking multiple layers of patterned paper creates sophisticated three-dimensional microfluidic networks that can support the development of analytical and bioanalytical assays. Paper-based microfluidic devices are inexpensive, portable, easy to use, and require no external equipment to operate. As a result, they hold great promise as a platform for point-of-care diagnostics. In order to properly evaluate the utility and analytical performance of paper-based devices, suitable methods must be developed to ensure their manufacture is reproducible and at a scale that is appropriate for laboratory settings. In this manuscript, a method to fabricate a general device architecture that can be used for paper-based immunoassays is described. We use a form of additive manufacturing (multi-layer lamination) to prepare devices that comprise multiple layers of patterned paper and patterned adhesive. In addition to demonstrating the proper use of these three-dimensional paper-based microfluidic devices with an immunoassay for human chorionic gonadotropin (hCG), errors in the manufacturing process that may result in device failures are discussed. We expect this approach to manufacturing paper-based devices will find broad utility in the development of analytical applications designed specifically for limited-resource settings.

We detail a method to fabricate three-dimensional paper-based microfluidic devices for use in the development of immunoassays. Our approach to device assembly is a type of multilayer, additive manufacturing. We demonstrate a sandwich immunoassay to provide representative results for these types of paper-based devices.

Paper wicks fluids autonomously due to capillary action. By patterning paper with hydrophobic barriers, the transport of fluids can be controlled and ...

Microfluidic Platform with Multiplexed Electronic Detection for Spatial Tracking of Particles

Microfluidic processing of biological samples typically involves differential manipulations of suspended particles under various force fields in order to spatially fractionate the sample based on a biological property of interest. For the resultant spatial distribution to be used as the assay readout, microfluidic devices are often subjected to microscopic analysis requiring complex instrumentation with higher cost and reduced portability. To address this limitation, we have developed an integrated electronic sensing technology for multiplexed detection of particles at different locations on a microfluidic chip. Our technology, called Microfluidic CODES, combines Resistive Pulse Sensing with Code Division Multiple Access to compress 2D spatial information into a 1D electrical signal. In this paper, we present a practical demonstration of the Microfluidic CODES technology to detect and size cultured cancer cells distributed over multiple microfluidic channels. As validated by the high-speed microscopy, our technology can accurately analyze dense cell populations all electronically without the need for an external instrument. As such, the Microfluidic CODES can potentially enable low-cost integrated lab-on-a-chip devices that are well suited for the point-of-care testing of biological samples.

We demonstrate a microfluidic platform with an integrated surface electrode network that combines resistive pulse sensing (RPS) with code division multiple access (CDMA), to multiplex detection and sizing of particles in multiple microfluidic channels.

Microfluidic processing of biological samples typically involves differential manipulations of suspended particles under various force fields in order to ...

Fabrication and Validation of an Organ-on-chip System with Integrated Electrodes to Directly Quantify Transendothelial Electrical Resistance

Organs-on-chips, in vitro models involving the culture of (human) tissues inside microfluidic devices, are rapidly emerging and promise to provide useful research tools for studying human health and disease. To characterize the barrier function of cell layers cultured inside organ-on-chip devices, often transendothelial or transepithelial electrical resistance (TEER) is measured. To this end, electrodes are usually integrated into the chip by micromachining methods to provide more stable measurements than is achieved with manual insertion of electrodes into the inlets of the chip. However, these electrodes frequently hamper visual inspection of the studied cell layer or require expensive cleanroom processes for fabrication. To overcome these limitations, the device described here contains four easily integrated electrodes that are placed and fixed outside of the culture area, making visual inspection possible. Using these four electrodes the resistance of six measurement paths can be quantified, from which the TEER can be directly isolated, independent of the resistance of culture medium-filled microchannels. The blood-brain barrier was replicated in this device and its TEER was monitored to show the device applicability. This chip, the integrated electrodes and the TEER determination method are generally applicable in organs-on-chips, both to mimic other organs or to be incorporated into existing organ-on-chip systems.

This publication describes the fabrication of an organ-on-chip device with integrated electrodes for direct quantification of transendothelial electrical resistance (TEER). For validation, the blood-brain barrier was mimicked inside this microfluidic device and its barrier function was monitored. The presented methods for electrode integration and direct TEER quantification are generally applicable.

Organs-on-chips, in vitro models involving the culture of (human) tissues inside microfluidic devices, are rapidly emerging and promise to provide useful ...