我们感谢在康奈尔大学教授保罗·麦克尤恩早期讨论。这项工作是支持的启动基金提供的密歇根大学和美国国家科学基金会可扩展的纳米制造计划（DMR-1120187）。这项工作在美国密歇根大学，由美国国家科学基金会资助的美国国家纳米技术基础设施网络的成员劳瑞加工设施。

1。单壁碳纳米管生长催化剂图案<ol><li>首先，用低压化学气相沉积（CVD）生长为500nm的Si 3 N 4 / 500nm的SiO 2膜的顶部的硅晶片。 </li><li>旋涂一层光致抗蚀剂（PR），以500rpm，持续5秒，然后4,000 rpm离心40秒。 </li><li>晶片在115℃下烘烤90秒。 </li><li>与催化剂（ 图1）的矩形的凹坑中使用的光掩模和晶片暴露在UV（365 nm）的辐照度为300毫焦耳/厘米2，持续0.3秒。曝光后晶片在115℃下烘烤90秒。 </li></ol> 提示:5微米×5微米，10微米×5微米等设计不同大小的坑， 如考虑到单壁碳纳米管的化学气相沉积（CVD）的生长过程中变异。 <ol start="5"><li>开发晶片开发商70秒轻轻晃动晶片经由过程。 </li><li> DI年末冲洗晶圆r为2分钟，然后用氮气（N 2）枪吹干。 </li><li>开发晶片加载到电子束蒸发器腔和存款0.5 nm的铁（Fe）的上面的腔室的压力为10 -6托。 </li><li>切块的晶圆切割成较小的模具1.5厘米×1.5厘米。 </li><li>将光致抗蚀剂通过浸渍在温水中丙酮和异丙醇（IPA），每次10分钟。这留下铁催化剂纳米管的生长的长方形坑。 </li></ol> 2。 CVD生长碳纳米管<ol><li>将催化剂涂覆的模具的CVD生长炉中的石英管。 </li></ol> 提示:确定纳米管的生长的甜蜜点。增长是均匀的面积为2英寸x 2英寸下游炉（ 图2c）。 <ol start="2"><li>基板在空气中在880℃下退火一小时，以除去光致抗蚀剂残余物（ 图2a）。允许降温。 </li><li>用Ar吹扫室坤3的SLM（标准升每分钟）的5分钟。 </li><li>斜坡上升至800℃的炉中，在管的中心，同时保持1 SLM的氩气（ 图2b）的流动。 </li><li>流0.2 SLM的氢气为5分钟，以减少的催化剂粒子， 即，将铁的氧化物铁。 </li><li>介绍5.5的SCCM（标准立方厘米每分钟），乙烯（C 2 H 4）35分钟增长碳纳米管。保持H 2流量0.2 SLM整个过程。从这个配方获得单壁碳纳米管的长度是20微米（微米）。 </li><li>允许炉冷却至室温下，用一个小的氩气流量。 </li></ol> 3。单壁碳纳米管场效应晶体管制造<ol><li>设计的光掩模（ 图1）定义的源极-漏极电极的电流-电压（IV）特性的碳纳米管。 </li></ol> 提示:延长电极接触垫远APAR吨模具，使他们仍然可以访问后甚至放下一个微流体邮票上的活跃碳纳米管区域。 <ol start="2"><li>按照步骤1.2 -1.6定义接触面积金属沉积。 </li><li>存款的Ti /金0.5 nm/50的纳米电子束蒸发室在10 -6托的源漏接触。 </li><li>让丙酮过夜金属升空的模具。升空后，浸在IPA模具10分钟，然后吹干，用N 2枪。 </li><li>不要毯子为500nm的电子束蒸发沉积SiO 2的栅极电介质，在10 -6托。 </li><li>设计的光掩模（ 图1）限定的栅极电极。 </li><li>按照步骤1.2-1.6定义区域栅极金属沉积。 </li><li>蒸发50 nm/50的Cr / Au纳米电子束蒸发器的顶部栅电极。按照步骤3.4金属升空。 </li></ol> 提示:使用粗壮的镀铬层，以增加强度Øf暂停顶栅。门的尺寸也悬挂成功的关键。 <ol start="9"><li>毯上沉积一层薄的SiO 2（20纳米）的顶部电极钝化。 </li><li>光罩设计（ 图1），打开一个战壕在SiO 2，访问碳纳米管通道。焊垫设计在相同的掩模蚀刻区域打开远离单壁碳纳米管的通道访问源，漏和栅电极的区域。 </li><li>按照步骤1.2-1.6图案公关开辟沟渠湿蚀刻的SiO 2。 </li><li>湿法腐蚀的蒸发SiO 2的使用1:20的BHF溶液在3分钟和30秒。的Si 3 N 4提供了一个蚀刻停止层，以防止进一步渗透的BHF。 </li></ol> 提示:建议蚀刻时间校准。 <ol start="13"><li>在去离子水中彻底冲洗装置，然后在IPA中浸入5分钟。使用N 2枪吹干。该设备STructure由于粗壮的镀铬层保持完好。 </li></ol> 4。化学修饰碳纳米管侧壁<ol><li>准备6毫1 pyrenebutanoic的酸琥珀酰亚胺酯（PBSE）在二甲基甲酰胺（DMF）中的溶液中。 </li><li>孵育单壁碳纳米管FET在这个链接的分子溶液在室温下1小时死亡。 </li><li>彻底清洗模具，在DMF洗去过量的试剂。吹干设备。 </li><li>准备一个20mg/ml的溶液中生物素-3，6 - （生物素PEO胺-BPA）在去离子水中的生物素化的单壁碳纳米管dioxaoctanediamine。 </li><li>电路小片在该溶液中孵育18小时后，模具在去离子水中彻底冲洗并吹干。重视BPA的PBSE连接分子。 </li><li>准备1mg/ml的链霉亲和素溶液pH值在7.2 PBS溶液链霉亲和素结合。 </li><li>对于静态测量，模具在链霉抗生物素蛋白溶液中孵育20分钟，完全官能化的生物素化的单壁碳纳米管。索罗ughly冲洗并吹干模具。实时检测，杜绝PDMS流道（步骤6）然后流的链霉亲和素溶液中离子强度高的背景链霉亲和素结合（ 图3）。 </li></ol> 注:我们冲洗模具，在模具使用挤压瓶胶DI水（50毫升）。然后，我们的电路小片移动到另一培养皿中含有去离子水，1分钟左右，电路小片移动。我们重复这两个步骤总的8-10倍。 5。制备的聚二甲基硅氧烷（PDMS）流体腔模具<ol><li>在一个称量杯，倒入9份（重量）的PDMS单体，并添加1份（重量）的固化剂，充分混合两个。 </li><li>脱气的混合物25分钟，在干燥器中。气泡会通过混合物上升而离开。 </li></ol> 提示:如果混合物开始发泡，发泄室，让它沉淀下来的几个秒钟，然后再次脱气。 <ol start="3"><li>在培养皿中放置一个新的硅片。脱气后的PDMS混合物倒入在它的上面有一个PDMS层晶片上方5mm。 </li><li>陪替氏培养皿在70℃的烘箱中放置1小时。 </li><li>取出培养皿中，让它冷却下来。使用手术刀切出一块长方形的PDMS，将其拉出使用镊子。 </li></ol> 提示:PDMS侧直接与硅片接触 ​​是干净的，非常平坦。此方将单壁碳纳米管场效应管模具接触。要小心，不要污染。 <ol start="6"><li>上下颠倒放置的矩形模具，并在使用活检冲床（3毫米直径）从平到另一侧钻了一个洞。这将确保没有粗糙的边缘上侧扁PDMS（ 图4a）。 </li><li>的PDMS室放置在设备顶部的仔细对齐，它的有源区上方的所制作的单壁碳纳米管场效应管模具（&lt;强大的图4a，b）条。轻轻点击它，以确保模具邮票。侧扁债券到模具，，提供防漏室。 </li></ol> 提示:这可以通过用裸眼或用光学显微镜有足够的工作空间。 PDMS如果不沾（一般如果电路小片和/或PDMS印章不干净的），氧等离子体（20瓦，15秒），以协助对PDMS粘接。使用血浆权力高于这会导致更强的粘接，但是，我们已经看到翻录电极，而在这种情况下删除的PDMS。 <ol start="8"><li>取出后PDMS印章电气测试和之前下化学官能一步的模具浸入去离子水，轻轻抬起邮票。 </li></ol> 6。制备微流通道<ol><li>取一干净的硅晶片，并将其放置在电热板上，在200℃下进行5分钟以除去任何湿气。 </li><li>旋涂SU-8 2015 500转（100转/秒斜坡率），持续5秒，然后1,250 rpm离心30秒，在300转/秒的斜坡率。这给出了30μm厚的SU-8层上的硅晶片。 </li><li>软烤晶圆在95℃下5分钟。 </li><li>设计的光掩模（ 图4c）定义的300μm宽的流路图案的单壁碳纳米管的面积之上。 </li></ol> 提示:为避免结构倒塌的通道宽度:，足够高度比为10:1（300微米:在这种情况下，30微米）。 <ol start="5"><li>使用365纳米的紫外线光刻法定义的流路，用紫外曝光时间为0.9秒。 </li><li>烘烤后的电路小片在95℃下5分钟。 </li><li>发展伴随着轻轻摇动5分钟的SU-8开发的模式。 </li><li>冲洗晶圆IPA和N 2枪吹干。 </li><li>按照步骤5.1-5.2准备PDMS混合物。 </li><li>随着硅烷化剂TRIC 2-3下拉晶片放置在干燥器中的SU-8的模具hloro（3，3，3 - 三氟丙基）硅烷在培养皿中。打开真空泵在真空中1小时让晶片坐。 </li><li>脱气后的PDMS混合物倒入到晶片上，并在70℃的烘箱中加热1小时。 </li><li>切PDMS模具（SU-8负）使用手术刀。 </li><li>将PDMS印章倒过来，用活检冲头（0.75毫米直径）钻一个孔，在每一端的流道。确保从平面侧（一侧与硅片接触 ​​）到其他钻孔，以避免粗糙的边缘（ 图4e）。 </li><li> PDMS流室放置在设备顶部的仔细对准顶部制作的单壁碳纳米管的场效应管的模具，在显微镜下的有源区上。轻轻点击它，以确保模具邮票。侧扁债券到模具，，提供防漏室。 （ 图4c和4d） </li><li>推入聚乙烯管插入孔中并将 ​​另一端连接到流体的源极和漏极注射器（FIGURE 4E）。 </li><li>将注射器的注射泵系统，以保持控制，通过该信道（ 图6）的流体流动。 </li></ol> 7。直流电气测量设置的<ol><li>连接电压数据采集卡的端口的单壁碳纳米管场效应管的源极和栅极接触。 </li><li>通过电流的前置放大器，数据采集卡输入端口连接漏极接触。 </li><li>应用30毫伏（MV）源接触，扫的栅极电压，并记录电流漏（ 图5a）。 </li></ol> 提示:测量溶液中，保持栅极电压扫描参数| 0.7伏以内，以避免漏电和金属栅电极和解决方案之间的反应。 8。 AC电气测量设置<ol><li>来自锁相放大器的外部调制信号的端口上的频率发生器，用于设置连接的Ref信号AM调制freque的NCY输出。 </li><li>连接的单壁碳纳米管的场效应管的源极接触的频率发生器和直流电压从数据采集卡的使用偏置T形（ 图5b和5c）的AM调制的RF输出端口。 </li><li>将门接触电压端口的数据采集卡。 </li><li>漏极接触连接到锁相放大器读交流电流通过碳纳米管。通过数据采集输入端口读取的电流的幅度和相位。 </li><li>源的直流电压保持在0伏和AM信号的频率在200千赫（kHz）。 </li><li>扫栅极电压和测量的漏电流。 </li><li>我在不同的频率组合V G扫描频率增加，并重复步骤8.6。 </li></ol> 9。电气测量解决方案（无流量） <ol><li>准备:1 mM氯化钠，氯化钠10毫米和100毫米氯化钠盐溶液从5M的NaCl原液。 </li><li>以SWNT设备从步骤3.13。步骤4.1-4.5获得biotinylate的开展D设备。 </li><li>按照步骤5把PDMS室，设备顶部的。 </li><li>填充室DI水用吸管。 </li><li>按照步骤8混频测量不同频率。 </li><li>重复9.4三种不同的盐溶液，10 mM氯化钠，氯化钠100毫米1毫米氯化钠。 </li></ol> 提示:使用吸管撤回先前的解决方案，然后冲洗室多次与新的解决方案。随时切换，从低到高浓度的解决方案。 <ol start="7"><li>将PDMS印章用镊子轻轻提起邮票。 </li><li>用去离子水彻底冲洗设备。 </li><li>开展链霉亲和素结合在4.6-4.7解释。 </li><li>重复步骤9.3-9.6。 </li></ol> 10。电气测量解决方案（实时流） <ol><li>准备100 mM氯化钠盐溶液从5 M氯化钠股票解决方案开始。 </li><li>单壁碳纳米管的设备从STEp 3.13。开展步骤4.1-4.5获得生物素标记的设备。 </li><li>将以下的步骤（6）的移动设备上的微流体流动通道中..撤回模式操作的微流体通道的一端连接一个空的注射器。在其另一端连接注射器100mM NaCl的溶液中与背景。 </li><li>电气测量的详细步骤8。固定的频率（F = 10兆赫）和栅极偏置电压（V = 0）。 </li><li>开始注射泵在停药模式（流速= 0.4毫升-1小时），并监测随时间的电流信号。在100mM NaCl的链霉抗生物素蛋白-生物素的结合（ 图6）和监控信号的变化，开关1mg/ml的链霉抗生物素蛋白的溶液。 </li></ol>

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A., Dekker, C. <a target="_blank" href="http://www.ncbi.nlm.nih.gov/entrez/query.fcgi?db=PubMed&cmd=Search&doptcmdl=Citation&defaultField=Title+Word&term=Enzyme-coated+carbon+nanotubes+as+single-molecule+biosensors.">Enzyme-coated carbon nanotubes as single-molecule biosensors.</a> Nano Letters. 3, 727-730 (2003).</li><li>Snow, E. S., Perkins, F. K., Houser, E. J., Badescu, S. C., Reinecke, T. L. <a target="_blank" href="http://www.ncbi.nlm.nih.gov/entrez/query.fcgi?db=PubMed&cmd=Search&doptcmdl=Citation&defaultField=Title+Word&term=Chemical+detection+with+a+single-walled+carbon+nanotube+capacitor.">Chemical detection with a single-walled carbon nanotube capacitor.</a> Science. 307, 1942-1945 (2005).</li><li>Kong, J., Franklin, N. R., et al. <a target="_blank" href="http://www.ncbi.nlm.nih.gov/entrez/query.fcgi?db=PubMed&cmd=Search&doptcmdl=Citation&defaultField=Title+Word&term=Nanotube+molecular+wires+as+chemical+sensors.">Nanotube molecular wires as chemical sensors.</a> Science. 287, 622-625 (2000).</li><li>Star, A., Tu, E., et al. <a target="_blank" href="http://www.ncbi.nlm.nih.gov/entrez/query.fcgi?db=PubMed&cmd=Search&doptcmdl=Citation&defaultField=Title+Word&term=Label-free+detection+of+DNA+hybridization+using+carbon+nanotube+network+field-effect+transistors.">Label-free detection of DNA hybridization using carbon nanotube network field-effect transistors.</a> Proceedings of the National Academy of Sciences of the United States of America. 103, 921-926 (2006).</li><li>Patolsky, F., Zheng, G. F., Lieber, C. M. <a target="_blank" href="http://www.ncbi.nlm.nih.gov/entrez/query.fcgi?db=PubMed&cmd=Search&doptcmdl=Citation&defaultField=Title+Word&term=Nanowire-based+biosensors.">Nanowire-based biosensors.</a> Analytical Chemistry. 78, 4260-4269 (2006).</li><li>Stern, E., Klemic, J. F., et al. <a target="_blank" href="http://www.ncbi.nlm.nih.gov/entrez/query.fcgi?db=PubMed&cmd=Search&doptcmdl=Citation&defaultField=Title+Word&term=Label-free+immunodetection+with+CMOS-compatible+semiconducting+nanowires.">Label-free immunodetection with CMOS-compatible semiconducting nanowires.</a> Nature. 445, 519-522 (2007).</li><li>Krivitsky, V., Hsiung, L. -. C., et al. <a target="_blank" href="http://www.ncbi.nlm.nih.gov/entrez/query.fcgi?db=PubMed&cmd=Search&doptcmdl=Citation&defaultField=Title+Word&term=Nanowires+Forest-Based+On-Chip+Biomolecular+Filtering,+Separation+and+Preconcentration+Devices:+Nanowires+Do+it+All.">Nanowires Forest-Based On-Chip Biomolecular Filtering, Separation and Preconcentration Devices: Nanowires Do it All.</a> Nano Letters. 12, 4748-4756 (2012).</li><li>Kong, J., Soh, H. T., Cassell, A. M., Quate, C. F., Dai, H. J. <a target="_blank" href="http://www.ncbi.nlm.nih.gov/entrez/query.fcgi?db=PubMed&cmd=Search&doptcmdl=Citation&defaultField=Title+Word&term=Synthesis+of+individual+single-walled+carbon+nanotubes+on+patterned+silicon+wafers.">Synthesis of individual single-walled carbon nanotubes on patterned silicon wafers.</a> Nature. 395, 878-881 (1998).</li><li>Liu, G., Velasco, J., Bao, W. Z., Lau, C. N. <a target="_blank" href="http://www.ncbi.nlm.nih.gov/entrez/query.fcgi?db=PubMed&cmd=Search&doptcmdl=Citation&defaultField=Title+Word&term=Fabrication+of+graphene+p-n-p+junctions+with+contactless+top+gates.">Fabrication of graphene p-n-p junctions with contactless top gates.</a> Applied Physics Letters. 92, (2008).</li><li>Katz, E., Willner, I. <a target="_blank" href="http://www.ncbi.nlm.nih.gov/entrez/query.fcgi?db=PubMed&cmd=Search&doptcmdl=Citation&defaultField=Title+Word&term=Probing+biomolecular+interactions+at+conductive+and+semiconductive+surfaces+by+impedance+spectroscopy:+Routes+to+impedimetric+immunosensors,+DNA-Sensors,+and+enzyme+biosensors.">Probing biomolecular interactions at conductive and semiconductive surfaces by impedance spectroscopy: Routes to impedimetric immunosensors, DNA-Sensors, and enzyme biosensors.</a> Electroanalysis. 15, 913-947 (2003).</li><li>K'Owino, I. O., Sadik, O. 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作者宣称，他们有没有竞争经济利益。

碳纳米管的生长，不仅取决于炉况，但也基板清洁度。必须仔细校准，一旦固定，它们都或多或少的稳定的最佳气体流速，温度和压力的增长。即使这些条件得到满足，我们发现增长取决于图案的催化剂，面积，金额的催化剂和基板清洁度。因此，我们注册成立了多个催化剂坑的大小占增长的变异。一小时的高温退火步骤有助于从基片上除去任何污染物，例如:PR残余物等， 图2示出的条件，我们采用了单壁碳纳米管生长。 模具加工步骤从污染物的存在可能会导致在化学和生物传感中的杂散信号。因此，它是必要的化学官能化之前和之后彻底清洗衬底。每个功能化后的漂洗步骤有助于雷姆爱情任何多余的试剂，可能会坚持到活动区域附近的设备。我们还观察到，如果基板是不干净，PDMS印章不是无泄漏的，在测量过程中的流体压力与剥离。在这种情况下，很温柔O 2等离子体PDMS印章帮助附着力。太强大了O 2等离子体可能使PDMS印章棒很好，但很难从设备中删除，我们已经注意到翻录电极同时切除，使模具无法使用。如果PDMS印章和模具是干净的，它们之间的附着力好足够的生存也没有氧气等离子体处理的流体流量测量。不要使用单壁碳纳米管设备上，这会腐蚀碳纳米管O 2等离子体。 基于在溶液中的电气测量，压倒任何漏电流的检测信号。此泄漏发生，因为该解决方案也可以作为导体，其电阻下降而增加盐的浓度。因此，有必要将电极钝化步骤中的晶体管的设计。两个毯沉积在我们的制造协议（500纳米和20纳米SiO 2的 ）帮助减少泄漏源，漏极和栅极金属触点。此外，在采取任何电气测量之前，建议的栅 - 漏漏扫描，以确保无泄漏发生在预定的扫描电压范围。 流体流量的测量，这是必要的，以避免空气的流路中的陷阱。空气间隙导致的信号失真比在溶液中时，因为空气中的传感器的响应是不同的。在正向泵送模式中经常会遇到这个问题，阻碍流体流动。这避免了通过操作注射器泵在撤回模式。 用于高频的电气测量在高背景的离子溶液的响应下降30-40 MHz的范围以外由于谐振损失的测量设置。我们LIEVE这能够提高设计设备以较低的寄生效应。优化门单壁碳纳米管的距离和更小的BNC和SMA电缆可能有助于提高灵敏度。

当一个带电的分子结合到一个单壁碳纳米管或NW电子传感器，它可以捐赠/接受电子或作为本地静电门。在这两种情况下，可以改变结合的分子的电荷密度，在单壁碳纳米管或洒水通道，导致的传感器测得的直流电导的变化。 15-20分子种类繁多的已成功检测研究纳米传感器的直流特性，在这样的结合事件。即使负责检测的传感机理有很多优势，包括无标记检测21毫微摩尔灵敏度22，和电子读出能力15，它是有效的，只有在低离子强度的解决方案。在高离子强度的解决方案，直流检测离子筛选6-8阻碍。甲充电表面吸引，形成表面附近的双电层（EDL）从该溶液中的反离子。 EDL有效地筛选这些费用。作为T他的离子强度的溶液的增加，EDL的变窄和筛选的增加。该屏蔽效应的Debye屏蔽长度λD，其特征在于<img alt="公式1" fo:src="/files/ftp_upload/50438/50438eq1highres.jpg" src="/files/ftp_upload/50438/50438eq1.jpg" /> 式中，ε是介质的介电常数，K B为玻尔兹曼常数，T是温度，q是电子电荷，c为电解质溶液的离子强度。对于一个典型的100mM的缓冲溶液中，λD是约1纳米的表面电势在几个纳米的距离完全筛选。其结果是，大多数基于单壁碳纳米管或纳米线的纳米电子传感器操作无论是在干燥状态下20或低离子强度的解决方案5,15,17,21-22（C〜1纳米- 10毫摩尔），其他样品需要进行脱盐步骤15,23。点医疗诊断设备需要在生理离子强度，来样加工能力有限的病人站点。因此，减轻离子筛查效果是POC纳米电子生物传感器的开发和实施的关键。 我们减轻离子屏蔽效应，基于单壁碳纳米管的纳米电子传感器在兆赫频率范围经营。该协议提供了在这里制造的单壁碳纳米管晶体管为基础的纳米电子传感平台和高频混合测量生物分子检测。单壁碳纳米管生长的化学气相沉积图案与Fe催化剂24的基板上。对于我们的单壁碳纳米管晶体管，我们将暂停顶栅25以上500纳米的碳纳米管，这有助于提高高频传感器响应，也可以让一个紧凑的MICR邻流体腔密封装置。单壁碳纳米管晶体管的运行为高频率混频器9-11为了克服背景离子屏蔽效应。在高频时，溶液中的离子移动，没有足够的时间来形成的EDL和生物分子偶极子波动仍然可以门的单壁碳纳米管生成一个混合的电流，这是我们的传感信号。混频的产生是由于碳纳米管FET的非线性伏安特性 。我们的检测技术不同于常规技术充电检测和阻抗谱26-27。首先，我们检测生物分子的偶极子的相关费用，而不是在高频率。其次，高跨导的单壁碳纳米管的晶体管提供了一个内部​​的感测信号的增益。这省却了在高频阻抗测量的情况下，无需进行外部放大。最近，其他团体也高BA解决生物分子检测的ckground浓度23,28。然而，这些方法更复杂，需要复杂的制造或受体分子化学工程小心。我们的高频率的单壁碳纳米管传感器采用一个简单的设计，并采用混合属性的碳纳米管晶体管的固有频率。我们有能力，以减轻离子屏蔽效应，因此看好一个新的生物传感平台进行实时的护理点检测，需要直接在生理条件下运作的生物传感器。

<table border="1">
	<tbody>
		<tr>
			<td>&nbsp;</td>
			<td>&nbsp;</td>
			<td>&nbsp;</td>
			<td>REAGENTS</td>
		</tr>
		<tr>
			<td>&nbsp;</td>
			<td>&nbsp;</td>
			<td>&nbsp;</td>
			<td>Reagents which were provided within Lurie Nanofabrication Facility (University of Michigan) are marked as LNF in the catalogue column. Chemicals which require protective equipment (gloves, safety goggles, face mask, apron) and/or fume hood are denoted with PPE in comments section.</td>
		</tr>
		<tr>
			<td>Silicon wafers (P-type, &lt;100&gt;, 500-550 &mu;m thick)</td>
			<td>Silicon Valley Microelectronics</td>
			<td>&nbsp;</td>
			<td>&nbsp;</td>
		</tr>
		<tr>
			<td>SPR 220 3.0</td>
			<td>Dow (Rohm and Haas)</td>
			<td>Megaposit SPR</td>
			<td>PPE</td>
		</tr>
		<tr>
			<td>AZ 300MIF</td>
			<td>AZ Electronic Material Corporation</td>
			<td>&nbsp;</td>
			<td>PPE</td>
		</tr>
		<tr>
			<td>Acetone</td>
			<td>J T Baker</td>
			<td>9005-05</td>
			<td>PPE</td>
		</tr>
		<tr>
			<td>Isopropanol (IPA)</td>
			<td>J T Baker</td>
			<td>9079-05</td>
			<td>&nbsp;</td>
		</tr>
		<tr>
			<td>Buffered Hydrofluoric Acid</td>
			<td>Transene</td>
			<td>&nbsp;</td>
			<td>PPE</td>
		</tr>
		<tr>
			<td>1-Pyrene Butanoic Acid, succinimidyl ester</td>
			<td>Molecular Probes</td>
			<td>P130</td>
			<td>PPE</td>
		</tr>
		<tr>
			<td>Biotin PEO Amine</td>
			<td>Thermo Scientific</td>
			<td>EZ- Link PEG2 Biotin, # 21346</td>
			<td>PPE</td>
		</tr>
		<tr>
			<td>Streptavidin</td>
			<td>Invitrogen</td>
			<td>S 888</td>
			<td>PPE</td>
		</tr>
		<tr>
			<td>Dimethylformamide</td>
			<td>MP Biomedicals</td>
			<td>0219514791</td>
			<td>PPE</td>
		</tr>
		<tr>
			<td>Polydimethylsiloxane Elastomer Base and Curing Agent</td>
			<td>Dow Corning</td>
			<td>Sylgard 184 Elastomer Kit</td>
			<td>PPE</td>
		</tr>
		<tr>
			<td>SU-8 2015</td>
			<td>Microchem</td>
			<td>Y111064</td>
			<td>PPE</td>
		</tr>
		<tr>
			<td>SU-8 Developer</td>
			<td>Microchem</td>
			<td>Y020100</td>
			<td>PPE</td>
		</tr>
		<tr>
			<td>Silanizing agent</td>
			<td>Sigma Aldrich</td>
			<td>452807</td>
			<td>PPE</td>
		</tr>
		<tr>
			<td>Hydrogen</td>
			<td>Purity Plus</td>
			<td>LNF</td>
			<td>&nbsp;</td>
		</tr>
		<tr>
			<td>Ethylene</td>
			<td>Purity Plus</td>
			<td>LNF</td>
			<td>&nbsp;</td>
		</tr>
		<tr>
			<td>Argon</td>
			<td>Purity Plus</td>
			<td>LNF</td>
			<td>&nbsp;</td>
		</tr>
		<tr>
			<td>Phosphate Buffer Saline System</td>
			<td>Sigma Aldrich</td>
			<td>PBS1</td>
			<td>&nbsp;</td>
		</tr>
		<tr>
			<td>&nbsp;</td>
			<td>&nbsp;</td>
			<td>&nbsp;</td>
			<td>EQUIPMENT</td>
		</tr>
		<tr>
			<td>&nbsp;</td>
			<td>&nbsp;</td>
			<td>&nbsp;</td>
			<td>Equipment provided by Lurie Nanofabrication Facility (University of Michigan) is denoted as LNF in Catalogue column.</td>
		</tr>
		<tr>
			<td>GCA 200 Autostepper</td>
			<td>GCA</td>
			<td>LNF</td>
			<td>&nbsp;</td>
		</tr>
		<tr>
			<td>Low Pressure Chemical Vapor Deposition Tool</td>
			<td>Tempress</td>
			<td>LNF</td>
			<td>&nbsp;</td>
		</tr>
		<tr>
			<td>e-beam Evaporator</td>
			<td>Enerjet</td>
			<td>LNF</td>
			<td>&nbsp;</td>
		</tr>
		<tr>
			<td>CNT growth Furnace</td>
			<td>First Nano</td>
			<td>Easy Tube 3000 (LNF)</td>
			<td>&nbsp;</td>
		</tr>
		<tr>
			<td>Photomasks</td>
			<td>Nanofilm</td>
			<td>LNF</td>
			<td>&nbsp;</td>
		</tr>
		<tr>
			<td>Petri dish (150mm)</td>
			<td>&nbsp;</td>
			<td>LNF</td>
			<td>&nbsp;</td>
		</tr>
		<tr>
			<td>Desiccator</td>
			<td>Bel-Art</td>
			<td>F420100000</td>
			<td>&nbsp;</td>
		</tr>
		<tr>
			<td>Biopsy Punch</td>
			<td>Ted Pella</td>
			<td>15071/78</td>
			<td>&nbsp;</td>
		</tr>
		<tr>
			<td>Scalpel</td>
			<td>Ted Pella</td>
			<td>548</td>
			<td>&nbsp;</td>
		</tr>
		<tr>
			<td>Polyethylene Tubing PE-50</td>
			<td>VWR</td>
			<td>20903-414</td>
			<td>&nbsp;</td>
		</tr>
		<tr>
			<td>Syringe Pump</td>
			<td>New Era Pump Systems</td>
			<td>NE-1000</td>
			<td>&nbsp;</td>
		</tr>
		<tr>
			<td>Syringe</td>
			<td>Fisher Scientific</td>
			<td>BD Safety-Lok Syringes</td>
			<td>&nbsp;</td>
		</tr>
		<tr>
			<td>Syringe Needles</td>
			<td>Fisher Scientific</td>
			<td>14-821-13A</td>
			<td>&nbsp;</td>
		</tr>
		<tr>
			<td>DAQ card</td>
			<td>National Instruments</td>
			<td>779111-01</td>
			<td>&nbsp;</td>
		</tr>
		<tr>
			<td>GPIB connector</td>
			<td>National Instruments</td>
			<td>778032-51</td>
			<td>&nbsp;</td>
		</tr>
		<tr>
			<td>Lock-in Amplifier</td>
			<td>Stanford Research Systems</td>
			<td>SR 830</td>
			<td>&nbsp;</td>
		</tr>
		<tr>
			<td>Frequency Generator</td>
			<td>HP Agilent</td>
			<td>8648B, 9kHz -2GHz</td>
			<td>&nbsp;</td>
		</tr>
		<tr>
			<td>Bias Tee</td>
			<td>Picosecond</td>
			<td>5575A-104</td>
			<td>&nbsp;</td>
		</tr>
		<tr>
			<td>Current Preamplifier</td>
			<td>DL Instruments, LLC</td>
			<td>DL 1211</td>
			<td>&nbsp;</td>
		</tr>
		<tr>
			<td>BNC cables</td>
			<td>Allied Electronics</td>
			<td>665-xxxx</td>
			<td>&nbsp;</td>
		</tr>
		<tr>
			<td>SMA cables</td>
			<td>Sentro Tech Corp</td>
			<td>SCF65141</td>
			<td>&nbsp;</td>
		</tr>
	</tbody>
</table>

fabrication of carbon nanotube high-frequency nanoelectronic biosensor for sensing in high ionic strength solutions

独特的电子性质和高表面体积比单壁碳纳米管（SWNT）和半导体纳米线（NW）1-4让他们很好的候选人高灵敏度生物传感器。当一个带电的分子结合到这样的传感器表面，改变载流子浓度5中的传感器，其直流电导的变化。然而，在离子溶液中的带电表面也吸引从该溶液中的反离子，形成双电层（EDL）。此EDL有效屏幕关闭充电，在生理条件下〜100毫摩尔（MM），特性电荷屏蔽长度（德拜长度）小于一个纳米（nm）。因此，在高离子强度的解决方案，充电（DC）的检测是从根本上阻碍了6-8。 我们克服电荷屏蔽效应通过检测分子偶极子，而不是在高频收费，经营碳nanot的UBE场效应晶体管高频混频器9-11。在高频时，的AC驱动力可以不再克服溶液阻力和溶液中的离子，没有足够的时间形成的EDL。另外，混频技术使我们能够工作在频率足够高，以克服离子筛选，而检测到的感测信号在较低频率的11-12。此外，单壁碳纳米管的晶体管的高跨导，提供了一个内部​​的感测信号，避免了需要外部信号放大器的增益。 在这里，我们描述协议（一）制备单壁碳纳米管晶体管，（二）生物分子功能化碳纳米管13，（C）到设备设计和邮票聚二甲基硅氧烷（PDMS）微流控室14，（D）进行高频感应离子强度不同的解决方案11。

与一个暂停的顶栅晶体管的单壁碳纳米管的扫描电子显微镜图像如图7a所示。悬架25的栅极的尺寸是至关重要的。目前的设计尺寸（长x宽x厚= 25微米×1微米×100纳米）。栅电极由50纳米Cr/50纳米金，粗壮的镀铬层悬挂结构增加了更多的力量。也证实了顶部栅极和漏极（ 图7b）之间的泄漏电流的情况下的悬浮结构。 我们用生物素 - 链霉亲和素配体 - 受体系统评估我们的单壁碳纳米管传感器。为了表征的侧壁官能化的成功，我们监测每个官能化步骤之后在空气中的FET直流转换曲线， 图7c示出的转移曲线向右偏移生物素化后（红色）和链霉抗生物素蛋白结合（蓝色）。这可以归因于由electronegati的静电选通VE胺组PEO-胺素和链霉亲和素。 对于高频测量中，我们遵循的示意图如图5b所示。单壁碳纳米管晶体管非线性IV特性，混合在源极和栅极的高频输入，产生一个混合的电流输出，我的组合，这是我们的传感信号。 图7d显示混合作为栅极电压的函数测量一个典型的设备氯化钠100毫米。的混合电流的AM调制的输入调制频率，ω 米 ，由下式给出10-11 <img alt="公式1" fo:src="/files/ftp_upload/50438/50438eq2highres.jpg" src="/files/ftp_upload/50438/50438eq2.jpg" /> ，其中，m是调制深度， 伏交流输入振幅是AM和∂/∂V G的是该装置的跨导（在I的斜率<sUB&gt; DC-V G曲线图7D）。混合（M = 0.78，V AC = 20mV的）同意目前的结果与图中所示的模型。对于静态流体测量中，我们比较这种混合电流扫描功能化碳纳米管的高峰期。流量测量，我们固定载波频率调幅调制信号和固定栅极电压（V G = 0）和监视I 混合生物分子结合作为时间的函数，同时保持一个稳定的流体流动图7E-7F代表结果静态和流量测量。 对于生物分子的检测，它是必要的，在CNT被直接暴露到该溶液中即 SiO 2的过程中被完全蚀刻掉的BHF的蚀刻步骤。如果这个条件不能满足，化学修饰碳纳米管的连接分子是不可能的，因为不能叠加沿着碳纳米管侧壁。这显然是图7G所示，在这里我们看到前没有变化，甚至在去离子水结合后的SiO 2钝化设备。这也证明了我们的测量结果表明，成功的化学修饰以及高背景离子浓度的生物分子检测。在所有的测量中，我们观察到超过30MHz，这是由于从设定共振传感器的响应下降。 <img alt="图1" fo:content-width="4.5in" fo:src="/files/ftp_upload/50438/50438fig1highres.jpg" src="/files/ftp_upload/50438/50438fig1.jpg" /> 图1。纳米管的晶体管的制造处理流程（一）的制备过程中- （1）光掩模的第1层（PL-1）的催化剂沉积，（2）金属离地升空，（3）碳纳米管的生长，（4）PL-2的源极-漏极接触，金属离地升空（5），（6）的SiO 2沉积毯，（7）的PL-3栅极接触，（8）金属离地升空，（9）薄膜的SiO 2的厚沉积，（10）PL-4 BHF湿法刻蚀信道和（11）除去光致抗蚀剂后的最终的移动设备。颜色方案说明（二）设备结构示意图。 <img alt="图2" fo:content-width="3.75in" fo:src="/files/ftp_upload/50438/50438fig2highres.jpg" src="/files/ftp_upload/50438/50438fig2.jpg" /> 图2。碳纳米管的生长。（一）退火工序去除光刻胶残渣，（二）生长步骤碳纳米管的生长及（c）设备放置在生长炉。 <img alt="图3" fo:content-width="4.25in" fo:src="/files/ftp_upload/50438/50438fig3highres.jpg" src="/files/ftp_upload/50438/50438fig3.jpg" /> 图3。化学功能化CNT的流程图。 图4。 PDMS印章的测量解决方案。（A） - （B）静态（无流量）测量。（一）打孔和安装PMDS室（二）设备，设备上的流室示意图（三） - （E ）流量测量。（三）工艺流程PDMS流道采用SU-8模具。 （1）用于限定流路的光掩模，SU-8的模具（2）交联，（3）的PDMS上的SU-8（4）加盖到移动设备上的PDMS流动通道（四）在设备上的流路的示意图，并（五）在PDMS中，入口/出口孔冲孔冲压设备上的流路和聚乙烯管连接到进/出口。 <img alt="图5" fo:content宽度="5英寸"FO:SRC ="/ files/ftp_upload/50438/50438fig5highres.jpg"SRC ="/ files/ftp_upload/50438/50438fig5.jpg"/&gt; 图5。 （一）电气测量设置。直流测量原理图，（二）交流混合电流测量原理图及（c）实验装置的图像AM调频混合测量。 <a href="https://www-jove-com.remotexs.ntu.edu.sg/files/ftp_upload/50438/50438fig5large.jpg" target="_blank">点击这里查看大图</a> 。 <img alt="图6" fo:content-width="5in" fo:src="/files/ftp_upload/50438/50438fig6highres.jpg" src="/files/ftp_upload/50438/50438fig6.jpg" /> 图6。流量测量装置（A）图片的整个测量设置;（二）注射泵和探针台;及（c）图像的设备与PDMS流道，入口/出口流量管和电气p长袍。 <img alt="图7" fo:content-width="4.5in" fo:src="/files/ftp_upload/50438/50438fig7highres.jpg" src="/files/ftp_upload/50438/50438fig7.jpg" /> 图7。代表性的结果（一）的SEM图像的一个典型的悬浮顶栅装置，（二）的栅极-漏极泄漏确认悬挂结构，（c）本人 克直流 -V曲线为原始碳纳米管FET（黑色）后， 单壁碳纳米管的生物传感器。生物素化的（红色）和链霉抗生物素蛋白结合后（蓝色）在空气中测得，（四）的直流电流， 直流 （黑色，V SD = 10毫伏）及混合电流，我混合 （红，调制F = 200千赫）为一个函数的V 克 ，100mM NaCl的溶液中的装置。理论混合得到利用该模型在方程（1）中还示出用于比较（▲）。（五）吾混合 -V 克电流生物素（黑色）和链霉亲和素-生物素结合的单壁碳纳米管（红色） 在 f = 10 MHz的氯化钠100毫米，（六）实时流量测量检测链霉亲和素氯化钠100毫米及（g）在结合后的信号变化结合VES完全钝化控制装置在去离子水中，在不同的频率。 <a href="https://www-jove-com.remotexs.ntu.edu.sg/files/ftp_upload/50438/50438fig7large.jpg" target="_blank">点击这里查看大图</a> 。

Watch this Scientific Journal Video about Fabrication of Carbon Nanotube High-Frequency Nanoelectronic Biosensor for Sensing in High Ionic Strength Solutions at JoVE.com

Fabrication of Carbon Nanotube High-Frequency Nanoelectronic Biosensor for Sensing in High Ionic Strength Solutions

我们描述了基于碳纳米管的高频生物传感器的设备制造和测量协议。高频感应技术减轻了根本的离子（德拜）的屏蔽效应，并允许碳纳米管传感器工作在高离子强度的解决方案，在传统的电子生物传感器失败。我们的技术提供了一个独特的平台，为护理点（POC）在生理条件下运行的电子生物传感器。

制备碳纳米管高频纳电子生物传感器的传感高离子强度的解决方案

The unique electronic properties and high surface-to-volume ratios of single-walled carbon nanotubes (SWNT) and semiconductor nanowires (NW) 1-4 make them ...

fabrication-carbon-nanotube-high-frequency-nanoelectronic-biosensor

Nanyang Technological University

Research

JoVE Journal

Bioengineering

18.1K 次观看.  University of Michigan - Ann Arbor. 我们描述了基于碳纳米管的高频生物传感器的设备制造和测量协议。高频感应技术减轻了根本的离子（德拜）的屏蔽效应，并允许碳纳米管传感器工作在高离子强度的解决方案，在传统的电子生物传感器失败。我们的技术提供了一个独特的平台，为护理点（POC）在生理条件下运行的电子生物传感器。

视频: Fabrication of Carbon Nanotube High-Frequency Nanoelectronic Biosensor for Sensing in High Ionic Strength Solutions

Fabrication of Electrochemical-DNA Biosensors for the Reagentless Detection of Nucleic Acids, Proteins and Small Molecules

As medicine is currently practiced, doctors send specimens to a central laboratory for testing and thus must wait hours or days to 
receive the results. Many patients would be better served by rapid, bedside tests. To this end our laboratory and others have developed a versatile, reagentless 
biosensor platform that supports the quantitative, reagentless, electrochemical detection of nucleic acids (DNA, RNA), proteins (including antibodies) and small 
molecules analytes directly in unprocessed clinical and environmental samples. In this video, we demonstrate the preparation and use of several biosensors in this 
"E-DNA" class. In particular, we fabricate and demonstrate sensors for the detection of a target DNA sequence in a polymerase chain reaction mixture, an HIV-specific antibody and the drug cocaine. The preparation procedure requires only three hours of hands-on effort followed by an overnight incubation, and their use requires only minutes.

"E-DNA" sensors, reagentless, electrochemical biosensors that perform well even when challenged directly in blood and other complex matrices, have been adapted to the detection of a wide range of nucleic acid, protein and small molecule analytes. Here we present a general procedure for the fabrication and use of such sensors.

无试剂检测核酸，蛋白质和小分子的电化学DNA生物传感器的制备

As medicine is currently practiced, doctors send specimens to a central laboratory for testing and thus must wait hours or days to 
receive the results. ...

科研

生物工程

Hollow Microneedle-based Sensor for Multiplexed Transdermal Electrochemical Sensing

The development of a minimally invasive multiplexed monitoring system for rapid analysis of biologically-relevant molecules could offer individuals suffering from chronic medical conditions facile assessment of their immediate physiological state. Furthermore, it could serve as a research tool for analysis of complex, multifactorial medical conditions. In order for such a multianalyte sensor to be realized, it must be minimally invasive, sampling of interstitial fluid must occur without pain or harm to the user, and analysis must be rapid as well as selective.
Initially developed for pain-free drug delivery, microneedles have been used to deliver vaccines and pharmacologic agents (e.g., insulin) through the skin.1-2 Since these devices access the interstitial space, microneedles that are integrated with microelectrodes can be used as transdermal electrochemical sensors. Selective detection of glucose, glutamate, lactate, hydrogen peroxide, and ascorbic acid has been demonstrated using integrated microneedle-electrode devices with carbon fibers, modified carbon pastes, and platinum-coated polymer microneedles serving as transducing elements.3-7,8
This microneedle sensor technology has enabled a novel and sophisticated analytical approach for in situ and simultaneous detection of multiple analytes. Multiplexing offers the possibility of monitoring complex microenvironments, which are otherwise difficult to characterize in a rapid and minimally invasive manner. For example, this technology could be utilized for simultaneous monitoring of extracellular levels of, glucose, lactate and pH,9 which are important metabolic indicators of disease states7,10-14 (e.g., cancer proliferation) and exercise-induced acidosis.15

This article details the construction of a multiplexed microneedle-based sensor. The device is being developed for in situ sampling and electrochemical analysis of multiple analytes in a rapid and selective manner. We envision clinical medicine and biomedical research uses for these microneedle-based sensors.

复透皮电化学传感的空心微针为基础的传感器

The development of a minimally invasive multiplexed monitoring system for rapid analysis of biologically-relevant molecules could offer individuals ...

Biosensor for Detection of Antibiotic Resistant Staphylococcus Bacteria

A structurally transformed lytic bacteriophage having a broad host range of Staphylococcus aureus strains and a penicillin-binding protein (PBP 2a) antibody conjugated latex beads have been utilized to create a biosensor designed for discrimination of methicillin resistant (MRSA) and sensitive (MSSA) S. aureus species 1,2. The lytic phages have been converted into phage spheroids by contact with water-chloroform interface. Phage spheroid monolayers have been moved onto a biosensor surface by Langmuir-Blodgett (LB) technique 3. The created biosensors have been examined by a quartz crystal microbalance with dissipation tracking (QCM-D) to evaluate bacteria-phage interactions. Bacteria-spheroid interactions led to reduced resonance frequency and a rise in dissipation energy for both MRSA and MSSA strains. After the bacterial binding, these sensors have been further exposed to the penicillin-binding protein antibody latex beads. Sensors analyzed with MRSA responded to PBP 2a antibody beads; although sensors inspected with MSSA gave no response. This experimental distinction determines an unambiguous discrimination between methicillin resistant and sensitive S. aureus strains. Equally bound and unbound bacteriophages suppress bacterial growth on surfaces and in water suspensions. Once lytic phages are changed into spheroids, they retain their strong lytic activity and show high bacterial capture capability. The phage and phage spheroids can be utilized for testing and sterilization of antibiotic resistant microorganisms. Other applications may include use in bacteriophage therapy and antimicrobial surfaces.

Lytic phage biosensors and antibody beads are able to discriminate between methicillin resistant (MRSA) and sensitive staphylococcus bacteria. The phages were immobilized by a Langmuir-Blodgett method onto a surface of a quartz crystal microbalance sensor and worked as broad range staphylococcus probes. Antibody beads recognize MRSA.

生物传感器检测抗生素耐药金黄色葡萄球菌细菌

A structurally transformed lytic bacteriophage having a broad host range of Staphylococcus aureus strains and a penicillin-binding protein (PBP 2a) ...

Electric Cell-substrate Impedance Sensing for the Quantification of Endothelial Proliferation, Barrier Function, and Motility

Electric Cell-substrate Impedance Sensing (ECIS) is an in vitro impedance measuring system to quantify the behavior of cells within adherent cell layers. To this end, cells are grown in special culture chambers on top of opposing, circular gold electrodes. A constant small alternating current is applied between the electrodes and the potential across is measured. The insulating properties of the cell membrane create a resistance towards the electrical current flow resulting in an increased electrical potential between the electrodes. Measuring cellular impedance in this manner allows the automated study of cell attachment, growth, morphology, function, and motility. Although the ECIS measurement itself is straightforward and easy to learn, the underlying theory is complex and selection of the right settings and correct analysis and interpretation of the data is not self-evident. Yet, a clear protocol describing the individual steps from the experimental design to preparation, realization, and analysis of the experiment is not available. In this article the basic measurement principle as well as possible applications, experimental considerations, advantages and limitations of the ECIS system are discussed. A guide is provided for the study of cell attachment, spreading and proliferation; quantification of cell behavior in a confluent layer, with regard to barrier function, cell motility, quality of cell-cell and cell-substrate adhesions; and quantification of wound healing and cellular responses to vasoactive stimuli. Representative results are discussed based on human microvascular (MVEC) and human umbilical vein endothelial cells (HUVEC), but are applicable to all adherent growing cells.

This protocol reviews Electric Cell-substrate Impedance Sensing, a method to record and analyze the impedance spectrum of adherent cells for the quantification of cell attachment, proliferation, motility, and cellular responses to pharmacological and toxic stimuli. Detection of endothelial permeability and assessment of cell-cell and cell-substrate contacts are emphasized.

电动细胞基质阻抗传感的内皮细胞增殖，屏障功能的定量和能动性

Electric Cell-substrate Impedance Sensing (ECIS) is an in vitro impedance measuring system to quantify the behavior of cells within adherent cell layers. ...

Targeted Plasma Membrane Delivery of a Hydrophobic Cargo Encapsulated in a Liquid Crystal Nanoparticle Carrier

The controlled delivery of drug/imaging agents to cells is critical for the development of therapeutics and for the study of cellular signaling processes. Recently, nanoparticles (NPs) have shown significant promise in the development of such delivery systems. Here, a liquid crystal NP (LCNP)-based delivery system has been employed for the controlled delivery of a water-insoluble dye, 3,3&#8242;-dioctadecyloxacarbocyanine perchlorate (DiO), from within the NP core to the hydrophobic region of a plasma membrane bilayer. During the synthesis of the NPs, the dye was efficiently incorporated into the hydrophobic LCNP core, as confirmed by multiple spectroscopic analyses. Conjugation of a PEGylated cholesterol derivative to the NP surface (DiO-LCNP-PEG-Chol) enabled the binding of the dye-loaded NPs to the plasma membrane in HEK 293T/17 cells. Time-resolved laser scanning confocal microscopy and F&#246;rster resonance energy transfer (FRET) imaging confirmed the passive efflux of DiO from the LCNP core and its insertion into the plasma membrane bilayer. Finally, the delivery of DiO as a LCNP-PEG-Chol attenuated the cytotoxicity of DiO; the NP form of DiO exhibited ~30-40% less toxicity compared to DiOfree delivered from bulk solution. This approach demonstrates the utility of the LCNP platform as an efficient modality for the membrane-specific delivery and modulation of hydrophobic molecular cargos.

A liquid crystal nanoparticle (LCNP) nanocarrier is exploited as a vehicle for the controlled delivery of a hydrophobic cargo to the plasma membrane of living cells.

疏水货物封装有针对性的质膜交付在液晶纳米颗粒载体

The controlled delivery of drug/imaging agents to cells is critical for the development of therapeutics and for the study of cellular signaling processes. ...

Dry Film Photoresist-based Electrochemical Microfluidic Biosensor Platform: Device Fabrication, On-chip Assay Preparation, and System Operation

In recent years, biomarker diagnostics became an indispensable tool for the diagnosis of human disease, especially for the point-of-care diagnostics. An easy-to-use and low-cost sensor platform is highly desired to measure various types of analytes (e.g., biomarkers, hormones, and drugs) quantitatively and specifically. For this reason, dry film photoresist technology - enabling cheap, facile, and high-throughput fabrication - was used to manufacture the microfluidic biosensor presented here. Depending on the bioassay used afterwards, the versatile platform is capable of detecting various types of biomolecules. For the fabrication of the device, platinum electrodes are structured on a flexible polyimide (PI) foil in the only clean-room process step. The PI foil serves as a substrate for the electrodes, which are insulated with an epoxy-based photoresist. The microfluidic channel is subsequently generated by the development and lamination of dry film photoresist (DFR) foils onto the PI wafer. By using a hydrophobic stopping barrier in the channel, the channel is separated into two specific areas: an immobilization section for the enzyme-linked assay and an electrochemical measurement cell for the amperometric signal readout.
The on-chip bioassay immobilization is performed by the adsorption of the biomolecules to the channel surface. The glucose oxidase enzyme is used as a transducer for electrochemical signal generation. In the presence of the substrate, glucose, hydrogen peroxide is produced, which is detected at the platinum working electrode. The stop-flow technique is applied to obtain signal amplification along with rapid detection. Different biomolecules can quantitatively be measured by means of the introduced microfluidic system, giving an indication of different types of diseases, or, in regard to therapeutic drug monitoring, facilitating a personalized therapy.

A microfluidic biosensor platform was designed and fabricated using low-cost dry film photoresist technology for the rapid and sensitive quantification of various analytes. This single-use system allows for the electrochemical readout of on-chip-immobilized enzyme-linked assays by means of the stop-flow technique.

干膜 Photoresist-based 电化学微流控生物传感器平台: 器件制备、片上检测制备及系统运行

In recent years, biomarker diagnostics became an indispensable tool for the diagnosis of human disease, especially for the point-of-care diagnostics. An ...

Development of an Electrochemical DNA Biosensor to Detect a Foodborne Pathogen

Vibrio parahaemolyticus (V. parahaemolyticus) is a common foodborne pathogen that contributes to a large proportion of public health problems globally, significantly affecting the rate of human mortality and morbidity. Conventional methods for the detection of V. parahaemolyticus such as culture-based methods, immunological assays, and molecular-based methods require complicated sample handling and are time-consuming, tedious, and costly. Recently, biosensors have proven to be a promising and comprehensive detection method with the advantages of fast detection, cost-effectiveness, and practicality. This research focuses on developing a rapid method of detecting V. parahaemolyticus with high selectivity and sensitivity using the principles of DNA hybridization. In the work, characterization of synthesized polylactic acid-stabilized gold nanoparticles (PLA-AuNPs) was achieved using X-ray Diffraction (XRD), Ultraviolet-visible Spectroscopy (UV-Vis), Transmission Electron Microscopy (TEM), Field-emission Scanning Electron Microscopy (FESEM), and Cyclic Voltammetry (CV). We also carried out further testing of stability, sensitivity, and reproducibility of the PLA-AuNPs. We found that the PLA-AuNPs formed a sound structure of stabilized nanoparticles in aqueous solution. We also observed that the sensitivity improved as a result of the smaller charge transfer resistance (Rct) value and an increase of active surface area (0.41 cm2). The development of our DNA biosensor was based on modification of a screen-printed carbon electrode (SPCE) with PLA-AuNPs and using methylene blue (MB) as the redox indicator. We assessed the immobilization and hybridization events by differential pulse voltammetry (DPV). We found that complementary, non-complementary, and mismatched oligonucleotides were specifically distinguished by the fabricated biosensor. It also showed reliably sensitive detection in cross-reactivity studies against various food-borne pathogens and in the identification of V. parahaemolyticus in fresh cockles.

A protocol for the development of an electrochemical DNA biosensor comprising a polylactic acid-stabilized, gold nanoparticles-modified, screen-printed carbon electrode to detect Vibrio parahaemolyticus is presented.

电化学 DNA 生物传感器检测食源性致病菌的研究进展

Vibrio parahaemolyticus (V. parahaemolyticus) is a common foodborne pathogen that contributes to a large proportion of public health problems globally, ...

Improved 3D Hydrogel Cultures of Primary Glial Cells for In Vitro Modelling of Neuroinflammation

In the central nervous system, numerous acute injuries and neurodegenerative disorders, as well as implanted devices or biomaterials engineered to enhance function result in the same outcome: excess inflammation leads to gliosis, cytotoxicity, and/or formation of a glial scar that collectively exacerbate injury or prevent healthy recovery. With the intent of creating a system to model glial scar formation and study inflammatory processes, we have generated a 3D cell scaffold capable of housing primary cultured glial cells: microglia that regulate the foreign body response and initiate the inflammatory event, astrocytes that respond to form a fibrous scar, and oligodendrocytes that are typically vulnerable to inflammatory injury. The present work provides a detailed step-by-step method for the fabrication, culture, and microscopic characterization of a hyaluronic acid-based 3D hydrogel scaffold with encapsulated rat brain-derived glial cells. Further, protocols for characterization of cell encapsulation and the hydrogel scaffold by confocal immunofluorescence and scanning electron microscopy are demonstrated, as well as the capacity to modify the scaffold with bioactive substrates, with incorporation of a commercial basal lamina mixture to improved cell integration.

Improved 3D Hydrogel Cultures of Primary Glial Cells for In Vitro Modelling of Neuroinflammation

Herein, we present a protocol for the 3D culture of rat brain-derived glia cells, including astrocytes, microglia, and oligodendrocytes. We demonstrate primary cell culture, methacrylated hyaluronic acid (HAMA) hydrogel synthesis, HAMAphoto-polymerization and cell encapsulation, and sample processing for confocal and scanning electron microscopic imaging.

改良的3D 水凝胶培养的原发胶质细胞的体外Neuroinflammation 模型的建立

In the central nervous system, numerous acute injuries and neurodegenerative disorders, as well as implanted devices or biomaterials engineered to enhance ...

Spontaneous Formation and Rearrangement of Artificial Lipid Nanotube Networks as a Bottom-Up Model for Endoplasmic Reticulum

We present a convenient method to form a bottom-up structural organelle model for the endoplasmic reticulum (ER). The model consists of highly dense lipidic nanotubes that are, in terms of morphology and dynamics, reminiscent of ER. The networks are derived from phospholipid double bilayer membrane patches adhering to a transparent Al2O3 substrate. The adhesion is mediated by Ca2+ in the ambient buffer. Subsequent depletion of Ca2+ by means of BAPTA/EDTA causes retraction of the membrane, resulting in spontaneous lipid nanotube network formation. The method only comprises phospholipids and microfabricated surfaces for simple formation of an ER model and does not require the addition of proteins or chemical energy (e.g., GTP or ATP). In contrast to the 3D morphology of the cellular endoplasmic reticulum, the model is two-dimensional (albeit the nanotube dimensions, geometry, structure, and dynamics are maintained). This unique in vitro ER model consists of only a few components, is easy to construct, and can be observed under a light microscope. The resulting structure can be further decorated for additional functionality, such as the addition of ER-associated proteins or particles to study transport phenomena among the tubes. The artificial networks described here are suitable structural models for the cellular ER, whose unique characteristic morphology has been shown to be related to its biological function, whereas details regarding formation of the tubular domain and rearrangements within are still not completely understood. We note that this method uses Al2O3 thin-film-coated microscopy coverslips, which are commercially available but require special orders. Therefore, it is advisable to have access to a microfabrication facility for preparation.

Solid-supported, protein-free, double phospholipid bilayer membranes (DLBM) can be transformed into complex and dynamic lipid nanotube networks and can serve as 2D bottom-up models of the endoplasmic reticulum.

人工脂质纳米管网作为内质神经的底部模型的自发形成与重排

We present a convenient method to form a bottom-up structural organelle model for the endoplasmic reticulum (ER). The model consists of highly dense ...

Fabrication of Ti3C2 MXene Microelectrode Arrays for In Vivo Neural Recording

Implantable microelectrode technologies have been widely used to elucidate neural dynamics at the microscale to gain a deeper understanding of the neural underpinnings of brain disease and injury. As electrodes are miniaturized to the scale of individual cells, a corresponding rise in the interface impedance limits the quality of recorded signals. Additionally, conventional electrode materials are stiff, resulting in a significant mechanical mismatch between the electrode and the surrounding brain tissue, which elicits an inflammatory response that eventually leads to a degradation of the device performance. To address these challenges, we have developed a process to fabricate flexible microelectrodes based on Ti3C2 MXene, a recently discovered nanomaterial that possesses remarkably high volumetric capacitance, electrical conductivity, surface functionality, and processability in aqueous dispersions. Flexible arrays of Ti3C2 MXene microelectrodes have remarkably low impedance due to the high conductivity and high specific surface area of the Ti3C2 MXene films, and they have proven to be exquisitely sensitive for recording neuronal activity. In this protocol, we describe a novel method for micropatterning Ti3C2 MXene into microelectrode arrays on flexible polymeric substrates and outline their use for in vivo micro-electrocorticography recording. This method can easily be extended to create MXene electrode arrays of arbitrary size or geometry for a range of other applications in bioelectronics and it can also be adapted for use with other conductive inks besides Ti3C2 MXene. This protocol enables simple and scalable fabrication of microelectrodes from solution-based conductive inks, and specifically allows harnessing the unique properties of hydrophilic Ti3C2 MXene to overcome many of the barriers that have long hindered the widespread adoption of carbon-based nanomaterials for high-fidelity neural microelectrodes.

Fabrication of Ti3C2 MXene Microelectrode Arrays for In Vivo Neural Recording

We describe here a method for fabricating Ti3C2 MXene microelectrode arrays and utilizing them for in vivo neural recording.

用于Vivo神经记录的 Ti3C2 MXene 微电极阵列的制造

Implantable microelectrode technologies have been widely used to elucidate neural dynamics at the microscale to gain a deeper understanding of the neural ...