英文参考文献及译文

致谢

外文资料原文

OLED-Polypropylene Bio-CD Sensor

Srikanth Vengasandra, Yuankun Cai, David Grewell, Joseph Shinar, and Ruth Shinarc

Department of Agricultural & Biosystems Engineering

Ames Laboratory - USDOE and Department of Physics & Astronomy MicroelectronicsResearchCenter and Department of Electrical and Computer Engineering IowaStateUniversity, Ames, IA50011, U.S.A

ABSTRACT

With the goal of developing microfluidic platforms for sensing applications, flash-free micropatterns were embossed in polypropylene surfaces with ultrasonic heating for a biosensing application. The embossed features were designed to act as reservoirs, valves, and reaction chambers to allow, in combination with a compact sensing platform, the monitoring of analyte levels using a standard PC-CD player. To generate the compact sensor, as an example, we chose the photoluminescence (PL)-based detection of lactate and glucose using an OLED-based sensing platform. Once embossed, the surface energy of the plastic substrate was chemically modified to make it hydrophilic. Reagents, placed in separate reservoirs, were directed through burst valves towards a reaction chamber via CD rotation. Lactate or glucose were monitored by measuring the effect of the related dissolved oxygen level on the PL decay time of an oxygen-sensitive dye, following analyte oxidation catalyzed by a suitable specific oxidase enzyme. The results demonstrate the potential of integrating OLEDs as excitation sources in PL-based sensors with microfluidic CD- based platforms, including for simultaneous multiple analyses.

Keywords: Organic light-emitting diode, OLED, lab-on-a-CD, glucose sensor, lactate sensor

1. INTRODUCTION

Biomedical microelectromechanical system (MEMS)-based sensing platforms fabricated on plasticsubstrates have the potential of e.g., being low cost, disposable, cross-contamination free, andsensitive. Additionally, such sensors show promise for high throughput and multianalyte detection. The advantages of such a lab-on-CD-based biosensing platform include its simplicity in terms ofusage for a wide range of solutions, versatility in terms of multianalyte detection feasibility, andcompact size. Moreover, valving is easily implemented in such CDs.

This paper describes the use of ultrasonic micro-embossing to generate microfluidic channels, valves, reservoirs, and reaction chambers in a polypropylene (PP) PC compact disc (CD). The ultrasonic micro-embossing was used as a source of localized heat. The advantages of this technique include short cycle times, ease of de-embossing and low residual stresses. Moreover, this approach is applicable to batch and continuous manufacturing, and it is simpler relative to the more common fabrication methods, injection molding and hot embossing. Microfluidic CD architectures were generated in materials such as polycarbonate [1], polystyrene (PS) [2], polydimethylsiloxane (PDMS) [3], and PP [4].

As shown in this work, the CD can be integrated with a photoluminescence (PL)-based OLED sensing platform to generate a compact device for monitoring e.g., lactate and glucose. The CD materials are typically of relatively low surface energy making them hydrophobic. However, many of the functions of microfluidic devices rely on hydrophilic properties so that the channel walls can promote capillary action and allow proper fluid flow. Thus, to increase the surface energy chemical treatment and oxidization by ozone or plasma are often used.

In the example shown, analytes such as glucose and lactate were monitored by utilizing an oxygen-sensitive dye embedded in a thin film. Glucose and lactate were oxidized in the presence of specific enzymes, i.e., glucose oxidase (GOx) and lactate oxidase (LOx), respectively, and oxygen. Oxygen is consumed in such reactions and in solution under specific experimental conditions the final dissolved oxygen (DO) level is related to the initial analyte concentration [5].

The enzymes can sometimes be embedded in a thin film; alternatively, they can be dissolved in solution. The consumption of DO in the oxidation reactions results in an increase in the PL intensity and the PL decay time of the oxygen-sensitive dye. In the preliminary measurements shown below, the OLED pixel array and the sensing film were structurally integrated by attaching two glass substrates, on which they were separately fabricated, back-to-back. The PL was monitored using a photomultiplier tube (PMT); small-size Si photodiode arrays, compatible with the design of the OLED pixels, are also usable and will lead to more compact, field-deployable sensors. A more compact sensor can be obtained also by integrating, in addition to the OLED excitation source and the sensing film, a thin film-based photodetector (PD). Such PDs based on amorphous or nanocrystalline Si are currently under development [6,7], however, their current slow speed does not allow monitoring oxygen in the t mode. Organic PDs are also suitable for such integration and possibly, for measuring of suitable luminophores.

2. EXPERIMENTAL PROCEDURE

2.1. Materials

Foamed and extruded sheets of PP were obtained from Trexel Corporation (MA). The foaming level specified by Trexel was approximately 15% to 20%. The PP sheets had a thickness of about 5 mm and were cut to the size of a standard CD with common shears. The PP–CDs were then used for ultrasonic micro-embossing.

The oxygen sensitive dye used, Pt octaethyporphyrin (PtOEP), was obtained from H. W. Sands. It was embedded in a film of PS (molecular weight 45000) obtained from Sigma-Aldrich. GOx (from Aspergillus niger), LOx (from pichia pastoris), glucose and L-lactate were purchased from Sigma-Aldrich. All reagents were dissolved in a phosphate buffered saline (PBS) solution. The OLED excitation source used was based on tris(quinolinolate) Al (Alq3).

2.2. CD preparation

Micro-embossing was performed in a foamed material using a Branson 2000 Series ultrasonic system that operated at 20 kHz using a titanium horn to generate the micro-features. The ultrasonic embossing was performed at 40 peak-to-peak amplitude and 0.5 s of heating time. These conditions were based on previous studies [4]. A cross section of a typical feature is seen in Fig. 1.The image on the left is an embossed feature on a microcellular foam substrate; in contrast, the image on the right is an embossed feature on standard PP. It is seen, as previously noted [4], that by using the foamed substrate the undesired flash is reduced. Thus, results presented here are only for structures in foamed substrates.

Fig. 1. Cross section of typical features: Left: micro-embossing of foamed PP to generate flash-free micro-patterns. Right: embossed features in a standard PP and the generated flash.

Fig. 2 shows the microchannels, reservoirs, and burst valves generated in the glucose/lactate Bio-CD. The CD contains four such sections that were used to detect different concentrations of glucose simultaneously. Valving was achieved by use of

burst valves. That is, valves in which capillary forces pin liquids at an enlargement in a microfluidic channel. A pressure generated by rotation at a “burst frequency,” which depends on the rotation speed and channel size, overcomes the capillary pressure, enabling fluid flow.

In order to increase the surface energy of the PP surface, treatments with a silicone surfactant and a proprietary organic system, which is believed to contain a surfactant, were evaluated. Each of the two resulting chemical coatings was individually studied and both produced the desired results. The coating solutions were obtained from Goulston Technologies, Inc (Monroe, NC) and were diluted with DI water at a ratio of 1:10; the dilution was applied directly onto the PP bio-CD sample substrate and allowed to cure for 15 minutes under clean, humid conditions. The treated surfaces were then rinsed with DI water and blown dry under air pressure. Contact angles for the pretreated and untreated PP sample surfaces were measured and compared to assess the increase in the surface energies.

To enable monitoring of the PL, the bottom part of the reaction chamber on the CD was cut and a sensing film deposited on a thin glass substrate was attached. This replacement to a transparent bottom of the reaction chamber was necessary to enable PL monitoring using a PD in a “back detection” geometry. In this geometry the PD is behind the OLED array, detecting the PL signal that reaches it through the gaps between the OLED pixels. The PL signal, which reflected back from the bio-CD was captured and measured by a Hamamatsu R6000 PMT operated at 900 V.

Fig. 2.Schematic (not to scale) of the Bio-CD components for monitoring glucose or lactate.

2.3. OLED preparation

The OLEDs were prepared as an encapsulated matrix array of 2 x 2 mm2 square pixels resultingfrom mutually perpendicular stripes of etched ~100 nm thick indium tin oxide (ITO) (the anode) and ~100 nm Al (the cathode). The organic layers sandwiched between the electrodes were deposited by thermal vacuum evaporation, as described previously [8, 9]. The OLEDs were operated in a pulsed mode with typically a forward bias of ~20 V and a pulse width of 100 with a repetition rate of 50 Hz.

2.4. Analyte detection

To evaluate the Bio-CD-based sensor in terms of flow, mixing, operation of the valves, and response time, the analytes were monitoring in sealed cells as described next. In the initial studies, the analytes (enzymes) were placed in the reaction chamber and the enzymes (analytes) in the reservoirs. Upon rotation at ~1000 rpm the valve bursts and the buffer solutions of the enzyme and analyte were mixed. The reaction was performed in sealed cells, where dissolved oxygen (DO) is not replenished by ambient oxygen, to simplify the analysis and enable the use of a modified Stern-V olmer equation [5], which directly correlates the final DO level with the initial analyte concentration. Analyte concentrations were in the range of 0-0.5 mM and the final volume contained in the reaction chamber was 200 (i.e., full capacity). As previously reported [10], smaller volumes can also be used, though the signal intensities are reduced, as expected. Control experiments were also designed for this study: the buffer, buffered GOx, and a mixture of buffered GOx and the oxygen-sensitive dye were characterized in terms of detected signal intensity and/or PL decay time.

3. RESULTS AND DISCUSSION

3.1. Surface energy manipulation

Fig. 3 demonstrates the flow of blue colored DI water in the surfactant coated PP bio-CD. Pictures a to d were captured consecutively at 2 s intervals following injection of the solution (see Fig. 2). As seen, within 8 s, the water flowed from the burst valve to the reaction chamber. This indicates that the surface modification increased the surface energy within the channels and chambers embossed in the CD, since prior to this treatment we did not observe flow under similar experimental conditions.

Fig. 3. Hydrophilic polypropylene CD surface prepared by using a surfactant based chemical coating. Images …a? through …d? were captured at 2-second intervals.

To quantify the effect of the surface treatment, the DI water contact angle was measured for untreated and treated surfaces. It was found that each of the two coatings increased the surface energy of the PP material from 29 dynes/cm to ~48 dynes/cm. Although the exact composition of the coating is proprietary, it is believed to contain an ethylene vinyl alcohol copolymer mixed with a surfactant, such as polyethylene glycol diolate, nonylphenoxypoly(ethyleneoxy) ethanol, triethylene glycol divinyl ether, and their combinations [11].

3.2. Analyte monitoring

In preliminary results, a solution of 18.2 mg/mL of the oxygen-sensitive dye tris (4,7-diphenyl-1, 10-phenanthroline) Ru chloride (Ru(dpp)) was used together with a blue OLED based on 4,4?-bis(2,2?-diphenylvinyl)-1,1?-biphenyl (DPVBi) to monitor glucose. The effect of the solution volume (which was then limited to a maximal value of 15 μL) in the reaction chamber on the PL signal intensity was monitored. It was shown that the signal intensity is generally proportional to the volume of the reagents mixture with a resolution of 3 μL under the non-optimized experimental conditions employed. A correlation between the integrated signal intensity and the glucose concentration was obtained. However, as the background light that reaches the PD, which stems mostly from the tail of the electroluminescence (EL) of the OLED, reduces the detection sensitivity, we opted to monitor the PL decay time (τ), which is insensitive to the EL as the analyte is monitored during the off time of a pulsed OLED and the EL decay time is < 100 ns.

It was possible to follow τ by using PtOEP, whose τ ranges from ~95 μs in oxygen-free

atmosphere to ~5μs in 100% O2 . In aqueous solution at ~23oC, where the DO level is ~8.5 wt. ppm, τ is 25-30μs, depending on the sensor film. Under comparable experimental conditions and sensor films, the calibration lines for glucose and lactate are similar as the DO level is monitored. However, each analyte is oxidized only in the presence of its specific enzyme and therefore there is no interference between the analytes.

As shown by Cai et al. [5], calibration lines for experiments performed in sealed containers obey a modified Stern-V olmer equation:

where I0 andτ0 are the unquenched PL intensity and decay time, respec tively, and KSV is a film-and temperature-dependent constant.

Fig. 4 shows the modified Stern-V olmer plot for the preliminary results obtained using a sensor based on an Alq3-based OLED, a PtOEP:PS sensing film that served as the bottom of a reaction chamber in the PP CD, and a PMT PD. The reagents were mixed by opening a burst valve.

Fig. 4 . Modified Stern-V olmer calibration line for lactate in a sealed cell at 23℃monitored on a

PP CD-based sensor.

following rotation of the CD. The measured values of t were ~30, ~50, and ~90us for solutions at equilibrium with air at ambient temperature, a solution containing 0.1 mM lactate, and a solution containing 0.25 mM or larger analyte concentrations, respectively. The latter are solutions devoid of oxygen, which was completely consumed by the oxidation reaction, and as the reactions were performed in seals cells

the in-diffusion of O2 from air was minimal. These preliminary results show the promise of the combination of the OLED-CD approach. The addition of a thin film PD is expected to lead to compact, field-deployable devices.

It was also possible to measure four different concentrations of glucose by utilizing the four separate sections on the CD. The four sections should enable also simultaneous detection of different analytes on the same CD when using a compatible array of small size photodiodes.

4. CONCLUSIONS

In summary, the advantage of the OLED-lab-on-CD based biosensing platform is its simplicity in terms of fabrication, integration, and usage, as well as its versatility in terms of (compact) size, design, and multianalyte detection feasibility. Additional studies are needed to optimize the sensor in terms of e.g., response time and analyte volume and flow. Additionally, the sensor should be optimized for simultaneous detection of multiple analytes and integration with a thin film PD that will lead to the envisioned compact device.

ACKNOWLEDGEMENTS

Ames Laboratory is operated by IowaStateUniversity for the US Department of Energy (USDOE)under Contract DE-AC 02-07CH11358. This work was partially supported by NSF, ISU, and theOffice of Basic Energy Sciences, USDOE.

外文资料译文

基于OLED的聚丙烯生物光盘传感器

Srikanth Vengasandra, Yuankun Cai, David Grewell, Joseph Shinar, and Ruth Shinar

农业与生物系统工程中心

阿姆斯实验室-USDOE和物理学与天文学系

微电子研究中心、电子与计算机工程系，爱荷华州立大学，50011，美国

摘要

为了研制出微流平台的传感器设备，通过超声波加热把微型无闪光设备表面附上聚丙烯薄膜，得到生物光盘传感器设备。压制的特点是设计了储液槽、阀门、反应室，其共同组成了结合紧凑的遥感平台，通过一台个人电脑光盘播放器监测分析物水平。为了使传感器结构紧凑，我们选择光致发光为基础基于OLED感应平台的检测乳酸和葡萄糖。镀膜之后，表面能量的塑料基板表现出亲水性。试剂放置在单独的位置，通过通道向反应室由光盘旋转混合。在适当氧化酶催化作用下，通过测量相关的溶解氧水平对发光衰减时间的氧气敏感染料的影响来测量乳酸或血糖浓度。结果表明有机电致发光器件作为基于光谱激励源传感器，微流光盘平台有很大潜力，可以同时分析多种分析物。

关键词：有机发光二极管，发光二极管，基于光盘实验，葡萄糖传感器，乳酸传感器

1．简介

生物微机电系统（微机电系统）为基础的感应平台在制作塑胶基板上有很大的潜力，例如成本低、可一次性使用、无交叉污染、较好敏感性。此外，这种传感器也表现出较高一次性监测量和多种分析物检测的能力。这样一个基于光盘实验的传感平台有很大优势，它使用范围广并且可简单地解决多种问题，可以分析检测多种分析物，并且尺寸更紧凑。此外，检测通道很容易通过这种光盘实现。

本文介绍了使用超声波压印聚丙烯在光盘里产生的微流通道、阀门、储层、反应点。超声微压印是局部热的来源，这种技术的优点包括循环时间短，易压印和低残余应力。此外，这种方法适用于批量和连续生产，它是注塑成型和热压相对简单并且更常见的制作方法。微流控光盘结构产生材料是聚碳酸酯、聚苯乙烯）、

聚二甲基硅氧烷（硅橡胶）和聚丙烯。

就像这项工作中所表现的那样，该光盘可以与光致发光为基础的有机发光传感器感应平台集成在一起，来构造一个紧凑的监测装置，例如检测乳酸和葡萄糖。光盘材料通常具有较低的表面能使它们具有疏水性。然而，许多功能微设备依靠亲水性能使通道的墙壁可以促进微通道允许适当的流体流动。因此臭氧和等离子体等常用来化学处理增加表面能的氧化性。

在示例中，在薄膜中添加敏感材料来对组分中葡萄糖和乳酸进行监测。葡萄糖和乳酸在特定的酶作用下氧化，即葡萄糖氧化酶（葡萄糖）和乳酸氧化酶（液氧），还有氧气。氧在这样的反应中消耗，通过计算在特定的实验条件的最终溶解氧来对分析物浓度做初步分析。

酶也可以被嵌入在一个薄膜中，它们也能够溶解在溶液中。消耗在氧化反应的结果增加发光强度和发光衰减时间的氧气敏感的染料。在初步测量所示，该像素阵列和传感膜结构集成在玻璃基板上，对他们进行单独制造使它们背靠背挨着。发光监测使用光电倍增管，它是小型硅光电二极管阵列，兼容设计的有机发光二极管像素也可用来使其更紧凑，传感器可实时调试。也可通过整合发光激发传感膜、薄胶片探测器构造一个更紧凑的传感器。这种综合布线系统基于非晶或纳米晶硅，目前正在发展中，然而目前发展的速度缓慢，不允许监测氧的模式。有机综合布线系统也适用于这种融合和可能，测量合适的发光体。

2. 实验过程

2.1 材料

发泡板材挤出聚丙烯，来自Trexel公司，发泡的特定水平大约是15%到20%。聚丙烯板厚度约5毫米大小和标准尺寸的光盘一样，聚丙烯光盘用于超声波压制。氧气敏感的染料，十乙基卟啉钯，它溶解在聚苯乙烯（分子量45000）中用来镀膜；葡萄糖氧化酶，液氧（来自毕赤酵母公司），葡萄糖和乳酸均购自Sigma-Aldrich；所有试剂溶解于磷酸盐缓冲盐水溶液；有机电致发光器件的激发源是基于Alq3。

2.2 准备光盘

轧花是在泡沫材料使用2000系列超声波系统，运行在20千赫使用钛角产生的微观特征，超声波压花进行40峰峰值振幅和0.5秒的加热时间，确定这些条件的依据是以往的研究，期中一个截面的典型的特征如图1。图像的左边是一个压印的微孔泡沫衬底；相对图像右边是一个标准的压印特征，就像前面指出的那样，

通过使用发泡基材，不受欢迎的闪光减少，因此这是唯一结构合适的发泡基材。

图1：截面的典型特征：

左：轧花聚丙烯发泡产生闪光微观形态。右：浮雕的特点，在标准和产生的闪光。

图2显示了微通道、储液槽和通道产生的葡萄糖/乳酸生物光盘。该光盘包含四部分，分别同时检测不同浓度葡萄糖。汽门通过使用脉冲阀获得，也就是说，阀针的液体在管道内被强制扩大在微通道，压力通过一个快速高速的旋转产生，这取决于旋转速度和通道的大小，克服毛细管压力使流体流动。

为增加聚丙烯表面能量，提出了一种方法，在有机硅表面添加活性剂和专有的有机系统。对每一个产生的化学涂料都进行单独研究，都产生预期的结果。该涂层解决方案来自于古斯顿技术公司，稀释去离子水比例为1：10，直接应用到聚丙烯生物的光盘样品基质在湿润的条件下清洁处理15分钟。表面处理后用去离子水冲洗并干燥的空气吹干。计算以适当的角度来预处理和处理的样品表面，以增加表面能。

为了能够监测光谱，反应室的光盘被切断，一个传感薄膜沉积在一个薄玻璃基板附近。这种置换到一个透明的反应室底部，它能够使光电探测器在几何位置的后侧检测光谱。在背后的有机发光二极管阵列检测穿过有机电致发光器件阵列的光谱信号。从生物光盘反射回来被俘获的光致发光信号通过滨松R6000光电倍增管测量，工作电压为900 V。

图2.示意图（不按比例）的生物光盘组成监测葡萄糖或乳酸。

2.3 OLED 准备

OLED是一个封装的矩阵阵列的2×2平方毫米正方形像素，以相互垂直的条纹蚀刻100纳米厚的铟锡氧化物为阳极和100纳米厚铝为阴极。有机层之间夹在电极之间，通过沉积热真空蒸发镀膜，具体如前面描述。OLED的运行的脉冲模式正向偏置电压20伏特，脉冲宽度为100，重复率为50赫兹。

2.4 分析物检测

评价基于生物光盘传感器在流动、混合、阀门的操作和响应时间，监测分析物需要各自密封，接下来将做描述。在最初的研究中，分析物（酶）被放置在反应室，酶（分析物）放在储液槽。在1000转每分钟的转速下缓冲溶液中的酶和分析物混合。反应是在密封的小隔间中并且缺乏溶解氧并隔绝氧气的环境下进行，以简化分析，依据Stern-V olmer方程可直接计算出初步分析物浓度。分析物浓度范围在0-0.5mM，最后的量在反应室中为200（即充分反应）。如以前报告所述，小反应量也可以测量，信号强度会向预期一样降低。研究中实验控制变量为缓冲区、缓冲、混合缓冲葡萄糖，氧气敏感的染料的影响体现在检测信号的强度和发光衰减时间。

3．结果和讨论

3.1 表面能操纵

图3表明流动蓝色的去离子水的表面活性剂涂层聚丙烯生物光盘。图片a到d 是在2秒的间隔连续捕获的.注射后的结果（参见图2）。可以看出，在8秒内，水流从管道到反应室。这表明，表面限制增加了范围内的和压制在光盘上的小格子的表面能量，因为在此之前类似的实验条件下我们并没有观察到流动。

图3. 亲水性聚丙烯光盘表面采用表面活性剂为主的化工涂料

图像a到d显示每2秒时间间隔

为了量化表面处理的影响，定量计算水和处理、未处理表面的接触角。结果发现，每种涂料增加了表面能的聚丙烯材料从29达因/厘米~48达因/厘米。虽然确切组成的涂层是专有的，它被认为含有乙烯乙烯醇共聚物与表面活性剂混合，如聚乙二醇、氧化乙烯、乙醇、三乙二醇二乙烯基醚以及它们的组合。

3.2 分析物监测

初步结果，一种方案以18.2毫克/毫升的氧气敏感的染料三氯化钌和基于2,2'-二苯乙烯-1,1'-联苯的蓝光有机电致发光器件监测血糖，监测实验用量（最大值为15μL）对结果的影响。结果表明，信号发光强度和所用试剂的量成正比例关系，并且在相同的实验条件下3μL的体积取得最佳的发光效果，获得了完整的信号强的和葡萄糖浓度的关系曲线。然而，由于背景光影响光电探测器在一定程度上影响了试验器件发光光谱，降低了检测灵敏度，我们选择监测发光衰减时间（τ），

因为它受有机电致发光器件脉冲发光的影响较小，光谱脉冲发光延迟时间小于100纳秒。

接下来用PtOEP测发光时间t变化，在无氧环境下延迟时间t为95μs，在纯氧中延迟时间为5μs，具体取值取决于传感器的薄膜。通过实验对比实验条件和传感器薄膜，测得的葡萄糖和乳酸量和预期水平一致，由于每个分析物只在特定的酶作用下氧化，因此实验在各组分之间不存在干扰。

根据资料，密封状态下实验校准线的计算满足Stern-V olmer方程：

期中τ0是淬火的发光强度和衰减时间，Ksv是一个和薄膜、温度相关的常量。

图4显示了修正后的Stern-V olme图形，基于Alq3有机电致发光传感器，溶解在聚苯乙烯中的PtOEP作为传感薄膜镀在了光盘的上面，上面还有一个光电探测器，试剂通过打开一个通道混合。

图4. 基于光盘的传感器监测在23℃密封状态下的修正后的Stern-Volmer校准线

随着光盘的转动，测量时间为30us，50us，90us，并且随空气和温度稳定；变量设计为0.1mM乳酸和含0.25mM或更大的分析物的浓度。后面是在没有氧气的情况下的结论，这是完全消耗的氧化反应，在密封环境下空气中气体干扰最小。这些初步结果表明各种结合的可能性，另外一个镀膜的光电探测器体积要精巧，以满足各种设备。也可以测量四个不同浓度的葡萄糖利用在光盘上四个独立的空间，这四部分应使在使用小尺寸的光电二极管阵列时兼容。

4．结论

总之，基于OLED在光盘上实验的传感器的优势是制造时的简单精巧，容易集成和使用，同时它的紧凑性能够使用在设计和分析多种分析物上。另外在研究和优化传感器上仍需继续深入研究，比如考虑到反应时间和样品的体积、流量。此外，可进一步优化传感器，使其可检测多种分析物，这样可以设计出更紧凑的装置。

致谢

阿姆斯实验室属于爱荷华州立大学为美国能源部，合同号DE-AC 02-07ch11358，项目工作的部分支持来自美国国家科学基金会和国际基础能源科学。