一个完全植入式无线压力监测系统的开发

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1、This article comes from：Biomed Microdevices (2009) 11:259264DOI 10.1007/s10544-008-9232-1一个完全植入式无线压力监测系统的开发摘要：一个完全植入式无线压力传感器监测系统的开发是为了检测体内膀胱的压力。该系统包括一个小型的商业压力模具经导管连接到电子放大设备，一个微控制器，无线发射器，电池和个人数字助理（PDA）或接收无线数据的计算机。该传感器是完全植入的，并且每秒传送一次压力数据，压力检测评估范围为1.5 psi，分辨率为0.02 psi。在体外，设备的校准测量表现出高度的线性度和良好的时间响应。该装置植入

2、一些猪的体内进行研究，历时超过3天。该系统能适用于其他方面的压力读数，以及其他生命体征的测量，它迈出了一个无处不在的发展远程医疗和远程病人监控平台的第一步。关键词：微机电系统、压力传感器、植入、患者的监测、遥测、远程医疗、膀胱、无线1介绍：目前，医学界对在家和在医院的病人的远程医疗和远程监控产生了浓厚的兴趣。当前，病人监护仪器和实现手法是繁琐和局限的。例如，在重症监护病房，血压监测可用动脉线进行连续监测。这是一个放置在动脉的导管，并且外部传感器能检测到压力。它的局限性是高度可变的准确性，患者往往服用镇静剂，以防止因运动对自身造成伤害。另一方面，在标准的基础护理中，虽然对病人是完全非侵入性，并且

3、免除了他们的负担，但是，病人血压标准测量与非连续点的测量通常采取每2-12小时一次。重要的生命体征测量值间的显影可能被错过。目前，还没有连续监测生命体征的，并且可提供无极端入侵和或麻烦的，如同临床医生的设备。一个连续的，能够实时检测的，并且监测没有显著减少病人的舒适或限制其自身的运动的设备将在性能和舒适的集约化和标准治疗上缩小差距。一个简单而划算的的解决办法是利用无线遥控可植入微系统。无线遥测技术释放了被束缚的患者，使他们脱离了大医院的监控，他们可以参加医院无线传感器网络，该网络通过最大限度地减少人员的工作负荷，增加获得的数据量，并简化其存储和处理，进而可能会提升监控效率。在大多数情况下，无线

4、植入式压力传感器发展的重点是由无线电供电设备射频（RF）感应，使无限期植入和手术操作时没有必要交换电池。此外，设备总体积最小化，因为电池通常是最大的组成部分。一些团体已经开发和测试出能够检测股动脉动态血压或动物模型主动脉的设备（纳杰菲和Ludomirsky2004； Schlierf等。2007年）。然而，传输范围通常局限于厘米（纳杰菲和Ludomirsky2004；Schlierf等。2007）和传感器射频感应时只能传输数据。这往往限制了离散时间点的测量或在病人身上装上天线在所有的时间连续测量（CardioMEMS2007；沃尔顿和克鲁姆 2005年）。这里，我们提出了一个不同的方法来监测流

5、动压力，其中包括：微压模具，电子放大器，微控制器，无线发射器，电池，和通讯的个人数字助理（PDA）或计算机。2方法和材料：压力传感平台分为三个部分：压力传感导管导线；传感器节点（图1）；PDA或电脑接收器。以上各部分的制造将在下面部分讨论：2.1导管导线：该导管导线装载了压力传感器，并且连接到传感器节点上（图2）。它由一个用紫外线环氧树脂（Masterbond UV10）印刷在测量0.650.65毫米的陶瓷印刷电路板（PCB）的压阻压力传感器（硅微5108）构成。模具设计参考了它的大小，感应距离，精确度，灵敏度具有1.6 mV/ psi的音频/视频。芯片上的应变片配置在温度补偿的惠斯登电桥。然

6、后，该芯片（西邦德7402C）焊接到基板焊盘上。四个单独绝缘铂铱（铂铱），通过一个7.5French（2.5毫米）的电线焊接到连接到接触焊盘的电路板。紫外线环氧树脂应用于基板焊盘和芯片上，并进行8分钟固化，以防止任何铂铱电线或wirebonds破坏连接。有四个楔形孔金帽削减了紫外线对环氧树脂印刷电路板压贴的压紧膜，以保护芯片。四铂铱线被缠绕在一个高强度聚酯绝缘的磁芯，并通过聚乙烯管材。该装配用硅胶填充焊接螺纹管。图1图1：在设备充分植入并充分包装后，该压力传感器安置在7.5French的导管中，它直接植入膀胱或腹腔。导管的另一端连接到传感器节点，这些点由单片机和无线发射器，电子产品放大器和电池

7、组成。该设备是包裹在LDPE薄膜上并用医用级硅橡胶制造成型。图2图2：艺术家的设计引领导管导线技术的尖端。（a）给出了无包装的领先技术，商业压紧膜焊接到PCB板上。四铂铱通过导管电线焊接到PCB板上。（b）描绘了包装的领先技术。金盖板，保护芯片和丝焊。缠铅玻璃纸; PDMS塑造其周围，以增加生物的相容性。2.2传感器节点：该传感器节点由三个部分组成：电子放大器，微控制器和无线发射器，电池。导管导线的一端焊接到一个特定的电路板上。这个电路板是四元组微功耗，单电源运算放大器（德州仪器TLV2764），一个2.5伏的电压调节器芯片（模拟装置REF192）和一个单刀双掷（SPDT）磁簧开关打开和关闭设

8、备（哈姆林）（林2007）。芯片组的电压调节器的电源电压供电的设备和其他电子产品电压设置到2.5 V。以防止因电池电压的变化引起压紧膜信号的变化。运算放大器被配置为无任何偏移量将传感器电压从放大桥放大到300倍。生理相关的压力测量范围为1.5 psi的表压，并与器件的灵敏度，电源电压和放大器，设备输出为1.2伏/ psi和生理压力范围为1.8V。该放大电路的输出被连接到微控制器和无线发射器（Mica2Dot（Crossbow MPR510CA），以下简称为the dot mote），其中发射功率为433兆赫。我们编程的微控制器以获取和传输数据，同时最大限度地提高电池的寿命在以下三个方面：首先，

9、微控制器使传感器的测量脉冲周期为30微妙，在这之后整个设备进入睡眠模式。第二，测量时只有每秒一次。最后，由于传输消耗最大电量，采样的数据存储在本地the dot mote中，并且每测量30次传输一次（林等人。2007）。这些技术是能量消耗从3兆焦耳降到625J（林2007年，林等人2007年）。电池使用的是3.7伏，850毫安时锂聚合物电池（电池美国）。该装置的使用寿命是387300次测量，或者在电池电压低于电源电压情况下，大于四天采样率不变。一旦组装完成，传感器节点包裹25微米厚的低密度聚乙烯（LDPE，塑料薄板供应）和压缩，再PDMS成型。此后，该装置浸入聚二甲基硅氧烷的第二硅层，以堵塞第

10、一层硅橡胶的任何漏洞。期间和之后的包装过程中，电池无法充电或更换，所以是钕磁铁被堆放在模具上并激活磁性开关和在它24小时治愈后关闭设备。2.3无线通信：该点配有互补的通信接收机基站(Crossbow MIB510CA)，该基站连接到计算机。发送的dot mote是十六进制格式的，其中包括一个时间戳，一个独特的ID标签，剩余的电池电压，放大的压力数据。 LabVIEW的（美国国家仪器公司）的程序来读取和转换成数据包，该数据包保存在一个文本文件中，并能实时绘制图形曲线。2.4体外试验：一旦一个导管导线制作完成，每根导线都要进行单独测试，并且将其放置在密封压力腔中测试其性能，这是通过压力腔盖电连接到

11、电线上。导线外部供电（安捷伦E3630A）和高精度万用表测试输出电压（吉时利2000年）。压力是在大气压力恒定下举行了30分钟，而生物医学Microdevices的（2009）NISTcalibrated压力表（欧米茄DPG5600B-30A）11:259-264261器件的输出和压力读数每5分钟记录一次。体腔连接到了一个压缩的氮气瓶，通过压力调节器将压力上升到1.0 或1.5 psi并保持三十分钟。对于第一个十分钟，压力和电压读数被全部带走，在后来的二十分钟，每五分钟检测一次。压力容器释放出大气压力和电压，前十分钟，压力彻底读出，在后来的二十分钟，每五分钟读取一次。每根导线校准三次，以测试环

12、境对导线及其包装的影响。他在空中检测，首先作为一个控件，然后导线的尖端放置在容器的烧杯的水中，开始测试之后，让它在水下保持四天。如果四天后输出的量值和时间响应和以前一样，这就搭配了一个传感器节点。经过一个传感器节点配对的导管导线，整个装置进行了包装前的再次测试。导线被放置在密封的水压力容器。该传感器节点连接到连接外部压力腔的导线，并有电池或者DC电源供电。LabVIEW计算机程序运算和存储无线压力数据。压力是从0-1.5psi逐步加大以测试实验所需压力范围，每步保持五分钟，每两分钟进行一次压力读数。为了检验该设备分辨率，增量为0.02 psi的压力变化从0到0.1psi和0.1psi的增量从0

13、.1到1.5psi。在每次试验结束后，标准曲线生成相关的电压数值通过电脑记录到压力腔内。另一个测试封装设备是将它淹没 2.5 gal 染色的水中并传输数据，直到电池耗尽。之后进行数据分析，以寻找任何短路的迹象，描绘任何传感器漂移现象和量化设备的全寿命。包装后来被去除，看是否有漏水的痕迹。2.5 体内测试：经加州大学洛杉矶分校医学中心 IRB # 2004年-185-11批准，成年母猪被用于作为体内测试。一个设备植入膀胱，另一个放置腹腔内作为参考。当传感器节点被放置在皮下组织中时，导管的导线被放在那些场所。手术后，猪被关在动物园的围栏中，从而使在围栏外的计算机和设置好的无线电接收器收集数据。在这

14、种情况下，猪都有充分的意识到和动态。在接下来的 2-4 天，猪被处死，然后分离出设备。简单进行尸检，以寻找组织炎症或任何免疫反应的有机硅包装。然后检查设备是否有任何损坏，泄漏或如有必要的话，再寻找任何其它故障点。3结果及讨论：组装完毕的铅导线和传感器节点的测试显示了传感器的快速线性响应 (1.34 psi/V）。在初次测试之后，对设备进行拆卸和重新组装。观察到偏移量略有改变。一旦该设备完全包装和准备植入，这种变化通过计算的压力校准的比较曲线和压力表的实际压力和调整校准曲线的偏移值进行补偿。在测试聚二甲基硅氧烷包装的完整性时，设备在不失灵的情况下运作，直到电池设备使用107 h而耗尽。在前两天半

15、时间内，输出的电压变化小于0.0003。然而，在接下来的 2 天，电压稳步下降直至设备停止运作。一旦停止运作，设备就检测出无液态水或气态水，或水的PDMS渗透层，这样，数据就不会因短路丢失了。4结论：总之，我们已建立了完全植入体内的无线压力传感器，其在短期内应用于泌尿系统的研究和病人监护仪上。体外测试演示其快速的时间响应和其高线性。通过膀胱和猪腹腔模型的体内试验，压力传感系统能够成功记录医学相关的数据，其中包含像排尿这种生理活动。这个平台可以扩展其他的传感模式，如测核心温度的热敏电阻，白金和银电极测血液或组织氧气电压，铅植入动脉以获得心率血压分析参数。导管进一步小型化到点，这样，它能放入针中，

16、进而可以消除人体对大手术的需要。为了这次试验，电子和无线传输单元被保存在内部，以免动物对其造成损坏。对于人类的应用，它更实用的做法是将设备固定在身体表面。在医院中，像它这样，传感器平台的发展将会更加实际和普遍。附件2：外文原文This article comes from：Biomed Microdevices (2009) 11:259264DOI 10.1007/s10544-008-9232-1Development of a fully implantable wireless pressuremonitoring systemAbstract： A fully implantable

17、 wireless pressure sensor system was developed to monitor bladder pressures in vivo. The system comprises a small commercial pressure die connected via catheter to amplifying electronics, a microcontroller, wireless transmitter, battery, and a personal digital assistant (PDA) or computer to receive

18、the wireless data. The sensor is fully implantable and transmits pressure data once every second with a pressure detection range of 1.5 psi gauge and a resolution of 0.02 psi. In vitro calibration measurements of the device showed a high degree of linearity and excellent temporal response. The impla

19、nted device perfored continuously in vivo in several porcine studies lasting over 3 days. This system can be adapted for other pressure readings, as well as other vital sign measurements; it represents the first step in developing a ubiquitous sensing platform for telemedicine and remote patient mon

20、itoring.Keywords： MEMS . Pressure sensor . Implantable . Patient monitoring . Telemetry . Telemedicine . Bladder .Wireless1 IntroductionThere has been significant interest in the medical community in telemedicine and remote patient monitoring at home and in the hospital . Current patient monitoring

21、instrumentation and practices can be cumbersome and restrictive. For example, in the intensive care unit, blood pressure monitoring can be monitored continuously with an arterial line. This is a catheter that is placed in the artery, and an external transducer detects the pressure. The limitations o

22、f this are that the accuracy is highly variable, and the patient is often sedated to prevent him from injuring himself from movement. On the other hand, in standard floor care, while completely non-invasive and burden-free to the patient, standard blood pressure measurements with a cuff are non-cont

23、inuous point measurements typically taken every 212 h. The development of critical vital signs between measurements could be missed. Currently, there is no device which provides clinicians with continuous monitoring of vital signs without being extremely invasive and/or cumbersome.A device capable o

24、f continuous and real time measurement and monitoring without significantly reducing the patients comfort or restricting his movement would fill the gaps in performance and comfort between intensive and standard care. A simple and cost effective solution is to utilize implantable microsystems utiliz

25、ing wireless telemetry. Wireless telemetry frees the patient from being tethered to large hospital monitors and can participate in a hospital sensor network, which could increase monitoring efficiency by minimizing staff work load, increasing the amount of data obtained, and streamlining its storage

26、 and processing.For the most part, wireless implantable pressure sensor development has focused on devices powered by radio frequency (RF) induction, which enables indefinite implantation and operation without the need for subsequent surgeries to exchange batteries. Also, the total device volume is

27、minimized, as the battery is typically the largest component. Several groups have developed and tested devices that detect dynamic blood pressure in the femoral artery or aorta of animal models (Najafi and Ludomirsky 2004; Schlierf et al. 2007). However the transmission range is often limited to cen

28、timeters (Najafi and Ludomirsky 2004; Schlierf et al. 2007) and the sensor can only transmit data when it is exposed to RF energy. This often limits the measurements to discrete points in time or tethers the patient to an antenna at all times for continuous measurements (CardioMEMS 2007; Walton and

29、Krum 2005).Here we present a different approach to monitor ambulatory pressures, which consists of a micromachined pressure die, amplifying electronics, microcontroller, wireless transmitter, and battery, and communicates with a personal digital assistant (PDA) or computer.2 Methods and materialsThe

30、 pressure sensing platform is divided into three parts: the pressure sensing catheter lead, the sensor node (Fig. 1), and the PDA or computer receiver. The fabrication of each of these parts is discussed below:2.1 Catheter leadThe catheter lead houses the pressure sensor and connects it to the senso

31、r node (Fig. 2). It consists of a piezoresistive pressure sensor (Silicon Microstructures 5108) measuring 0.650.65 mm affixed onto a ceramic printed circuit board (PCB) with UV epoxy (Masterbond UV10). The die was chosen for its size, sensing range, and precision, with a sensitivity of 1.6 mV/psi/V.

32、 The strain gauges on the die are configured in a temperature-compensated Wheatstone bridge. The chip was then wirebonded (West Bond 7402C) to contact pads on the substrate board. Fourindividually insulated platinumiridium (PtIr) wires threaded through a 7.5 French (2.5 mm) catheter were soldered to

33、 leads on the board connected to the contact pads. UV epoxy was applied over all contact and solder pads on the substrate board and chip and cured for 8 min to prevent any of the PtIr wires or wirebonds from breaking contact. A gold cap with four wedge-shaped holes cut out of the side was affixed wi

34、th UV epoxy onto the PCB over the pressure die to protect the chip. The four PtIr wires were wound around a high-tensile insulating polyester core and threaded through tygon tubing. This assembly was threaded through silicone tubing prior to soldering.Fig. 1 The implanted device after fully being fu

35、lly packaged. The pressure sensor is housed at the end of a 7.5 French catheter, which is implanted directly into the bladder or peritoneal cavity. The other end of the catheter is connected into the sensor node, which consists of the dot mote (microcontroller and wireless transmitter), the amplifyi

36、ng electronics, and battery. The device is wrapped in LDPE film and molded in medical-grade PDMS.Fig. 2 Artists rendition of catheter lead tip. (a) Shows the lead tip unpackaged. A commercial pressure die is affixed and wirebonded onto a PCB substrate. Four PtIr wires fed through the catheter are so

37、ldered onto the PCB. (b) Depicts the lead after packaging. A gold lid covers and protects the chip and wirebonds. Cellophane is wrapped around the lead; PDMS is molded around it to make it biocompatible2.2 Sensor node The sensor node consists of three components: amplifying electronics, microcontrol

38、ler and wireless transmitter, and the battery. The free end of the catheter lead is soldered onto a custom-designed circuit board. On this circuit board are a quad micro-power, single supply operational amplifier (Texas Instruments TLV2764), a 2.5 V voltage regulator chip (Analog Devices REF192), an

39、d a single pole, doublethrow(SPDT) magnetic reed switch to turn the device on and off (Hamlin) (Lin 2007). The voltage regulator chip sets the supply voltage powering the device and other electronics to 2.5 V to prevent any variations in signal from the pressure die due to variations in battery volt

40、age. The operational amplifiers were configured to null any offset from the sensor bridge and amplify the bridge voltage by a factor of 300. The physiologically-relevant pressure measurement range was 1.5 psi gauge pressure, and with the device sensitivity, supply voltage, and amplification, the dev

41、ice output was 1.2 V/psi and 1.8 V for the physiological pressure range.The output of the amplifying circuit was connected to the microcontroller and wireless transmitter (Mica2Dot (Crossbow MPR510CA), hereafter referred to as the dot mote), which transmits at 433 MHz. We programmed the microcontrol

42、ler to acquire and transmit data while maximizing battery life in three ways: first, the microcontroller pulses the sensor for only 30 s each measurement cycle, after which the entire device goes into sleep mode. Second, the measurements are taken only once per second. Finally, since the greatest po

43、wer draw comes from transmission, the sampled data is stored locally on the dot mote and is transmitted every 30 measurements (Lin et al. 2007). These techniques reduce the energy consumption from 3 mJ per measurement to 625 J (Lin 2007; Lin et al. 2007). The battery used is a 3.7 V, 850 mAH lithium

44、-polymer battery (Batteries America). The device was observed to have a lifetime of 387,300 measurements or 4 days at this sampling rate before the battery voltage dropped below the supply voltage of the device.Once fully fabricated, the sensor node was wrapped in 25m-thick low density polyethylene

45、(LDPE, Plastic Sheeting Supply) and compression-molded in PDMS. Afterwards, the device was dipped into PDMS for a second silicone layer to plug any holes in the first PDMS layer. During and after the packaging process, the battery cannot be charged or replaced, so neodymium magnets were stacked on t

46、op of the mold to activate the magnetic switch and turn off the device while it cured for 24 h.2.3 Wireless communicationThe dot mote communicates with a complementary receiver station (Crossbow MIB510CA), which is connected to a computer. The dot mote sends data in a hex format that includes a time

47、stamp, a unique ID tag, the remaining battery voltage, and the amplified pressure data. LabVIEW (National Instruments) was programmed to read and convert the data packets, which are stored in a text file and graphed in real time.2.4 In vitro testsOnce fabrication of each catheter lead was completed,

48、 the lead alone was tested and characterized by placing it in a sealed pressure chamber (Binks). It was electrically connected to wires threaded through the lid of the pressure chamber. The lead was externally powered (Agilent E3630A) and the output voltage was read by a high precision multimeter (K

49、eithley 2000). The pressure was held constant at atmospheric pressure for 30 min while the device output and pressure readings from an NISTcalibrated pressure gauge (Omega DPG5600B-30A) were recorded every 5 min. The chamber was connected to a cylinder of compressed nitrogen through a pressure regul

50、ator and the pressure was raised in 1.0 or 1.5 psi increments and held for 30 min. For the first 10 min, pressure and voltage readings were taken every minute and every 5 min for the following 20 min. At the conclusion of the calibration, the pressure vessel was vented back to atmospheric pressure a

51、nd voltage and pressure were read every minute for 10 min and every 5 min for the following 20 min. The calibration was performed three times for each lead to test the effects of the environment on the lead and its packaging. It was tested in air first as a control, and then the tip of the lead was

52、placed in a beaker of water inside the chamber. It was left submerged for 4 days when it was tested again. If the output magnitude and temporal response stayed consistent after the fourth day, it was paired with asensor node.After a sensor node was paired to a catheter lead, the entire device was te

53、sted again before packaging. The lead was placed in water within the sealed pressure vessel. The sensor node was connected to the lead outside the pressure chamber and powered either by the battery or DC power source. The LabVIEW program ran on the computer and stored the wireless pressure data. The

54、 pressure was incrementally stepped up from 01.5 psi to test the required pressure range and held at each step for 5 min with pressure readings taken every 2 min. To test the resolution of the device, the incremental pressure changes were 0.02 psi from 00.1 psi and 0.1 psi for 0.11.5 psi. At the con

55、clusion of each test, a calibration curve was generated that related the voltage recorded on the computer to the pressure inside the chamber.Another test of the packaged device was performed by submerging it in a 2.5 gal bucket of dyed water and left to transmit data until the battery drained out. A

56、fterwards, the data was analyzed to look for any signs of short circuits, characterize any sensor drift, and quantify the full lifetime of the device. The packaging was later removed to look for any signs of water leakage.2.5 In vivo testsAdult female swine were used for in vivo testing as approved

57、by UCLA Medical Center IRB #2004-185-11. One device was implanted into the bladder and another was placed inside the peritoneal cavity as a reference. The tips of the catheter leads were placed in those spaces while the sensor nodes were placed in a subcutaneous pocket. Following surgery, the pigs w

58、ere kept in a holding pen in a vivarium while the computer and radio receiver were set up outside the pen to collect the data. At this point, the pigs were fully conscious and ambulatory. After 24 days, the pigs were sacrificed and the devices were explanted. A brief necropsy was performed to look f

59、or tissue inflammation or any immune response to the silicone packaging. The devices were later inspected for any damage, leaking, or any other failure points if necessary.3 Results & discussionTesting of the fully assembled lead and sensor node showed a rapid and linear response (1.34 psi/V) of the

60、 sensor. When the device was disassembled and reassembled following the initial tests, the offset was observed to change slightly. Once the device was fully packaged and ready to be implanted, this change was compensated by comparing the calculated pressure from the calibration curve and the actual

61、pressure from a pressure gauge and adjusting the offset value on the calibration curve.While testing the integrity of the PDMS packaging, the device performed without malfunction until the battery drained for a lifetime of 107 h. The output voltage stayed constant for the initial 2.5 days with a var

62、iance of 0.0003. However, for the next 2 days, the voltage steadily decreased until the device ceased functioning. An inspection of the device once it had ceased functioning showed no liquid or gaseous water or water penetrated the PDMS layer and the data was free of short circuits.4 ConclusionIn su

63、mmary, we have constructed a fully implantable wireless in vivo pressure sensor for use in short term urological studies and patient monitoring. In vitro testingdemonstrates its quick temporal response and its high linearity. Through in vivo tests in the bladder and peritoneal cavity of porcine mode

64、ls, the pressure sensing system was able to successfully record medically relevant data, which contained physiological events like urination. This platform can be expanded with additional sensing modalities, such as a thermistor for core temperature, platinum and silver electrodes for blood or tissu

65、e oxygen tension, and analysis of blood pressure to obtain heart rate if the lead was implanted into an artery. Further miniaturization of the catheter to a point such that it can fit inside a needle can eliminate the need for major surgery. While the electronics and wireless transmission unit were kept internally for this test because of fears of damage to it by the animal, for human applications, it may be more practical to keep it outside yet secured to the body. This would make deployment of a sensor platform like this more practical and ubiquitous within a hospital.