Sensors 2014, 14, 20620-20644; doi:10.3390/s141120620
OPEN ACCESS

sensors
ISSN 1424-8220
www.mdpi.com/journal/sensors
Review

Chronically Implanted Pressure Sensors: Challenges and Stateof the Field
Lawrence Yu 1, Brian J. Kim 1 and Ellis Meng 1,2,*
1

2

Department of Biomedical Engineering, University of Southern California, 1042 Downey Way,
DRB-140, Los Angeles, CA 90089-1111, USA; E-Mails: lawrence.yu@usc.edu (L.Y.);
brianjk@usc.edu (B.J.K.)
Ming Hsieh Department of Electrical Engineering, University of Southern California,
3740 McClintock Ave, EEB-100, Los Angeles, CA 90089-2560, USA* Author to whom correspondence should be addressed; E-Mail: ellis.meng@usc.edu;
Tel.: +1-213-740-6952; Fax: +1-213-821-3897.
External Editors: Andrew J. Mason and Wen Li
Received: 27 June 2014; in revised form: 14 October 2014 / Accepted: 21 October 2014 /Published: 31 October 2014

Abstract: Several conditions and diseases are linked to the elevation or depression ofinternal pressures from a healthy, normal range, motivating the need for chronic
implantable pressure sensors. A simple implantable pressure transduction system consists ofa pressure-sensing element with a method to transmit the data to an external unit. The
biological environment presents a host of engineering issues that must be considered forlong term monitoring. Therefore, the design of such systems must carefully consider
interactions between the implanted system and the body, including biocompatibility,
surgical placement, and patient comfort. Here we review research developments on
implantable sensors for chronic pressure monitoring within the body, focusing on generaldesign requirements for implantable pressure sensors as well as specifications for different
medical applications. We also discuss recent efforts to address biocompatibility, efficienttelemetry, and drift management, and explore emerging trends.
Keywords: implantable pressure sensors; implantable pressure transducers; biocompatibility;
telemetry; drift

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1. Introduction
Pressure in various organs of the body (e.g., brain, eye, heart, bladder) is highly regulated and its
value can provide an indication of patient health or disease progression. In diseases where the ability toregulate these pressures is lost, impairment of function or death may result; even at the microscopic
level, cell damage may be induced in the presence of abnormal pressure. Monitoring the absolutevalue of and variation of internal pressures allows diagnosis and tracking of the progress of
medical intervention.
In current medical practice, it is common to obtain a single patient pressure data point in the clinic.
However, this information only provides a limited snapshot of the dynamic pressure profile within theorgan of interest and does not capture peak or trough values or changes in the profile which may
provide important clues to inform appropriate intervention strategies [1,2]. Also, measurements
obtained may be perturbed by the stress experienced by some patients when visiting clinical settings [3].
Therefore advances in sensing technology to enable convenient, accurate, and continuous pressure
monitoring that can extend to settings outside the clinic may enable more effective disease management.
Noninvasive methods for measuring pressure have been explored extensively. These methods infer
pressure as tissues interact with high-energy waves imposed on the body in the form of sound orelectromagnetic radiation (e.g., light, X-ray). However, these methods do not provide adequate
precision and accuracy [4]. Catheterization enables targeted probing of pressure in specific body regions,and this invasive method grants the accuracy needed for effective chronic monitoring, currently serving
as the gold standard for patient care in many conditions. However, prolonged catheterization of bed
bound patients hinders normal, ambulatory behavior (e.g., psychological white-coat effects that shows
false positives) and the catheter breach across the skin leaves the patient prone to infection [2,5–7].
Fully implantable sensors provide targeted measurement of pressure in the body without theinfection risks posed by catheters or wires. Devices that integrate a sensing element with telemetry,archaically named endoradiosondes or transensors, have been in development for over half acentury [8–10]. Such devices enable a clinician to easily monitor the patient over the lifetime of theimplanted sensor. The rapid improvement of microfabrication technologies has enabled the productionof low cost, highly accurate sensors that may be safely implanted into patients for chronic monitoring.
Over the last few decades, there has been steady progress in the development of technologies toachieve clinically relevant implantable pressure sensors. The reader is referred to past reviews onpressure sensor types for biomedical applications for more comprehensive coverage of related
developments [11–13]. In this paper, we focus the discussion on implantable pressure sensors forlong-term pressure monitoring within the body, including their requirements and specificconsiderations for different pressure sensing applications. Finally, we highlight recent work and trends
in chronic pressure sensing development.
1.1. Challenges for Implantable Pressure Sensors
The in vivo environment presents a host of engineering challenges that must be considered in thedevelopment of implantable pressure sensors. The sensing element must be able to transduce pressure
over the range of clinically normal and abnormal values with sufficient resolution to inform treatment.

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A telemetry system is required to transmit the data to an external reader and may also serve to power
the implanted system. For achieving both, considerations of the interactions between the implanted systemand the body, including biocompatibility, surgical placement, and patient comfort are essential toachieve a feasible design. Key requirements for chronic implantable pressure sensors are discussed.
1.1.1. Calibration
The prevailing method of pressure measurement for in vivo applications is differential (gauge) inwhich atmospheric pressure is designated as the baseline or “zero” for the measurement. This reference
pressure value is not only dependent on the atmospheric temperature but a variety of other factors suchas temperature and even time of day; these variations give rise to large baseline (or zero) drift. Anothertype of differential pressure measurement is taken as the difference in pressure taken at two points,
where an example might be the pressure from outside the bladder subtracted from the pressure
inside [14]. Absolute pressure measurement is the measurement of pressure difference compared to aperfect vacuum. The challenge of creating an absolute pressure sensor partially stems from thechallenge of creating a perfect vacuum and ensuring that it remains relatively stable over an extended
period of time [15]. Currently, all in vivo pressure measurements are reported as gauge, so a calibration(“zeroing”) procedure must be performed prior to implantation in order to report a gauge measurementwith an absolute pressure transducer.
1.1.2. Integration with the Body
For chronically implanted sensors, the interaction with the body’s warm, saline environment can bea cause for concern especially for electrically active circuits and systems which are vulnerable to ioniccontamination [16,17]. The electrolyte rich aqueous environment within the body is conductive andcan cause inadvertent shorts or electrical leakage if electronic circuits are exposed. Proper hermetic
encapsulation is required to protect the electronics from water intrusion that can result in sensor drift
and device failure. Furthermore, ingress of gas (e.g., oxygen) can cause oxidation of interconnect
metals such as solders, which in turn can lead to attachment failure [18]. The materials incorporated
should also be resistant to corrosion [17].
Implantable sensors must also manage the immune response occurring at its surfaces. The pressure
sensing function may be adversely affected by the body’s action to isolate and expel any foreign
bodies [19]. Potential mitigation strategies include the selection of biocompatible materials that have a
track record of use in FDA-approved implants or designs that minimize tissue irritation [16]. Any
material or construction strategy will need to be evaluated in biocompatibility studies on the final
device. Also, the achievable performance in the presence of the inevitable fibrous tissue encapsulationor other cellular or blood-based biofouling needs to be carefully evaluated, especially if the device
operates in a mechanical mode.
In addition to the foreign body response of the body to the implanted device, the sensor must be
carefully designed to minimize any damage to tissues during implantation and due to its chronicpresence. Such disruptions may exacerbate the normal foreign body response and further worsen thechronic in vivo performance of the implant. Designs that anticipate and mitigate the immune response
can improve sensor integration and preserve long-term performance.

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1.1.3. Telemetry
Wireless telemetry can be used to provide power to and/or receive pressure data from the body.
This is commonly accomplished using radio frequency (RF)-based antennae or inductively coupledtransmission [20]. Telemetry may take on one of two forms for implantable sensors: (1) active
telemetry, where the implanted system contains active electronics (e.g., amplifiers, microcontrollers),
for which a power source is necessary, either through an onboard battery or transmitted power signal;or (2) passive telemetry, where the implanted system is completely passive (no power necessary) and isusually interrogated externally as load changes to an external coil (e.g., resonant frequency shifts of LC
tanks). For the active case, long term monitoring requires long lasting power sources, which may berechargeable. However, implantable power sources are still large in size. This disadvantage may beoverlooked in the case where the enabled signal processing can achieve improved signal stability in thelong term. Passive devices may be preferred for advantages in simplicity of circuitry on the implantsize and smaller implant footprint. In either case, issues related to communication over distances
persist; RF signals quickly dissipate in soft tissue and bone while inductive coupling requires proper
placement and alignment of internal and external coils for proper signal transfer. A large body of work
exists that describes methods to improve the coupling efficiency between the external coil andimplanted device, regardless of its position [21–31].
1.1.4. Drift and Long-Term Accuracy
To achieve reliable long-term measurements, implanted sensors should possess a stable,consistent response over their lifetime. Shifts in response should only be due to pressure and not otherinterfering factors. Signal drift encountered in sensors can be divided into two categories:(1) offset drift, where the baseline measurement slowly drifts to obscure the desired pressure
measurement and (2) sensitivity drift, where the devices face slow reductions in sensitivity overtime [32]. The causes of drift are attributed to changes in the sensor performance independent of theenvironment (i.e., material aging and mechanical fatigue) and changes in the environment in which thedevice is implanted. The in vivo environment is dynamic; pressure is just one of many parameters thatchange over time. Assuming that a device incorporates appropriate material choices and sensing
operational principles geared towards implantable applications, a drift and poor long-term accuracy
stem largely from interactions with the wet, implanted environment such as unavoidable tissue
encapsulation. For this reason, it is crucial for implanted sensors to have drift compensation circuits or
zeroing functions that can allow for correction to provide reliable measurements over the periodof implantation.
2. Pressure Measurement Principles
A brief overview of common pressure sensing modalities suitable for implantable systems ispresented here.

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2.1. Sensing Methods
2.1.1. Membrane-Based Sensors
An overwhelming majority of implantable pressure sensors utilize a simple design consistingof a membrane and a sealed cavity; the membrane element responds to and deflects underpressure [8–10,33–45]. Because of the simplicity of the design, geometry and materials have been
adapted for a wide variety of applications in addition to in vivo environments [46–50]. Furthermore,
piezoelectric or capacitive sensing of membrane deflection can be integrated easily with telemetryelectronics.
The miniaturization advantages of microelectromechanical systems (MEMS) technology can scalethese deformable membranes into a small form factor and create complete micro-scale systems that caneasily be implanted within the body. MEMS-based membrane pressure sensors most commonly use
either capacitive or piezoresistive methods to transduce the membrane deflection into an electricalsignal. In the case of capacitive sensing, an electrode is placed on the top (deformable) and bottom(rigid) surfaces of the sealed cavity. Deflection of the membrane causes changes in the capacitancemeasured between the electrode pair (Figure 1a). For piezoresistive sensing, a piezoresistor ispatterned onto the membrane surface, and deflection of the membrane is transduced into a change inresistance, usually measured via a bridge circuit.
Figure 1. (a) Schematic of operation of a capacitive-based membrane pressure sensor.
Diaphragm deflects under pressure, changing the effective distance between two parallel
plates, and thus increases the measured capacitance across the plates; (b) Circuit model ofpassive LC tank, commonly used for inductively coupled telemetry of capacitive-based sensors.

The selection of one method over the other is dictated by the collective combination of advantages
and disadvantages. Capacitive sensing allows for high pressure sensitivity and low temperature drift
compared to piezoresistive, but suffers from the need for on-site detection/amplification circuitry due tosmall active capacitances [42,51]. Arraying capacitive sensors can improve sensitivity [52,53]. In
piezoresistive-based membrane pressure sensors, the amplifier circuitry can be placed on the external
circuit, but suffers from low-pressure sensitivity and high temperature drift, requiring temperaturecompensation. The use of a passive LC resonating circuit, or LC tank (Figure 1b), allows for simple
wireless transmission of capacitive changes induced by pressure compared to wireless strategies forpiezoresistive sensors; this allows reduction in implant size and simplifies the required sensing
circuitry [54,55] (Figure 2a,b).

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2.1.2. Alternate Transduction Schemes
Other transduction schemes have also been investigated to address the challenges of pressure sensors
outlined above. To improve biocompatibility and tissue integration, alternative transduction schemes that
do not require hermetic sealing and utilize biocompatible polymers have been investigated, such as theuse of localized gas bubbles that can be harnessed for pressure transduction [56] (Figure 2d).
Alternatives to telemetry have been explored as well, such as the implementation of a microfabricatedspiral bourdon tube [57] and manometer that can be interrogated via visual inspection or through theuse of X-ray imaging [38,58]. It is also interesting to note that the nervous system utilizes strain
sensitive cells called baroreceptors; these have inspired the development of biomimetic strain sensors
for applications in pressure transduction [59,60].
Figure 2. Examples of implantable pressure sensor systems currently in development:(a) A capacitive-based sensor for measurements of intraocular pressure [39];(b) Capacitive-based sensor developed by CardioMEMS for measurements of bloodpressure; (c) Inductor-based sensor with piezoelectric energy harvester for measurementsof bladder pressure (scale bar 10 mm) [45]; (d) Micro-bubble based pressure sensor formeasurements of intracranial and bladder pressure (scale bar 100 µm) [61].

2.2. In Vivo Pressure Sensor Requirements
For chronic implanted sensors, the design is guided by a number of requirements beyond the initial
selection of a pressure transduction scheme. Select requirements are reviewed and discussed.
2.2.1. Size
The size of the implant is application dependent and should be sufficiently miniaturized so as toallow appropriate sensor placement in the body (i.e., compatible with implantation techniques andanatomical restrictions) and integration with telemetry components (which further increase the footprint).
By leveraging advances in micromachining technologies, MEMS-based sensors have been developedthat minimize the overall footprint and mass. The overall system size can be further reduced by

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combining sensors with application specific integrated circuits (ASICs) instead of discrete electronic
components. However, inductive powering coils and batteries are still relatively large and impose
limits to miniaturization.
2.2.2. Range and Precision
A number of pressures within the body are relevant indicators in various conditions and diseases.
Taken together, these pressures span the range of −10 to 200 mmHg when taken in reference toatmospheric pressure. Pressure ranges of interest, including abnormal pressure ranges, for in different
areas of the body are listed in Table 1 and graphically presented in Figure 3. Sensors should be able tomeasure both clinically relevant normal and abnormal pressure ranges. A commonly accepted
specification for sensor precision across the majority of the in vivo applications mentioned in Table 1
is a measurement with deviation of ±1 mmHg, or 5%–10% of the clinically normal range.
Figure 3. Relevant pressure ranges for in vivo pressure monitoring for diagnostic
applications [62–67].

2.2.3. Calibration
Ideally, the measured baseline pressure measurement should remain constant under constant pressure
conditions. In practice, this value can drift significantly over the lifetime of the implanted sensor. The
baseline drift in the commonly used gauge pressure measurement may be attributed to temperature,biofouling, and even static discharge effects [68–70]. As an example, the observed range for baseline
drift for a commonly used commercial sensor was ±5 mmHg, or up to 30% of the measurement range
for certain clinical applications [71,72]. Paradoxically, this is at odds with requirements for precision,and limits the utility of long-term recordings of implanted pressure sensors [5,73–75]. This issue
requires further attention to promote clinical adoption of implanted transducers.
2.2.4. Frequency Response
Nearly all fluid systems in the body are dynamic; fluids are constantly generated and drained, both
actions that contribute to the pressure profile over time. Pressure transducers should be able to resolveand distinguish normal and abnormal physiological hydrodynamics. The dynamics of internalpressures in the body vary and the frequency response of the sensor should therefore be selectedaccordingly (Figure 4). A highly responsive sensor will enable data acquisition and Fourier analysis of
waveforms, where flow and other useful biometrics may be inferred [76–78]. Sensors with poor

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frequency response and low slew rates will not properly capture transient changes in pressure, which canbe critical indications of disease state [78–81].
Figure 4. Relevant frequency bandwidths for varying pressure signals in vivo [23,32,82–85].

2.2.5. Materials and Packaging Considerations
In considering materials and packaging methods for implantable pressure sensors, several points
must be considered: (1) acceptable tissue integration (i.e., low cytotoxicity); (2) hermetic encapsulation
for protection of active circuits and (3) long-term mechanical stability in the in vivo environment. Dueto the inevitable foreign body response against the implanted sensor, device materials must be chosen
to try and limit biofouling for more reliable, long-term performance of the device [16]. Sterilization is
also an important factor to consider when choosing materials; devices should be able to withstandautoclave sterilization heat or chemical sterilization processes (e.g., ethylene oxide or plasma) [86,87].
Further, additional considerations for device design (e.g., removing sharp corners) can also benefit
integration [88]. Efforts to encapsulate devices in drug-based coatings (e.g., anti-inflammatory) havealso shown to limit the immune response [19,89]. For sensors with active circuitry, proper hermetic
packaging is necessary to impede electrolyte intrusion to prevent device failure and sensor drift. For
membrane based transducers that rely on the mechanical stability of the membrane material, material
selection must be carefully considered for long-term operation [17,90].
2.2.6. Telemetry and Circuitry
The design of electronics for implant telemetry requires careful management of a number of
tradeoffs that affect power consumption, antenna size, and transmission frequency. Electromagnetic
radiation characteristics must carefully prioritize safety and performance while addressing thechallenges of power and data transmission in the in vivo environment. Standards have been set that
limit tissue heating and radiation specific absorption rate (SAR), [91,92]. A licensed frequency bandfor communications, medical implants communications service (MICS), encourages development oftelemetry systems within 402–405 MHz, but newer studies suggest operating at higher frequency(2+ GHz) for optimum performance [21]. Dielectric heating during MRI examinations should beconsidered (especially if the implant is situated in a critical location in the body), so analysis of MR
heating should be done prior to implantation [93]. Another area of concern for implants would beinteractions during X-ray examination, which can damage microcontroller and non-volatile storage.

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3. Application Specific Requirements
In addition to the requirements listed above, there are several criteria that are application specific.
The etiology and current gold standard of measurement is detailed in the following sections for themost common clinical pressure sensing applications. Critical parameters for proper pressure sensor
operation within the reviewed applications are summarized in the following table:
Table 1. Table of relevant parameters in the development of pressure sensors for
implantable applications.
Application

Pressure Range(mmHg) [Ref.]

Pressure Resolution(mmHg) [Ref.]

Frequency Bandwidth(Hz) [Ref.]

ICP−10 to 50 [66]

1 [74]

0–30 [82]

2 [94,95]

0–30 [85]

1 [32]

0–200 [32]

1 [96]

3–5 [23,83]

Intra-urethral
Catheter--

0–15 ([84])--

IOP
Blood Pressure
Bladder Pressure
Intra-abdominal Pressure

12–22, >22
abnormal [65]50–180 [32]
10–70, ~150
during voiding[63]
0.2–16.2 [62]PackagingConsiderations
Shunt integration,catheter
Needle delivery,
contact lensCatheter, stents

3.1. Intracranial Pressure
Intracranial hypertension is an acute and chronic condition with a variety of causes that include
traumatic brain injury, aneurysms, brain tumors, hydrocephalus, stroke, and meningitis. Standard
treatment is the relief of fluid pressure through a lumbar puncture (spinal tap) or external ventriculardrain (EVD) [78], where an opening in the skull is created and cerebrospinal fluid (CSF) is drained.
For hydrocephalic patients, fluid pressure is shunted, typically to the abdominal cavity, with a surgicallyimplanted catheter. Unfortunately, shunt failures are common, highlighting the importance of monitoring
and regulation of CSF pressure for treatment and recovery [66,97–99]. Most importantly, long-termmonitoring of ICP provides a quantitative method for diagnosis in contrast to vague symptoms (e.g.,headaches, nausea). Two commonly used EVDs are the Codman Microsensor ICP (Codman & Shurtleff,Inc., Raynham, MA, USA) and Camino (Integra LifeSciences Corporation, Plainsboro, NJ, USA), whichfeature a sensor-tipped catheter that enables location-specific pressure measurement in the intracranial
space [100]. Unfortunately, long-term monitoring with the EVD confines patients to the clinic and theextended period of opening of the blood-brain barrier increases risk of meningitis and other infections(exacerbating intracranial hypertension) [4].
Implantable telemetric pressure sensors promise improved precision of measurements without thedisadvantages current technologies. However, aggressive immune response within the blood brain barrier
can lead to significant biofouling and drift in sensor response [10,74,101]. Robust sensor design shouldseek to minimize immune response and drift to be effective for monitoring of intracranial hypertension.

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3.2. Intraocular Pressure
The monitoring of intraocular pressure (IOP) is crucial for the diagnosis and monitoring ofglaucoma. The Goldmann tonometer is the prevailing standard used to quickly and noninvasively infer
IOP for diagnosis of glaucoma, but unfortunately this only provides a snapshot of the dynamicallychanging intraocular pressure and is of limited use for monitoring of disease progression [2,13,95].
Implantable sensors enable the clinician to accurately monitor IOP for patient response to treatment(e.g., assessing efficacy of surgically implanted drainage devices during the course of recovery).
Fortunately, a multitude of treatment options exist for glaucoma, and the addition of long-term IOP
information would help clinicians more accurately tailor the treatment program for the patient.
The design of implantable sensors must take into account the frequent movement of the eye, as
implant placement may damage peripheral tissue and motion artifacts can obscure measurement [13].
The sensor should be placed so as not to block the visual path. Sensor design can take advantage of
optically clear tissue that allows for visual interrogation, which obviates the use of complicatedelectronics for telemetry for the construction of a completely passive sensor. Active sensors may alsoutilize photovoltaics for long-term powering within this environment. As such, the developmentof IOP sensors has resulted in interesting designs that exploit the advantages of the optically
clear environment.
3.3. Cardiovascular Pressure
Chronic blood pressure monitoring is critically needed to monitor various conditions of the
cardiovascular system (e.g., restenosis, hypertension, and heart failure) as well as to assess the efficacy
of surgical interventions as is the case in monitoring repaired aneurysms [102]. Though the long term
efficacy of blood pressure monitoring can be justified, the current gold standards of blood pressure
measurement, namely external pressure cuffs or intra-arterial catheter-based systems fall short due tolack of patient comfort, infrequent measurement, possible occlusion of blood flow, and long-term
complications such as trauma and infection. As these approaches require the patient to be tethered tothe measurement system in a clinical setting, inaccuracies can be introduced as was observed by a“white coat hypertension” due to psychological effects [103]. The long-term monitoring of bloodpressure can be greatly improved using implantable blood pressure sensors that allow for continual
monitoring without hindering daily activities.
Two industry pressure sensor developments of note are the Remon ImPressure/RemonCHF from
Remon Medical Technology (acquired by Boston Scientific, Natwick, MA, USA) for measurement ofpressures within an aneurysm sac following endovascular aneurysm repair (EVAR) [104] as well as
the first FDA approved implantable blood pressure device, the EndoSensor/CardioMEMS HF Systemfrom CardioMEMS (recently acquired by St. Jude Medical, Saint Paul, MN, USA) for detectionof heart failure within the pulmonary artery. Given the prevalence of heart disease, large strides
have been made in implantable cardiovascular pressure sensors, however there is still room for
improvement and technology advances.

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3.4. Bladder Pressure
Urinary incontinence is a common issue within the United States (affecting 13 million adults [105])
currently diagnosed using a series of tests collectively known as urodynamics. Among one of thesetests is a cystometry study, where a catheter-based pressure sensor is inserted intraurethrally to accessthe bladder and assess pressure variations. Unfortunately, though this is the current gold standard, thiscatheterization process is both painful and non-physiological, which can compromise the accuracy of
the test results. The fact that these tests are also conducted within the doctor’s office in a 20–40 min
span can provide unreliable results and underdiagnose various conditions as pressure variations may be
obscured by psychological white-coat effects and short measurement times [106,107]. In addition,measurements are plagued by catheterization, such as disturbances along the catheter line and different
resonance and response time delays due to fluid inertia. For chronic monitoring, these catheter-typedevices are not viable as they are prone to infection and stone formation within the bladder [108].
Another effort that can benefit from implantable pressure sensors within the bladder are devices forneuromodulation, where patients can undergo electrical stimulation-based rehabilitation to either
inhibit or promote bladder activity. Chronic bladder pressure measurements can enable closed loop
interventions to provide more efficient and reliable treatment. The development of discrete wirelessimplantable sensors for the bladder allows for the necessary chronic monitoring of the pressures within
the bladder, while maintaining the patient’s quality of life.
3.5. Intra-Abdominal Pressure
Abnormalities in pressure within the thoracic cavity have multifactorial etiologies (e.g., edema,pneumothorax, sleep apnea) and benefits greatly from long-term monitoring [77]. Current methods
utilize measurements of stomach distension or suprasternal placement of strain sensors [109] to infer
pressures within the thoracic cavity, but these noninvasive approaches lack accuracy and specificity,
limiting clinical use. Implanted pressure sensors must be able to withstand significant mechanicaldisplacement of organs in the thoracic cavity during respiration.
4. Sensing Technologies and State of the Field
The focus of this following discussion will be the current state of the field in the developmentof improved system packaging, telemetry methods, and drift compensation for implantablepressure sensors.
4.1. Material Selection and Hermetic Packaging
A wide range of materials including metals (e.g., nitinol, titanium, platinum), polymers (e.g., liquid
crystal polymers [31], Parylene, silicone rubber), and various inorganic (carbon nanotubes [15]) andceramics (e.g., alumina) exhibit low cytotoxicity and sufficient immuno-inertness within the body andare therefore potential candidates for selection as biocompatible materials for sensor construction [16].
However, many MEMS sensors rely on the bulk micromachining of silicon and the use ofnon-biocompatible photoresists and epoxies such as SU-8 [31]. If the use of these materials cannot be
avoided, it is necessary to protect the body from these materials using a hermetic packaging system.

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Caution must be exercised in the selection of biocompatible polymers for pressure sensing membranes,
such as silicone [43] or liquid crystal polymers [55]. These polymers have been found to have significant
water absorption once implanted within the body (or in benchtop soak tests) that have caused
significant drift or leaks within their sealed cavities.
Beyond the materials used for the sensing element and circuit system, biocompatible materials forinterconnects are also in development as crucial parts of the system. Typically lead free, gold-tin solders
have been used [31], along with conductive epoxies that are encapsulated to prevent leaching into thebody. Hermetic packaging can prevent failure of implanted active circuitry and conversely, prevent toxic
materials from circuitry from reaching the surrounding biotic environment via leaching. A common
approach to hermetic packaging uses a sealed metal casing [14,63,89] that houses the sensing andcircuitry components of the system, or other material systems such as anodically bonded
silicon-to-glass [41,110], sintered and fused ceramic enclosures [55], or hermetic welding [110]. For
some membrane-based pressure sensors, the hermetic packaging can interfere with proper pressuretransduction to the pressure sensitive membrane. In these cases, a window within the casing is created,typically with its own membrane that allows for deflections with pressure. The pressure is thenconducted to the sensing membrane of the pressure sensor using an incompressible fluid (e.g., silicone
gel or oil) that allows for low attenuation transfer of the pressure from the external package to theactual sensor [25,111].
To provide additional hermeticity and improved biocompatibility to a system, a secondary encapsulation
typically with a silicone rubber or a Parylene coating is applied [19,39,93,112–114]. These coatings
were found to increase the biocompatibility and reduce offset drift of these implantable pressure
sensors. However, the conformality of the coating must be carefully monitored, as silicone rubber
may have poor coverage and reveal cytotoxic regions at sharp edges of the device [115]. The
addition of a silicone or Parylene coating over the membrane structure was found to reduce
sensitivity [39,57,93,112,116], and thus degrading the signal-to-noise ratio and compromising themeasurement resolution. Furthermore, PDMS encapsulation was found to negatively impact frequency
response, causing a 1–2 s delay in the pressure measurement due to increased viscoelastic nature of themembrane [117]. Epoxies have also been used to encapsulate circuitry; some must be stained dark inorder to reduce light incidence on ASICs [89] that may affect their performance.
Implantable pressure systems may take advantage of drug coatings to reduce the adsorption ofproteins and other cellular materials (e.g., blood) on the membrane that could cause poor sensor
performance. Antithrombogenic or heparin coatings for their devices allowed for reduced clotting forsensors implanted within arteries [89]. In another effort, sodium 2-acrylamido-2-methylpropyl sulfonate(AMPS) coating was adhered to the silicone encapsulation layer of their device to prevent thedeposition of salts on their deflecting membrane to maintain consistent sensor operation [19].
Ideally, the pressure sensor format should minimize the need for bulky, hermetic packaging andutilize a material set that appropriately balances sensing and biocompatibility requirements. The
micro-bubble transducer sets out to achieve this goal by eliminating the hermetic packaging around thesensing element (bubble) and utilizing a transduction mechanism (electrochemical impedance) thatleverages the wet environment. The exclusive use of Parylene C and thin-film platinum reduces thecomplexities of multi-material implantable system that necessitates biocompatible encapsulation [61].

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4.2. Telemetry
An in-depth review of various active and passive telemetry methods for implanted sensors is beyondof the scope of the present work, for a further information the reader is directed to the following
publication [20]. Both active and passive telemetry methods have been explored that allow for
implanted pressure sensors to transmit sensor data outside of the body for monitoring or to allow
external units to send in power signals for active circuitry. In looking at the challenges associated withtelemetry for chronically implanted pressure sensors, two issues arise, namely: (1) powering methods ofthe implanted sensor and system for continual use in the long term and (2) improving the coupling
between the implanted and external coils as well as reducing the attenuation between them for more
efficient power and data transfer. Approaches within literature specific to implanted pressure sensors
that address these two issues will be presented within this section.
Currently, to transmit power to application specific integrated circuitry (ASICs) that allow for signal
processing within the implanted device, two main methods are employed: battery-based approaches thathave a limited lifetime, and RF based approaches that require coupling efficiency. For devices intended
for acute monitoring (e.g., 1 week [19]) or for applications where the data transmission frequency is
selected to maximize range instead of requiring a certain frequency for power transmission [14], abattery-based system can be ideal. To minimize power consumption and extend battery life, one approachis to use circuitry that turns the system on for 10 s every 5 min, in essence creating a “sleeping mode”
for the sensor to conserve battery life [118]. However, these methods do not provide truly continuous
monitoring of pressure and are appropriate when occasional sampling is sufficient.
For longer implant durations (months to years), an RF-based inductive coupling approach isnormally used to transmit power signals between implanted and external coils. The coupling efficiency
between the implanted and external transmitter/receivers is of concern, as the RF signal needs to becontinuously transmitted in order for the sensor to operate in a continuous mode. Also, an externalbulky transmitter must be continuously worn. In an effort to reduce the bulky equipment required forcontinuously monitoring systems to improve patient quality of life, a hybrid system was createdthat combined the battery and RF charging components [26,27,83]. This implantable RF-based
rechargeable battery design allows for the sensor to operate and transmit signals using a battery during
the day eliminating the need for an externally worn transmitter, and then recharge the battery at nightusing an external coil within the bed of the patient. Because of the reliance of battery operation during
normal measurements during the day, specific methods to increase the lifetime of the lithium battery
were implemented, such as the recharging of the battery to under the maximum power, such that the
number of recharge cycles was extended. Beyond that, the battery itself has the capacity to run theimplanted system for 48 h, but is recharged on a 24-h basis. A hybrid system such as this would be able
to eliminate the need for a continuous power signal and mandatory wearable external transmitter andprovide a more efficient strategy for long-term continuous pressure monitoring.
Implantable sensors for small animals or patients in an ambulatory setting can also have issues
with improper coupling between the implanted and external coils leading to poor power or datatransmission. Current efforts set to improve two aspects of transmission by either improving or
circumventing coupling issues between the two coils or also by reducing the attenuation-related effects
using novel coil designs. In the case of the former, there have been efforts in the development ofSensors 2014, 1420633

adaptive power control technology [25,30] that can enable the sensing mode if the power level is highenough to run the sensor and onboard ASIC system, or can store the transferred power to a capacitor if
the power received is not high enough. This technique circumvents poor coupling and uses smart
implant technologies to operate when the power becomes sufficient. In other efforts, one system utilizes
three external coils with a system that allows for the external coil that couples the best with the implanted
coil to switch on, while the others remain off [23]. This leads to short down time of the system while
the correct coil is being chosen, but can ensure proper coupling regardless of the position of theimplanted device. As mentioned previously, RF-battery hybrid systems eliminate the need forcontinuous RF coupling and can use a large external coils (within the bed of the patient, recharging of
the device occurs at night) that would reduce coupling factor variation.
To reduce attenuation, current devices set to improve the dielectric mismatch for implanted devices or
improve the quality factor of their transmitter coils. For example, certain LC tank-based systems faceissues once implanted due to changes in the dielectric constant of the surrounding wet environment
compared to air (εr ~ 1 in air and εr ~ 80 for saline) [55]. These changes in the dielectric constant of thesurrounding environment increases the total capacitance of the system and in essence decreases theresonant frequency as well as the quality factor of the system [55]. The silicone encapsulation
however, though it has been shown to causes changes to the sensitivity of the device, helps to decrease
these differences in dielectric constants.
Small ferrite rods can be used within in the coil to improve the quality factor of the coils for better
inductive coupling. Innovation in antenna technology includes the use of metal stents as more thanstructural elements, but also as built-in antennas (i.e., “stentenna”) [31,114]. A “pseudo” normal mode
antenna has been developed that has a unique circular polarization profile that decouples the foveation
related effects of poor power transfer based on device position [29].
Alternative forms of energy harvesting have been developed that allow the use of piezoelectrics(PZTs) to power some implantable pressure sensors [119]. An acoustic signal generated by an external
speaker was transmitted through the body to an implanted piezoelectric cantilever connected to apressure sensor within the bladder [45]. The low frequency of the acoustic signal allowed fortransduction through the body to the PZT with minimal signal loss, and no concerns of orientation oralignment related effects were observed as the acoustic wave field was large enough from the speakerto fully contain the PZT. This particular device decouples the dependence on alignment and separation
observed in RF-based telemetry devices and reduces the complexity of the implanted system byremoving the need for sophisticated onboard receiver circuitry. A similar technique is used for theRemon ImPressure device, where acoustic waves are used to both transmit power and data to theimplanted sensor. For IOP measurements, one technology implanted within the eye takes advantage ofa thin-film lithium ion battery and solar cell hybrid [120] that absorbs light entering the eye.
4.3. Drift Compensation Strategies
Significant advances in biocompatible fabrication and telemetry have been detailed and discussed
above, however drift remains as significant challenge. Sources of drift can be sorted to two categories:
those that affect baseline and those that affect sensitivity. The toolbox of techniques to combat drift isever expanding, as these methods are borrowed and shared amongst all types of transducers.

Sensors 2014, 14

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A common method to remove drift in the baseline measurement involves the comparison of sensor
signal to that of a controlled reference. For gauge membrane-based pressure sensors, the effect of theenvironment on the mechanical properties of the sensing membrane can be the source of baseline driftthat can compromise the long-term stability and performance of the implanted device [89,115].
The introduction of a capacitor that is unresponsive to changes in pressure and implanted alongside thesensing element can be used as a reference in a differential signal circuit that subtracts out common
effects of drift. Averaging the output from several sensors in an array is a variation of this concept
for drift compensation, with possible benefits of extended range and sensitivity. As discussed
previously, atmospheric pressure has been used as the baseline reference pressure for many sensing
modalities [72,77]. However, several efforts have been made to replace this gauge measurement with adifferential type that makes use of another sensor placed at a nearby low-pressure location (e.g.,difference in pressure inside vs. outside a bladder wall). Benefits of this type of measurement havebeen adopted with moderate success in improving baseline drift [14].
The addition of temperature data may assist the calibration of sensor response in the in vivo
environment. Efforts to improve the chronic performance of these sensors make use of an
on-board temperature compensation circuit [42,108,121] or external signal processing circuitry [89,118].
Appropriate material selection for certain sensing modalities such as SAW type transducers allow for
the use of temperature-insensitive quartz, which circumvents the need for compensation [121,122].
Drift in sensitivity for the commonly used capacitive pressure sensor is primarily caused by theaccumulation of biological material on the surface of the device, affecting the mechanical properties ofmembrane. Removal of the biofouling material is challenging without detrimental effects to thesurrounding tissue, and many strategies focus on mechanical approaches rather than chemical [123].
The addition of microstructures to mechanically disturb and break up the cellular matrix have been
proposed and developed for other in vivo applications may be applied to implantable pressuretransducers [124]. Ultrasound has also been used to clear biofilms [123], and may be possible to realize
on a SAW or piezoelectric membrane transducer by simply driving the membrane to move at ultrasonicfrequencies. Another strategy involves the use of aeration and bubbles to dislodge attached cellularmaterial [125], which can be implemented by adding electrodes on the surface of the implant toperform low-power electrolysis. Long-term water absorption can also play a significant role in affecting
mechanical response. Coatings to prevent biofouling and mitigate immune response would also serve toimprove drift in sensitivity by shielding the force-transducing membrane from the effects of the body.
The development of alternate transduction methods should seek to maintain sensitivity despite thepresence of biofouling.
5. Prospective Trends
Beyond the development of implantable pressure sensors within the aforementioned areas, trends forchronically implanted pressure sensors center on the use of novel biocompatible electronics
and alternative power sources. Briefly, there is considerable development of and vision forbiocompatible active electronics, that is organic transistors [86,126] and polymer-based MEMS
(i.e., Parylene C) [61,127]. These types of devices can remove the dependence on bulk silicon for theconstruction of micromachined devices to create implantable sensor and circuit systems that haveSensors 2014, 1420635

improved integration with the body once implanted. Taking this concept one step further, thedevelopment of biodegradable sensors and batteries [128,129] may enable implantable sensors that canhave a limited life-time (currently on the order of days), and then be completely absorbed by the bodyon demand to remove issues pertaining to a secondary surgery to remove the device following the endof a monitoring time frame or during the event of failure.
The development of alternative power sources outside of RF inductive powering is another area ofimprovement for implantable pressure sensors. Energy harvesting and power-MEMS is a large andgrowing area within the MEMS field, with a large portfolio of novel biocompatible devices for use in
implantable systems. Movement away from RF inductive powering towards energy harvesting
techniques in vivo can sever the tether to external units or coils to create a seamlessly integrated implantsystem. Along these lines, using these energy harvesting devices to power implanted body sensor
networks [83,130] and corresponding transmission of the sensed data to the cloud, can greatly improve
patient monitoring in the form of wireless health.
Lastly, the field is dominated by the use of membrane-based capacitive or piezoresistive sensing
systems for use in implantable pressure systems. Novel sensing mechanisms looking beyond thedependence on the mechanical deflection of a membrane can be insensitive to mechanical fatigue andmaterial stiffness changes, and potentially reduce issues related to drift.
6. Conclusions
From the onset of the development of implantable micromachined pressure sensors, the historicalchallenges of biocompatibility, efficient telemetry, and drift management have persisted. This review
highlighted current strategies to improve implantable pressure sensing technology, largely focusing on
biocompatible encapsulation along with drug coatings and the development of smart active circuitry to
improve telemetric coupling and eliminate drift. The large drift observed in many of the developed
sensors may overshadow the improvements gained in novel tissue integration strategies andpower/data transmission technology. Further work is needed to address drift management to improve
the long-term performance and gain widespread clinical acceptance of these monitoring tools.
Acknowledgments
This work was partially funded by the NSF under award numbers ECCS-1231994 and EFRI-1332394.
The authors would like to acknowledge members of the Biomedical Microsystems Laboratory at
the University of Southern California for their insightful thoughts and discussion in the preparation ofthis review.
Author Contributions
Lawrence Yu and Brian Kim contributed equally to the literature review. All authors contributed to
the writing of the final manuscript.
Conflicts of Interest
The authors declare no conflict of interest.

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