Implantable Devices: Issues and Challenges

Abstract

Ageing population and a multitude of neurological and cardiovascular illnesses that cannot be mitigated by medication alone have resulted in a significant growth in the number of patients that require implantable electronic devices. These range from sensors, gastric and cardiac pacemakers, cardioverter defibrillators, to deep brain, nerve, and bone stimulators. Long-term implants present specific engineering challenges, including low energy consumption and stable performance. Resorbable electronics may offer excellent short-term performance without the need for surgical removal. However, most electronic materials have poor bio- and cytocompatibility, resulting in immune reactions and infections. This paper reviews the current situation and highlights challenges for future advancements.

Full text

Electronics 2013, 2, 1-34; doi:10.3390/electronics2010001
electronics
ISSN 2079-9292
www.mdpi.com/journal/electronics
Review
Implantable Devices: Issues and Challenges
Kateryna Bazaka and Mohan V. Jacob *
Electronic Materials Research Lab, School of Engineering and Physical Sciences, James Cook
University, Townsville 4811, Australia; E-Mail: katia.bazaka@my.jcu.edu.au
* Author to whom correspondence should be addressed; E-Mail: mohan.jacob@jcu.edu.au;
Tel.: +61-7-4781-4379; Fax: +61-7-4781-5177.
Received: 9 October 2012; in revised form: 19 November 2012 / Accepted: 7 December 2012 /
Published: 21 December 2012
Abstract: Ageing population and a multitude of neurological and cardiovascular illnesses
that cannot be mitigated by medication alone have resulted in a significant growth in the
number of patients that require implantable electronic devices. These range from sensors,
gastric and cardiac pacemakers, cardioverter defibrillators, to deep brain, nerve, and bone
stimulators. Long-term implants present specific engineering challenges, including low
energy consumption and stable performance. Resorbable electronics may offer excellent
short-term performance without the need for surgical removal. However, most electronic
materials have poor bio- and cytocompatibility, resulting in immune reactions and
infections. This paper reviews the current situation and highlights challenges for
future advancements.
Keywords: implantable electronic device; bioresorbable electronics; radio-frequency (RF)
wireless powering; encapsulation
1. Introduction
Over the last 60 years, implantable electronic systems and devices have undergone a significant
transformation, becoming a valuable biomedical tool for monitoring, measuring and soliciting
physiological responses in vivo using wireless communication. The invention and subsequent
advancement of these devices have relied heavily on the growing knowledge regarding various aspects
of the human neuro-motor system, and the development of electronics technologies capable of
OPEN ACCESS

Electronics 2013, 2 2
interfacing with living tissues and organs at micro- and nano-scale. Increased in vivo stability,
miniaturization and lower energy requirement of modern electronics led to a multitude of miniature
wireless electronic devices, such as sensors, intelligent gastric and cardiac pacemakers, cochlear
implant, implantable cardioverter defibrillators, and deep brain, nerve, and bone stimulators being
implanted in patients worldwide [1–4]. Figure 1 shows several examples of electronic devices for
in vivo applications. According to Halperin et al. over 25 million US citizens were reliant on
implantable medical devices for life-critical functions [5], with the number of implantable cardioverter
defibrillator implants increasing tenfold between 1990 and 2002 [6]. Advances in semiconductor
technology, particular in the area of micro-electro-mechanical systems (MEMS) and microfluidic
lab-on-chip biomedical systems have allowed for the development of units for rapid diagnostics, and
precisely controlled pulsatile, rapid or sustained delivery of drugs and biomolecules and complex
therapeutics [7–11]. These systems have also been used for the development of tissue engineering
platforms and in regenerative medicine applications, particularly where muscular and nervous tissues are
concerned [12,13]. In addition to enhancing the survival rate and the quality of life of patients globally,
implantable electronic systems have contributed significantly to our appreciation of the biological
processes taking place within the human body, including the complex mechanisms of neural
communication and control, and greatly enhanced our understanding of how these are affected by
various diseases and treatments. Ex vivo, MEMS and dielectric elastomer actuators (DEAs) have been
used to investigate the manner in which biological cells modulate their behavior, express genes,
proliferate or differentiate in response to mechanical and electrical stimuli, knowledge which is
essential for adequate tissue engineering design [14–16]. In addition to playing a profound role in the
advancement of restorative medicine and biomedical sciences, implantable information and
communication technologies drive notable changes in the social and cultural attitudes of people
towards technology [17]. There, implantation is viewed beyond the medical context as a means to
enhance the abilities and experiences of healthy individuals.
In spite of substantial innovations in the fabrication and application of implantable biomedical
electronic systems since the first implantable heart pacemaker of 1958, the modern implants are still
faced with a number of challenges [18–20]. In terms of device production, there is a strong trend to
produce devices with ever diminishing size and weight in order to make them compatible with normal
human activities and enhance comfort for the host. Implants that weigh less than 2% of the patient’s
body weight are typically required [18]. When used, batteries, whether single-use or rechargeable,
significantly contribute to the overall weight and size of the device. Single-use, non-rechargeable
batteries, such as those used to support pulse generation in cardiac pacemakers and deep brain
stimulators, have a predetermined lifetime, at the end of which they have to be surgically replaced, at
high cost to the patient and the healthcare system. Rechargeable batteries, such as those used in
cochlear implants, can be powered or recharged transcutaneously using external signals, e.g., radio
frequency (RF), ultrasound, infrared light, low-frequency magnetic field, and so on. More recently,
internal charging using the energy produced by the physiological environment or natural body motion
has been investigated [21,22]. Further miniaturization can be attained by means of battery-less
implants, where energy harvested from natural or artificial power sources surrounding the patient is
used directly to power the device [23–26]. Inductive (or near field) and electromagnetic (or far field)
coupling are frequently used for remote powering of such battery-less devices [27]. In the former case,

Electronics 2013, 2 3
time-harmonic magnetic field generated by the low-frequency alternating current in the external coil
generates an alternating current in the implanted component [28], whereas in the latter, electromagnetic
waves propagate from the antenna in the far field region to power the implanted chip [29]. Biomedical
actuators that do not rely on the conventional wireless delivery, harvesting, accumulation and storage
of power in electrical form have also been investigated for such high-energy actuation applications as
drug release and mechanical adjustment in prosthetic devices [30]. Recently, Denisov and Yeatman
designed stepwise microactuators where incoming ultrasonic waves initiate vibrations in the
mechanical oscillator components of the device. The oscillations are then converted into stepwise
motion of a mechanical actuator through oblique impact, without the need to convert the energy from
mechanical to electrical form.
Figure 1. (A) Cross-section of a model of the modified hip implant with a metal head. The
temperature telemetry with thermistor, electronic circuit and power/data coil are placed
inside the neck of the implant [31]. (B) Retinal prosthesis: (a) the schematic version of a
minimally invasive approach; (b) photograph of an implant in the eye of a minipig [32,33].
(C) An active, flexible device for cardiac electrophysiological mapping: (a) circuit design;
(b) photographic image of the fabricated device in a slightly bent state; the inset shows a
magnified view of a pair of unit cells; (c) sequential images during the contraction cycle of
the heart, with blue lines emphasizing the degree of bending along the device and the black
arrow in the left-most image indicating a conventional pacing electrode [34]. (D) The
implantable microchip-based human parathyroid hormone drug delivery device
(54 mm × 31 mm × 11 mm, l × w × h) containing two microchips with 10 reservoirs each
(13.0 mm × 5.4 mm × 0.5 mm, l × w × h). Schematic cross section of microchip assembly
showing drug releasing from one reservoir [35].

Electronics 2013, 2 4
Concomitantly, there is a strong emphasis on increasing the functionality and reliability of these
electronic devices to support real-time complex in vivo stimulation, data collection, data compression
and fast wireless data transmission to external components of the system. This increasing complexity
of signal processing electronics further adds to the power budget of the device, which should remain
very low if the device is to remain operational for extended periods of time. For instance, a ultra wide
band technology offers high-speed data transfer between the implanted device, e.g., implantable
electronic cardiovascular devices, and the medical practitioner, and low interference potential, yet its
implementation is limited due to its high power consumption [36,37]. Although the wireless
programming is highly application-specific, differing significantly between that for a pacemaker and a
wireless pulse oxymeter, subjectivity of a particular wireless technology to interference is an important
factor to consider, given that the wireless devices operate within an electromagnetically shared
environment. Electromagnetic pulses, external electric fields and those from other indwelling electrical
devices can all generate interference. The ability of the implanted devices, such as pacemakers,
glucose-monitoring and insulin-delivery systems, neural stimulators, and smart prosthetics, to be easily
interrogated by health practitioners also makes these devices vulnerable to hacking [38,39]. In addition
to having access to sensitive patient data, the devices can potentially be reprogrammed, interfering
with the correct device operations. Therefore, security measures, including security check protocols,
firewalls, data encryption and restricted network access should be seriously considered.
Application of nano- and molecular-scale technologies for design and fabrication of the implantable
circuitry can lead to remarkable advancement in integration density and dynamic power dissipation,
enabling neuro-electronic interfacing and nano-bio-robotics [40]. However, current biomedical
nanotechnologies are still faced with challenges, such as lower reliability, relatively high stand-by
power consumption, and electron leakage due to insufficient insulation. Furthermore, in an effort to
improve the resolution of the biological signals being collected, the increasing number of electrodes
demands more power to be delivered to the electrode array, thus potentially increasing the thermal
energy dissipated within the implant circuitry [41,42]. Given the high cost and time associated with the
surgical implantation of the device and the recovery of the patient, long-term reliability of the device is
crucial. Peri-implant space is a chemically harsh environment, with the surface of the implant being
continuously attacked by the highly conductive and corrosive physiological medium which also carries
a variety of biochemically reactive organic molecules. The drive towards small, light and flexible
devices may undermine mechanical robustness of the implant; aggressive cleaning procedures used on
the devices prior to implantation may further contribute to weakening of the organic layers and
adhesives. The ensuing in vivo degradation and loss of integrity may be detrimental to the performance
of the device, potentially leading to the device failure, e.g., electrical shorting, and subsequent surgical
removal. The implanted device and its degradation by-products may stimulate activation of a range of
immune mechanisms, leading to inflammation, which in turn may further contribute to the implant
degradation. The toxicity of the leaching ions and fragments may hinder the recovery of damaged
tissues adjacent to the implanted device. Surface fouling and infections are also of great concern [43].
The abiotic surface of the implant presents a suitable ground for colonization by human pathogenic
bacteria. Once attached to the surface of the device, the bacterial cells may form a three dimensional
biofilm, which serves as a protection barrier against detachment, predation by host immune cells and
significantly reduces the efficacy of most systemic antibiotics [44,45].

Electronics 2013, 2 5
Achieving suitable biocompatibility is a complex matter, due to the dynamic multifaceted nature of
the host biological response to synthetic and organic materials used in device fabrication. Where
in vivo sensing or stimulation is required for a short period of time, resorbable implantable electronic
devices can provide a solution to overcome inflammation and infections associated with long-term
implant utilization. The premise is that the materials used in device fabrication are biodegradable, and
undergo controlled dissociation over time under normal in vivo physiological conditions. The
degradation by-products illicit minimal toxic response and are removed from the peri-implantation site
by means of normal metabolic activity [46]. However, fabricating a complex high-performing
electronic system from entirely biodegradable, non-toxic set of electronic materials is a difficult
undertaking, particularly at small scales. A combination of robust and reliable non-biodegradable
silicon electronics with bioresorbable polymer platform offers both the flexibility of the device and
sufficient bulk degradation that the immune response to the remaining material is minimal [47]. For
the technology to be clinically implemented, however, the challenges associated with integration of
sensitive electronics functions with the fabrication techniques used for production of biodegradable
component, and the control over degradation kinetics and biocompatibility of the device should be
addressed. In spite of many reports detailing the biological activity and degradation behavior of many
commonly used materials in vitro and in vivo, our appreciation of these complex processes is yet
to be adequate.
The aim of this review is to discuss the challenges faced by modern implantable electronic devices
and give a brief overview of the solutions that have been proposed, investigated and implemented in
order to overcome these challenges.
2. General Characteristics of Implant Systems
An implantable electrical system can perform a telemetry (sensing) function, whereby biological
data is collected, a teleactuation (stimulation) function, or both in a form of closed loop control.
Regardless of it intended function, an implantable system typically comprises of two fundamental
components, an indwelling module which resides within the host body, and an external device located
outside of the body. The external module is generally employed to transmit the information to and
from the internal module and deliver power to the indwelling component of the device. The indwelling
module may be fully electronic, or contain chemical, biological, or mechanical components. In sensing
electronic implants, the indwelling sensors detect, collect and translate the desired biological and
physiological parameters into electrical signals. These signals are then modulated by the interface
electronics and transferred by means of an inductive coupling link to the external receiver component,
where the data is recorded and analyzed. For example, a micro-accelerometer implanted directly on the
surface of the heart of patients who have just undergone coronary artery bypass graft (CABG) surgery
can be used to measure the heart wall motion as a means of early detection of surgery complications [9]. In
stimulation implantable systems, the external component is used to wirelessly transmit commands to
the indwelling component, where they are processed by the interface electronic circuitry to produce a
range of electrical stimuli. The produced electrical currents are then delivered to tissues and nervous
structure by means of electrodes. As an example, Enterra® therapy is used to treat diabetic gastroparesis by
applying electrical stimulation to the antrum via two indwelling unipolar intramuscular leads and a

Electronics 2013, 2 6
neurostimulator, where the stimulation parameters are adjusted noninvasively by using Medtronic
N’Vision clinician programmer [48]. A closed loop system encompasses both the indwelling sensing
and the stimulation components, and all information transfer and processing takes place within the
body of the patient. This type of an implantable electronics is commonly employed to maintain a
certain level of function within the human body, e.g., cardiac resynchronization, to enable automated
provision of medical care and prevention of critical incidents, e.g., sudden cardiac death [37].
Hemodynamic sensors have been integrated into implantable pacemakers to enable rate-responsive
pacing; these closed loop systems function on both sensed and paced ventricular beats, thus
surmounting the key constraint of the previous pacemaker systems, namely the need for permanent
ventricular pacing [49]. Concomitant sensing and stimulation provided by closed loop neuromodulation
devices provide a platform for enhancing therapies for neurological disease, while concurrently
assessing the instantaneous response of the neural system to stimuli [50].
Within the host body, the individual modules of the electrical system can reside intracavity, e.g., within
the intestinal, oral, or urinary systems etc., be implanted subcutaneously or deep within the tissues, or
be located on the external surfaces of the body [50–52]. Hard shell packaging is often used to protect
the electronic circuitry, whereas the remainder of the indwelling assembly may also have a soft
encapsulation layer [53]. The role of the protective casing and the encapsulant is two-fold. For one, the
hermetic protective casing ensures the in vivo integrity and reliability of electronic performance of the
devices over the life-time of the implant under the specific physiological conditions. This includes
protecting the device elements from the highly corrosive environment and ensuring no leakage current
flowing through the electrodes [53]. Secondly, the encapsulation layer performs a biocompatibility
function, protecting the host tissues from potentially harmful elements of the device. It can also
provide a soft low-friction conditioning layer, ensuring a smooth integration within host tissues.
Thirdly, the hard casing may offer mechanical support to devices that are submitted to a considerable
load or strain during extension/flexion and wear.
3. Typical Requirements of Implant Systems
When designing an implantable electronic system, several general requirements are to be addressed,
namely minimal size and weight, low power consumption, good reliability, high biocompatibility and
minimal toxicity, high data rate and data latency. As the case with any commercial product, the design
of the implantable devices is heavily influenced by the demands and preferences of their consumers. In
addition to being less invasive to the body of the patient during the implantation, smaller and lighter
devices are likely to result in less pain and discomfort to the host during healing and use. The
excessive size and weight may be detrimental to the healing process by putting pressure on the
adjacent tissues that have already been damaged as a result of surgery, contributing to the
inflammatory processes within the peri-implant space. Small and light devices are less restrictive in
terms of normal level of human activity, and thus afford better quality of life to the patients. The power
source and encapsulation components remain the major contributors to the overall weight and size of
the device, whereas the electric circuitry components have decreased dramatically with the
advancements in MEMS and nanotechnology. Coupling capacitors used to ensure charge-balance and
effectively minimize current leakage may further increase the volume of the implantable module [54,55].

Electronics 2013, 2 7
Lower power consumption is important in terms of both the long-term performance of the device
and the safety to the patient. Close proximity of the electrodes to living tissues places firm restrictions
on the amount of dissipation in power an implanted electronic system should not exceed, as extensive
dissipation may inflict damage onto these soft tissues [56]. In addition to thermally-induced damage [41],
the electrical stimulation-induced tissue injury (overstimulation) and damage due to the
electrochemical products released into physiological medium as a result of electrode corrosion should
be considered [54]. The energy use by interface electronics should also be minimized to ensure
longevity of the implants with single-use batteries, as the replacement of such a device would require a
costly and invasive surgical procedure [57]. Although using a rechargeable battery may address the
need for battery replacement surgical intervention, the need for frequent charging may be
inconvenient, time- and resource-consuming activity. For battery-less devices powered by an RF link,
the low power restriction is also applied to ensure the electromagnetic energy radiated or backscattered
by the device during wireless communication is in line with the IEEE human tissue exposure
standards [29]. Excessive electromagnetic fields can potentially undermine correct device functioning,
leading to temporary device malfunction or permanent damage. Indeed, device reliability is paramount,
as failure may not only cause discomfort, pain, or local damage to the peri-implant space, but may in
some cases result in the irreversible damage or death of the patient. Considering that many implants
are introduced deep into the tissues and cavities of the body, device maintenance is complicated, with
risks to the health of the patient. It is important to note that the presence of a neurostimulation system
may limit the electromagnetic diagnostics and treatment, e.g., magnetic resonance imaging (MRI), to
which the patient can be exposed [58]. Heating, magnetic field interactions, induced currents, and
interference with correct functioning of the implanted modules may result in considerable temporary or
permanent damage, e.g., transient dystonia, paralysis, coma, or death [59–61].
The electrode material and structure should be selected so that during stimulation, sufficient change
can be injected to elicit the desired response, and that the level of products from irreversible Faradaic
reactions that result from this stimulation are sufficiently low as not to damage the surrounding tissues
and the electrode itself [54,56]. Relatively low voltages of both spontaneous and evoked signals, as
well as those produces by the transducer necessitate particular care when designing methods for signal
detection, amplification, modulation and transfer. Spontaneous potentials, such as those detected using
electroencephalography, electrooculography, electromyography, or electrocardiography, occur naturally
within the body and are typically range from less than one μV to tens of mV range [62]. Potential
amplitudes of evoked responses, i.e. event related electrical potentials observed in the central nervous
system structures as a consequence of a stimulus, are even lower, falling in the less than one μV to
tens of μV region. The bandwidth of bioelectrical signals ranges from 0.01 Hz to 15 KHz, with low
frequency signal (less than 1 MHz) frequently used to wirelessly power up and transfer data from the
external module to the indwelling device [62]. Recently, however, implantable electrical systems that
function in the Medical Implants Communication Services (MICS) band (402–405 MHz) are being
developed, as this band has been expressly designated for implanted medical devices and is only
shared with meteorological aids [37].
Surgical placement, orientation, and extraction of the electrodes is intricate, particular where neural
system is concerned, and should be designed to synergistically interact with the available stimulation
parameter settings to attain the best remedial outcome for a patient [63]. Indeed, given the difficulty in

Electronics 2013, 2 8
revising the placement of indwelling electrodes, much care should be given to matching the electrode
configuration to the stimulation capacity of the stimuli generating module. In general, the specificity of
electrical stimulation is restricted due to electrode scaling and physical placement of these onto the
stimulated tissues, with some improvement obtained by manipulating the electrical current applied to
the tissues, and the volume of tissue being stimulated [64]. For example, current focusing and current
steering approaches in cochlear implant systems and deep brain stimulators employ current-controlled
stimulation using several autonomous sources of current to attain control over the volume of tissue
receiving stimuli [65,66]. Mechanical shaping and deep reactive ion etching were applied to the
implantable silicon-based probes used for neural stimulation to minimize the insertion force when
introducing multi-electrode arrays into the brain and spinal cord of the animals used in in vivo study [67].
The sensing and recording quality of these arrays were monitored over time, with neuronal spike
activity recorded up to 566 days after implantation. Such prolonged implantation has minimal impact
on the tissue architecture, as indicated by histopathology evaluation of neurons and astrocytes.
The capacity to simultaneously sense and stimulate is highly desirable, as it enables well-tailored,
prompt adaptive therapeutics and contributes to our understanding of natural and evoked neural
activity [68]. However, in practice, the ability of closed loop neuromodulation devices to detect brain
signals is limited, due relatively high amplitude of the stimulation potential compared to the field
potential signals used to sense brain activity [50]. Or, in the case of implantable cardiac defibrillators,
pacing at fast rates may delay or hinder detection of ventricular tachyarrhythmias [63]. Furthermore,
the processes resulting from the stimulation of neural networks are complex, involving both neural
excitation and inhibition. Experiments showed that at high frequencies, electrical stimulation resulted
in inhibition of subthalamic nucleus activity, while also directly exciting the cell and/or its axon [69].
Use of multiple sensors may raise the frequency of problems associated with hardware and software
integration, reduce long-term reliability and longevity of the device, and increase susceptibility to
oversensing of endogenous and exogenous signals, e.g., diaphragmatic myopotentials and
electromagnetic interference, respectively [63]. In themselves, complex algorithms may lead to
noncapture or oversensing of biological signals, potentially resulting in under or incorrect
diagnosis [63,70]. As a result, concurrent sensing and stimulation is often foregone if favor of detecting
and recording data regarding the immediate actuation performance, reducing neuromodulation
treatment to rigid stimulation system that relies heavily on the symptomatic assessment and actuation
tuning by the medical practitioner [55,68]. Although certain combinations of indwelling hardware,
e.g., high performance amplifiers, stimulation parameters and interpretation algorithms can minimize
residual stimulation disturbances, further research in this area is required [50].
The impedance disparity between the electrodes and the tissues contributes negatively to the ability
to detect neural signals, limiting the amount and usability of the information sensed. Microelectrode
impedance serves a key role in the monitoring of low amplitude and high-resolution extracellular
neural signals, and as such, changes in electrical interface impedance can be used as a preliminary
marker to infer long term electrode viability [71]. The impedance difference has been demonstrated to
increase with the length of implantation, whereby even those electrode designs that show adequate
performance under acute testing conditions may not necessarily show the same level and consistency
of signal capturing during chronic implantation [72]. For instance, an in vivo study involving
polyimide insulated tungsten microwire arrays implanted into the neural tissue of rats showed the first

Electronics 2013, 2 9
2–3 weeks post-implantation to be the most dynamic stage in the chronic electrode lifetime,
characterized by greater variations in the electrode impedance, functional electrode performance, and
the structural changes occurring at the electrode recording tips [71,73]. Longer term implantation was
associated with further electrode recording site deterioration, insulation damage and recession of the
recording surface. Similar results were observed in intracortical microelectrode arrays were implanted
into the pericruciate gyrus of cats, where the electrode-tissue interface changed daily over the first 1–2
weeks, then weekly for 1–2 months, stabilizing thereafter [74].
The mechanical tissue damage during the surgical insertion (acute trauma), as well as long term
contact of microelectrodes with electrically excitable tissues and micro movements associated with
electrode anchoring (chronic disturbance) induce activation of cells implicated in foreign body
response [75]. Mechanical mismatch between brain tissue and microelectrode material has also been
shown to affect the inflammatory response, with mechanically associated factors such as proteoglycans
and intermediate filaments shown to be important modulators of the response of the compliant
electrode material [76]. In the attempt to remove the foreign body, these cells release a host of
chemical and biological factors in the peri-implant space, some of which cytotoxic and neurotoxic
factors that contribute to localized neuronal degeneration and cell death [77]. Unable to enzymatically
degrade the implant material, the body responds by forming a thin layer of reactive glial tissue around
the implant to isolate the foreign matter from the surrounding tissues [78,79]. Such encapsulation is
detrimental to the ability of the electrode to sense signals, since it changes the diffusion properties of
nervous tissue (rendering it less permissive) and increases impedance [71,80], increases the distance
between the electrode and its nearest target neurons [74], and produces an inhibitory environment for
neurite extension, thus guiding regenerating neural processes away from the electrodes [72,81]. Gliosis
and enhanced formation of associated extracellular matrix molecules have been demonstrated to affect
molecule diffusion, and as such, neuron-glia communication, “cross-talk” between synapses,
extrasynaptic volume transmission, and tissue regeneration [80,82]. Even relatively small increases in
the separation between the sensing surface and the nerve tissue may be highly detrimental to the ability
of the former to detect a signal, since to adequately sense the neuronal spikes and local field potentials,
the distance between the neuronal ensembles and the target neurons should be within ~50 μm [77].
Local field potentials hold key information regarding functional behavior of neural networks that
correlates with disease symptoms, and can therefore be used as a biomarker [50].
Communication technologies used for data transfer to and from the indwelling device should
support high data rate, data latency, data accuracy and adequate data security, be reliable, and consume
minimal power [5]. The advancement of implantable devices used for sensing and stimulation resulted
in a considerable upsurge in the density of analysis and interpretation algorithms, consequently
contributing to the complexity and length of follow-up observations [37]. The extended battery life and
the increasing longevity of patients with indwelling medical devices further add to the ever increasing
number of implant carriers in follow-up. Given the limited amount of time and resources available to
medical practitioners, conventional follow-ups are followed by long periods of time when medical
personnel receive very little or no data on the wellbeing of the patient or the performance of the
indwelling module [83]. As a consequence, technologies that enable remote interrogation of indwelling
medical devices are attracting much attention [84]. Wireless remote monitoring facilitates collection of
technical information regarding the performance, attributes and settings of the implanted module, as

Electronics 2013, 2 10
well as the physiological parameters of the treated individual, and the outcomes that result from the
treatment [85]. The obvious benefits include the ability to promptly respond to the changes in the
clinical status of the patient, and minimize potentially harmful effects of implant malfunction or
failure; and ability to monitor the effectiveness of the treatment and alter the stimulation parameters
based on the data obtained. Furthermore, remote monitoring can effectively lessen the weight of in
clinic follow-up on the healthcare system, while maintaining or improving on the existing patient
safety standards [37]. The continuous stream of data can enhance the power of large-scale population
health bio statistical analysis, and thus contribute to the improvement in the quality of life of
the population.
4. Power Supply and Wireless Communication Technologies
The technologies used to supply power to the indwelling module can be broadly divided in single-use
non-rechargeable batteries and rechargeable batteries. The former can be commonly found in cardiac
pacemakers and deep brain stimulators, whereas the latter are frequently used to power cochlear
implants [86]. While single-use batteries require surgical removal to replace them, the rechargeable
batteries can be periodically recharged transcutaneously by means of wireless telemetry, which can
also be used to continuously powered up battery-less devices (without energy storage). Wireless
telemetry is also used to obtain the power status and performance of the non-rechargeable batteries.
Most commonly, the power is transmitted from the extracorporeal unit to the indwelling module via an
inductive coupling coil, which can be expressed as a lossy transformer. High wireless power transfer
efficiency is paramount to ensure minimal heating of the surrounding tissues, minimize the
interference with other devices and to reduce the size of the energy source [87,88]. Only a fraction of
alternating magnetic field generated by the coil within the external unit reaches the coil located within
the indwelling component, and is converted to alternating voltage [62]. The voltage is then rectified
and smoothed, and is fixed at a specific value suitable for the indwelling electronic circuitry. Aside
from loss associated with specificities of operating conditions (ambient environment), the power
transfer efficiency has been demonstrated to depend on the distance over which the magnetic field is
transmitted, i.e. the distance between the internal and external coils, the device geometry, and the
diameter of the coils [89]. Resonance-based wireless power delivery, where four high-quality (Q)
factor coils are employed instead of two, have been investigated for their improved energy transfer
efficiency and reduced dependence of the latter on the distance between the primary and secondary
coils [90]. The frequency chosen for the transmission is dependent on the type of the living tissues that
separate the indwelling module and the external component, specifically the frequency-dependent
attenuation by Foucault currents generated within the host tissues vary with the type of tissue [91].
Table 1 shows variations in electrical properties between biological tissues, measured ex vivo at 100 kHz.
Furthermore, addition of intermediate physical barriers, such as an encapsulation layer, has been
shown to further reduce the strength of the field, with encapsulant conductivity and thickness being
key determinants. Typically, using lower frequencies results in less loss compared to employing a
higher frequency field, however in real life most commercially available implantable devices use
higher frequencies to increase the data transfer rate [86]. The choice of frequencies is also affected by
the legislative regulations that specify the radiated power maximum to each frequency band.

Electronics 2013, 2 11
Table 1. Dielectric properties of tissues 1.
Tissue Type Relative Permittivity εr (×103) Conductivity σ (S/m)
Bone 0.28 0.0144
Liver 9.8–14 0.15–0.16
Spleen 3.3 0.62
Blood 2.7–4.0 0.55–0.68
Kidney 10.9–12.5 0.24–0.25
Retina 4.75 0.52
Bone (cancellous) 0.47 0.09
Bone (cortical) 0.23 0.02
Bone (marrow) 0.11 0.003
Cartilage 2.57 0.18
Skeletal muscle 14.4–27.3 0.38–0.65
Fat 0.09 0.02
Cerebrospinal fluid 0.1 2
Brain (grey matter) 3.8 0.17
Brain (white matter) 1.9–3.4 0.12–0.15
1 Measured ex vivo at 100 kHz, adapted from [92–94].
Less conventional energy harvesting methods that involve internal charging using the energy
produced by the physiological environment or natural body motion have also been reported, with
several examples presented in Figure 2 [21,22]. Sontag et al. suggested using highly dense
electroactive conjugated polymer brushes of poly(thiophene) and poly(phenylene) fabricated by means
of a surface-initiated Kumada-type polycondensation reaction to power up implantable devices [95].
Mercier et al. demonstrated energy extraction from the biologic battery in the inner ear, whereby the
electrochemical gradient within the ear is utilized as a power source for an anatomically sized,
ultra-low quiescent-power energy harvester chip integrated with a wireless sensor capable of
monitoring the ear electrochemical gradient [96]. When implanted in a guinea pig, the chip was able to
extract a minimum of 1.12 nW for up to 5 h, enabling a 2.4 GHz radio to transmit measurement of the
electrochemical potential every 40–360 s. Rapoport et al. reported the development of an implantable
fuel cell that generates power through glucose oxidation, producing 3.4 μW cm−2 and up to 180 μW cm−2
steady-state power and peak power, respectively [21]. Glucose is oxidized at the nanostructured surface
of an activated platinum anode, and oxygen is reduced to water at the surface of a self-assembled
network of single-walled carbon nanotubes embedded in film that forms the cathode [97,98]. The
half-opened geometry allowed the researchers to meet the requirement for simultaneous and
independent oxidation and reduction and thus avert electrochemical short circuits. The computational
investigations found that theoretically, glucose can be harvested from the cerebrospinal fluid to an
energy level of >1 mW without negative physiological consequences, thus confirming the potential of
this energy source to power brain-machine interfaces with low energy consumption. Glucose biofuel
cells with glucose oxidase and laccase mechanically incorporated into a conductive pure carbon
nanotube matrix were demonstrated to deliver a higher power density up to 1.3 mW cm−2 and an open
circuit voltage of 0.95 V [99]. Under physiological conditions of 5 × 10−3 mol−1 glucose and pH 7, the
devices remained stable for one month, delivering 1 mW cm−2 power density. Connected in series, two

Electronics 2013, 2 12
of these cells were able to deliver an open circuit voltage of 1.8 V with a maximum power of 3.25 mW
at 1.2 V, indicating the possibility of using these cells to power implanted biomedical devices that
typically require at least an operating voltage of 0.5–0.6 V. For example, a cytochrome P450—based
molecular biosensor used for drug sensing with temperature and pH monitoring was reported to have a
power consumption of 48 μW, with 32 μW required for the molecular detection, 2.5 μW for the pH
measurement, 1.4 μW for the control over the temperature sensor, and 12 μW for the multiplexing and
measurement reading [100]. Although in vitro studies return promising results, in vivo performance of
enzymatic biofuel cells is considerably lower. For instance, in vitro (at 4.7 × 10−3 mol−1 glucose, pH 7.2),
an intravenous implantable glucose/dioxygen hybrid enzyme-Pt micro-biofuel cell showed high
electrocatalytic performance with an open circuit voltage of 0.4 V and a maximum output power of
0.2 mW cm−2 at 0.25 V [101]. Once implanted into the jugular vein of a living rat, the device was able
to deliver an open circuit voltage of 125 mV at a maximum power density of 100 µW cm−2 at 80 mV.
Furthermore, the lifetime of the enzyme, and thus the long-term performance of the device remains an
issue, with a notable loss in the power generated with time in vivo [102].
The telemetric link can be used for bi-directional transfer of information, including the sensed and
recorded data about the patient and data regarding the condition of the indwelling module; the link also
enables wireless re-programming and communication between the multiple implanted modules
comprising a wireless network within the body of the patient. Typically, to enable powering/data
transfer using the same link, magnetic field modulations are employed to impress the data signal onto
the carrier signal used to power the device. The changes in the signal characteristics are detected and
interpreted by the indwelling module. The amplitude, phase and frequency of the signal can be
modulated. The amplitude modulation is one of the most popular techniques for short range
communications, and involves varying the amplitude of the signal from high to low, thus emulating the
zero/one logic of digital communication; it is described by the modulation depth, i.e., the extent to
which the amplitude was altered. In addition to the aforementioned amplitude modulation (AM) and
amplitude shift keying (ASK), frequency modulation (FM) and frequency shift keying (FSK) refers to
altering the frequency of the carrier signal, and phased shift key (PSK) involves changing the phase of
the carrier signal by 180° or less. Typically, the data rate attainable with the above modulations is
approximately 10% of the carrier frequency, however higher data rates can be achieved with more
sophisticated modulation approaches, e.g., by combining two modulation techniques. The choice of
modulation methodology used will also depend on the data transfer requirements of the implant, with
lower frequencies used for those with low data rate needs and higher frequencies for those demanding
large volume, ongoing data transmission. The restrictions within the system, e.g., availability of power
or bandwidth, will also affect the choice of the modulation approach. Appropriately chosen,
modulations can enhance the quality of the signal, improve the security of the patient-related data,
increase the quality of the signal, enable accurate transfer of data in the presence of noise and other
disturbances, and increase communication channel capacity. Communication channels can be
organized into additive white Gaussian noise channels (AWGN), band limited channels and fading
channels [62]. The former represents a channel model where white noise of a constant spectral density
and a Gaussian distribution of amplitude is added to the signal sent through the channel. In the band
limited channel model, the band width of the channel is smaller than that of the signal, resulting in the
elimination of the frequency components of the transmitted signal above the channel cutoff frequency.

Electronics 2013, 2 13
In the case of the fading channel, the amplitude and phase of the passing signal change rapidly,
attributed to fading due to multipath propagation and shadowing.
Figure 2. (A) Power extraction from cerebrospinal fluid by an implantable glucose fuel
cell: plausible site of implantation within the subarachnoid space and a micrograph of one
prototype, showing the metal layers of the anode (central electrode) and cathode contact
(outer ring) patterned on a silicon wafer [21]. (B) A photovoltaic-driven energy-autonomous
CMOS implantable sensor [103]. (C) An anatomically sized chip that harvests the energy
of the electrochemical potential in the guinea pig cochlea to power a wireless transmitter:
(a) plausible site of implantation within the mammalian ear; (b) cross-section of a typical
cochlear half-turn, showing the endolymphatic space (yellow) bordered by tight junctions
(red), the stria vascularis (green) and hair cells (blue), which are contacted by primary
auditory neurons (orange) [96].
An appropriate demodulation technique is also selected to minimize power consumption, reduce
interference, and ensure accurate translation of the message. The information transferred from the
indwelling module to the extracorporeal device is also modulated, with the electrical impedance of the
implanted electronic circuit being reflected back to the transmitter circuit via the same inductive
coupling link. The load shift key modulation (LSK) is attained by electronically switching the
impedance of the implant between two states. As with other modulation approaches, the data rate is
dependent on the carrier frequency. It is important to note that as the distance between the implanted
and external inductive coils increase, the magnetic field induced by the external coil progressively
transforms into an electric field, which cannot be modulated using LSK [62,86]. Thus, LSK may not

Electronics 2013, 2 14
be appropriate for deep tissue implants at certain frequencies. This is not the case for those implants
powered by the non-rechargeable batteries, where good transfer rate can be achieved at low
transmission power. As the advancement of wireless sensor network continue to develop, new
modulation techniques will need to be designed to address the needs of these complex systems. For
example, an intra-body area network (implant BAN) are being considered to establish timely, reliable,
and secure communication between indwelling devices, e.g., a cardiac implant, nerve sensor, and a
drug delivery pumps; or a series of diverse injectable microdevices used for multi-site stimulation
and sensing [104,105].
5. Remote Monitoring Technologies
Electronics systems used for diagnostics, e.g., endoscopic capsule, remain within the body of the
patient only a short time, and the patient is typically monitored by the physician at the clinic for the
duration of the procedure. Implantable electronics systems that are intended to reside within the body
of the patient for years, e.g., implantable cardiovascular devices, are reviewed intermittently, with
follow-up visits followed by extended periods of time when the medical practitioner receives no
information regarding the performance of the implantable system or the well-being of the patient. The
operating parameters of the indwelling device also remain static between the follow-up visits, which
may not reflect the needs and the clinical state of the patient. Then again, many scheduled follow-up
visits do not result in any changes being made to the device parameters and the patient requires no
medical intervention. A retrospective analysis of 1739 clinical visits by a random set of 169 patients
with implantable cardiovascular devices found that out of 1530 scheduled visits, 1197 visits resulted in
no relevant medical or device-related findings [106]. The non-scheduled visits, on the other hand, were
significantly more likely to result in identification of device- and/or patient-related problems and
require medical treatment, device re-programming, and hospitalization.
Remote monitoring can provide a robust system capable of timely capturing the device- or
patient-related issues, and ensuring that healthcare time and resources are spent where they are most
required [5]. Indeed, the same retrospective study found that a remote monitoring system was capable
of correctly detecting the vast majority of arrhythmias and/or device-related problems, potentially
missing an isolated pacing problem in less than 0.5% of all patients investigated [106]. Similarly, a
study involving the comparison between clinical traditional observation and remote measurements
found no statistical difference between the two conditions [107]. Remote monitoring involves a
periodic transfer of data, e.g., device parameters and functions, biological signals and clinical status of
the patient, from the indwelling module to a transmitter which is typically located outside of but in a
close proximity to the body [108]. There are a number of biological parameters that are monitored
depending on the patient, their medical condition and the type of indwelling device. For instance, in
those suffering from heart failure, common parameters to measure include: transthoracic impedance to
detect changes in fluid balance; electrocardiogram to identify the onset of atrial or ventricular
arrhythmias; blood pressure to manage hyper- or hypotension using adequate administration of
medicines, e.g., angiotensin-converting enzyme inhibitors and beta-blockers; temperature as an
indicator of potential infection; and blood oxygen saturation levels [20,109]. Upon receipt of the data
by the transmitter, the information is encrypted and securely sent to a central server of the

Electronics 2013, 2 15
manufacturer of the implantable system trans-telephonically or via web-based networks [37]. For
example, Home Monitoring technology introduced by Biotronik (Biotronik GmbH, Berlin, Germany)
in 2001 uses a device similar to a mobile phone to automatically transmit encrypted information from
implanted electronic cardiovascular devices, e.g., pacemakers, implantable cardioverter-defibrillators,
and heart failure devices, to a central server using mobile phone network. Other systems that use
standard and mobile telephonic communication channels include CareLink developed by Medtronic
(Medtronic Inc., Minneapolis, MN, USA), Housecall Plus used by St. Jude Medical (St. Jude Medical,
Sylmar, CA, USA), and Latitude by Boston Scientific/Guidant (Boston Scientific, St. Paul, MN,
USA) [110]. Currently available remote monitoring systems are manufacturer-specific, that is they can
only be employed to interrogate devices fabricated by the same manufacturer [108].
From there, the processed data can be accessed by relevant medical practitioners, and incorporated
into the hospital information system. The processing centre can also send the data on to the clinical
team responsible for the device using email, fax, SMS, etc. In addition to scheduled transmissions,
e.g., daily or weekly data transfers, a failure in the performance of the indwelling device or worsening
of the patient’s condition prompts an emergency data transmission to the server and subsequent
notification of the medical practitioner associated with the device. The early detection prevents or
minimizes the negative consequences of the event and increases the patient’s chances for survival and
recovery [111]. Furthermore, by analyzing data preceding an emergency event, medical practitioners
can identify the patterns and thus predict and potentially mitigate events leading to hospitalization. For
example, 123 patients implanted with cardiac resynchronization therapy devices with embedded Home
Monitoring capability were monitored over 12 months, at the end of which the data collected using
remote monitoring system was retrospectively analyzed against re-hospitalization and other clinical
events [109]. The transmitted data embraced several potential predictors of death or hospitalization,
including the onset of atrial and ventricular arrhythmias, extent of physical activity, mean heart rates
over 24 h and at rest, extent of cardiac resynchronization therapy delivered to the patient, and device
lead impedances. The study found that in 70% of the re-hospitalization cases, there was an increase in
mean heart rate at rest and in mean heart rate over 24 h within 7 days preceding the event. In 30% of
re-hospitalized patients, there was a notable decrease in the duration of daily physical activity, and in
43% of re-hospitalization incidents they were preceded by a reduction in the percentage of
resynchronization therapy delivered. Early detection of these patterns and timely response is likely to
result in a significant reduction in a number of hospital re-admissions, duration of hospital stay, and
patient mortality.
The benefits of employing the wireless body area network to advance healthcare quality are
undoubted, from ongoing health surveillance and patient- and progress-tailored rehabilitation to
emergency response systems and large-scale longitudinal medical and spatio-temporal social
studies [112,113]. Wearable sensors have been used to monitor motor fluctuations in patients suffering
from Parkinson’s disease, including estimating the severity of tremor, bradykinesia and dyskinesia
from accelerometer data features [114,115]. Compared to clinical visual observations, the sensor
network was able to accurately quantify the severity of the tremors with 87% accuracy, separating
resting and postural tremors, and discriminating tremors from other Parkinsonian motor symptoms
during daily activities [115]. In the case of major emergency events, e.g., natural mass-casualty
disaster, computer-enabled monitoring of clinical statuses of patients can facilitate prioritization of

Electronics 2013, 2 16
medical help to those who need it most, which can potentially save many lives [116,117]. In general,
the requirements for the sensor network will be influenced by the spatial and temporal scopes of the
study, the number of individuals/sensors for which the network is required, and the nature of the
wireless networking and sensing technologies that are being employed [113,117]. For instance,
availability of power and ergonomics of the system become more important as the temporal scope of
the study increases, whereas increasing the spatial scope such as in the case of epidemics study will
impact on the choice of communications infrastructure. For instance, miTag is a cost-effective scalable
wireless sensor platform developed to automatically track patients throughout the disaster response
process, from the scene through to ambulance and clinic [118]. It employs a 250 kbps 2.4 GHz IEEE
802.15.4 radio protocol, with 15 Bytes per second maximum data rate per miTag and an augmented
wireless range of 200 m indoor/400 m outdoor, which is similar to other specialized wireless sensing
platforms (known as motes) [117]. This platform can sustain a range of commodity sensors, including
GPS, pulse oximetry, blood pressure, temperature, electrocardiogram, to name a few, with the patient
data being relayed over a self-organizing wireless mesh network. In the pilot trials, the system was
shown to adequately and significantly increase the patient care capacity, reliably transmitting
patient-related data within radio-interference-rich critical care settings [118]. MEDiSN wireless sensor
network which uses miTag was also shown to tolerate high degrees of human mobility [117].
An exciting prospect, a sensor network of this size and complexity requires good understanding of
the network dynamics, including the capacity of a routing protocol to respond to node malfunction and
breakdown [111,119]. While employment of simulators and testbeds facilitate the advancement,
debugging, and spatio-temporal analyses of the sensor networks, e.g., determining the power
consumption, these tools are unable to fully account for the complexity of radio channel
characteristics, environmental stimuli, node mobility, and hardware failures of a real life
network [119–122]. Several passive external tools capable of observing, recording and reconstructing
of critical aspects of complex sensor networks in situ have been proposed, including LiveNet
developed by scientists at Harvard University [119,123]. LiveNet uses passive monitoring of radio
packets recorded by packet analyzers co-deployed with the network. Traces intercepted by multiple
packet analyzers (sniffers) are agglomerated to construct a global behavioral picture of the network to
which a range of analyses can be applied to determine application behavior, data transfer rates,
network topology, routing protocol dynamics, and packet loss [119]. When applied to a 184-node
sensor network testbed used to monitor vital signs of patients during a emergency drill, LiveNet was
able to correctly reconstruct the topology of the network, established bandwidth usage and routing
paths, discover hot-spot nodes and sources of packet loss [119]. Compared to traditional network
monitoring systems, LiveNet has a number of advantages, including (i) decoupling of packet
interception from trace analysis, allowing for the as-captured packet trace analysis; (ii) performance
and reliability of the network being investigated is preserved due to no changes to the network
required; (iii) the infrastructure for the non-intrusive monitoring is installed, reconfigured, and
dismantled independently from the network being monitored, and thus can be used on a need-to basis;
(iv) the monitoring system can be employed to monitor mobile or physically inaccessible sensor nodes.
Nonetheless, LiveNet and similar monitoring systems are faced with challenges, including
insufficient coverage of sniffer infrastructure with respect to total number of packets intercepted; the
trace merging being undermined by partial packet traces and insufficient sniffer time synchronization;

Electronics 2013, 2 17
intricate extraction of comprehensive data from the detailed traces [119]. Furthermore, as with any
other sensitive data storage and transmission, there are significant concerns regarding the security of
the wireless body area network [124]. The security and privacy of the data are in direct competition
with the practicality and usability requirements, particularly in light of the finite space and energy
resources available within the implanted module [116]. Data tampering, e.g., the deliberate destruction
or manipulation of data, can result in incorrect diagnosis and ineffective and/or wrong treatment of the
patient [124]. The intruder can attain patient-related data by intercepting the radio transmissions
between the sensor and the receiver, even when the data being transferred is encrypted [117]. This is
achieved by detecting the unique set of RF waveform features pertaining to a transmitter, so called
fingerprint, and the timing of each transmission [125]. These can then be used to associate each
message with a unique transmitter, thus providing an indication of the location and type of each sensor
that is communicating with this transmitter. To counteract this fingerprint and timing based snooping
attack, several approaches have been proposed, e.g., signal attenuation outside of the patient’s home to
increase the packet loss ration of the invader; periodically transmitting signals even if they do not
contain patient-related data; arbitrarily delaying messages to obscure the time at which the patient-related
data is being transmitted; making the transmitted fingerprint less discoverable; sending false messages
that imitate real events [117]. The open and dynamic nature of the wireless body area network and
distributed patient-related data storage often results in data being accidently lost, and therefore not
readily available for retrieval by clinical staff [116,123]. It has been suggested that the reliability of the
wireless sensor network can be improved by employing compressed sensing theory [126]. Compressed
sensing is believed to have the capacity to enhance processing capability, storage capacity, and time of
testing of wireless sensing networks [127]. It is based on the premise that sparse signals as those
observed in wireless sensor networks can be accurately reconstructed from a small fraction of random
linear measurements.
While remote monitoring technologies have been progressing steadily, remotely controlled
treatment delivery and reprogramming of indwelling electronic modules is hindered by the regulatory
issues, with law restrictions currently prohibiting remote reprogramming outside of the clinic [5,124].
Bodmer and Capkun suggested a number of potential security and privacy risks associated with the
remote reprogramming function embedded into cochlear implants [128]. A typical cochlear device
consists of a microphone worn behind the ear, an external speech processor, an indwelling signal
receiver/stimulator, and a remote control unit which allows the user to change some of the settings of
the implant [129]. Given the individual needs of each patient, the device is fine-tuned upon
implantation to ensure the appropriate performance [130]. Clearly necessary, the remote
reprogramming function also makes these devices potentially vulnerable to malicious interventions
that can range from turning off the implant to render the user deaf, to reprogramming the intensity of
cochlea stimulation where patient can hear sounds in the absence of corresponding external sound or
suffer from a painful perception of acoustic signals [128]. A less obvious intervention, the sound
processing unit can be reprogrammed to disregard the input from the microphone, replacing the
external sound with the sound generated by the intruder. Such message replacement invasion may be
particularly successful where the cochlear implant patient cannot receive visual confirmation of the
message. Malicious modifications of implanted cardiac devices may be even more devastating to the
health status of the patient, with potentially lethal outcomes [131]. This can be accomplished by giving

Electronics 2013, 2 18
an implanted cardiac device a command to stop operating when a patient has a cardiac episode or to
induce an episode by triggering defibrillation [132]. A battery-powered implantable cardioverter
defibrillator can also be saturated with external communication requests from the unauthorized device,
thus posing a risk of denial-of-service attack and potentially dangerously depleting the battery power
level of the device. Information extraction and invasion of privacy have also been identified as a
potential security issue [129]. Unauthorized scanning of people to determine the presence and type of
implantable medical device, such as a cochlear implant or implantable cardiac device, is a potential
threat, as such a discovery may result in people being discriminated against, with negative economic
and social consequences [132]. A tempered cochlear implant system can perform as a sound recording
and transmitting device, thus infringing the privacy of the user and those surrounding the bearer of the
implant. The implantable system can also be reconfigured to act as a tracking device.
A sophisticated authentication system that integrates all the implant components may aid in
ensuring the integrity of the communication and limiting the unauthorized access to the device. In
doing so, however, it may render the device less accessible by medical practitioners at clinics outside
of the patient’s clinical team [131]. For instance, in a time-sensitive situation when a patient loses
consciousness and is treated by an emergency unit in a country other than their own, an open access
device may allow rapid medical response and potentially save the life of the patient. To handle the
dynamic nature of the emergency response, an elliptic curve cryptography-based public key encryption
scheme can be employed for authentication, such as in the CodeBlue project [133]. Yet, this
authentication does not ensure the security of the data stored within the network, or the control to these
data. Another consideration is that the advanced security system is likely to increase the power budget
of the device and require some changes to the electronics design of the indwelling module. Halperin
et al. reported on a system of zero-power defense and prevention mechanisms that reside at the
interface between an implantable cardioverter defibrillator and the external components [134]. These
mechanisms are powered by externally delivered RF energy and not the primary battery of the
indwelling module, and can effectively mitigate the attacks from adversaries using custom and/or
commercial programmers. Upon an encounter of a security-sensitive event, the notification mechanism
sounds a warning to the implant bearer whilst symmetric encryption and authentication prevent
unauthorized access. The limit memory resources available to software developers necessitate the use
of lean, event driven concurrency models, significantly different to the conventional operating system
designs [117]. The limited computational power and restricted bandwidth result in the sensor nodes
engaging in a limited on-board processing to minimize information transmission requirements.
6. Biocompatibility and Implant Associated Infections
The reliability of implantable electronics has undergone significant improvement over the last 50
years, mostly due to the advances in encapsulation and packaging designed to protect the indwelling
module against the factors of the hostile environment [53]. Nonetheless, the issues with
biocompatibility and the propensity of the implanted constructs to get infected over time remain. The
highly invasive nature of many surgical procedures coupled with numerous serious health conditions
and inadequate immune response frequently observed in the implant recipients make these individuals
highly vulnerable to implant-associated infections [135]. Depending on the degree of severity, the

Electronics 2013, 2 19
complications may range from those that are painful and requiring localized antibiotic therapy to those
which necessitate complete removal of the infected device and systemic antimicrobial therapy [63]. If
left untreated, septicemia may develop, with potentially lethal consequences. The incidence of infections
associated with cardiac pacemaker implant ranges from 1%–19%, with 7%–8% attributed to
contamination during laboratory handling or the event of implantation [63]. Typically,
biomaterial-associated infections can develop along several pathways, with peri- and post-operative
contaminations being the most common route of introduction of the etiological agents [136,137].
Patient-specific factors including diabetes mellitus and long-term anti-inflammatory medication of the
patient using corticosteroids and other immunosuppressive drugs may slow down surgical site healing
and patient recovery, making the host more susceptible to developing an infection [138]. In addition,
the pathogens can originate elsewhere in the body, spreading to the implant site via blood to initiate
late hematogenous infection. These can include implanted central venous catheter used for
hemodialysis or other long-term access, a distant focus of primary infection, e.g., pneumonia, skin and
soft tissue infections, and invasive procedures unrelated to the implanted device [139,140]. This route
of infection is particularly relevant to the implants that are exposed to the blood stream [141,142].
The relatively high rate of implant-associated infections can be partially attributed to the fact that
many of the physico-chemical attributed of the implanted surfaces render them a highly suited
colonization ground for the bacteria. The non-living nature of the implant surface means that it does
not respond to being colonized nor does it produce chemical signals to notify the surrounding tissues
of the imminent danger. Certain combinations of surface properties can be employed to mitigate the
initial stages of bacterial attachment; however they are powerless against bacterial cells that manage to
adhere to the implant surface. Furthermore, bacterial cells have been demonstrated to release a vast
array of extracellular polymeric substances to pre-condition a surface otherwise not suited for
habitation or to form a three-dimensional polymer networks called biofilms. In their biofilm state,
bacterial cells are protected against predation by the host immune cells and the effect of systemic
drugs. Surfaces capable of both preventing the bacterial adhesion and replication, and eliminating
attached bacteria by releasing antibacterial drugs are receiving significant attention. The drug eluting
property can be imparted onto the implanted module via encapsulation or surface modification. In
addition to traditional antibiotics, a range of alternative antimicrobial agents have been considered,
including silver ions, nitric oxide, bioactive antibodies, and other bactericidal compounds [143,144].
In a study by Rohacek et al., the most common symptoms of infected antiarrhythmic devices were
pocket erythema and local pain, with 68% of the pathogens being coagulase-negative staphylococci,
followed by Staphylococcus aureus (23%), and 13% multipathogen infections [145–147]. Frequently,
the infections associated with implantable cardiac devices remain undetected for extended periods of
time, or even for the duration of the implantation [147]. Early detection and timely removal of the
implantable system (including potentially contaminated external modules) significantly increases the
chance of the patient to recover [148]. However, the re-implantation may not be as straight forward as
the infection will need to be under control prior to implantation [149]. Furthermore, a different site is
chosen for re-implantation since the status of the previous tissue may not be sufficient for the
successful healing. Persistent infections, particularly from non-retrieval of infected elements from the
patient’s body, have a significant rate of morbidity of over 60% [150,151]. The expense associated
with medical and surgical treatment of an infection around the implantable cardiac electronic device

Electronics 2013, 2 20
has been estimated to range from $25,000 for permanent pacemakers to $50,000 for implantable
cardioverter-defibrillators [138].
In addition to improving the hygiene during operative and post-operative procedures and
administering prophylactic antibiotic treatments, improved clinical outcomes can be attained by using
an AIGISRx antibacterial envelope (TYRX Pharma, Inc., Monmouth Junction, NJ, USA) consisting of
polypropylene mesh loaded with minocycline and rifampin [152]. Once implanted with a cardiac
implantable electronic device, the envelope is capable of progressively releasing these agents into the
generator pocket. Other products are employed to mitigate surgical site infections, e.g., arglaes wound
dressing (Medline Industries, Inc., Mundelein, IL, USA) which uses silver antimicrobial technology to
prevent bacterial infections through a continuous release of silver ions into the wound space. Silverlon
CA (Argentum Medical, Chicago, IL, USA), Aquacel Ag (Conva Tec USA, Skillman, NJ, USA), and
Silvercel (Systagenix, Quincy, MA, USA) also use silver antibacterial technology.
Inflammation is another important factor that can significantly undermine the utility of the
implanted module in vivo. When the properties of the abiotic material are not properly matched to the
characteristics of the surrounding tissues and cells, the integration and long-term utilization of such a
material may not be sufficiently successful. As mentioned earlier, the surfaces of the implants can be
contaminated with the bacteria or fragments of bacterial cells, which can induce an inflammatory
response. The physical presence of the implant, e.g., pressure it exudes onto the surrounding tissues
and organs, can also trigger inflammatory mechanisms where affected cells release a host of chemical
communication messengers. The surgical intervention to introduce the implant in the first place may
initiate the pro-inflammatory response, thus delaying tissue healing. Being exposed to a chemically
harsh environment may lead to implant degradation, with a cellular-mediated inflammatory response
which potentially resulting in a contained tissue loss in the immediate proximity to the implant [153].
In many cases, a correction surgery and extended post-operative care may ensue. Furthermore, the
loosening of the implant has been linked to the increased incidence of inflammations, complications,
and less successful functional performance of the implant. Debris-induced inflammation is a
multifaceted process, thus deciphering the etiology and pathology of the implant-induced inflammation
reactions is challenging and involves detection and interpretation of cellular events triggered by the
leaching and transfer of degradation particles [154,155]. The interconnected, often snowballing nature
of these cellular events and the multi-component nature of most implantable devices further complicates
matters. In addition, the installation techniques and level of activity of the patient influence the life-span
of the indwelling device and amount of degradation particles released. Finally, there are a plethora of
host-specific factors that play significant role in the implant tolerance [156], including genetic
predisposition, biomaterial hypersensitivity, chronic diseases, diet to name a few [157–159].
Given the multiple materials used for the fabrication of most implantable devices that include
polymers, ceramics, metals and composites, respective contributions of different types of degradation
fragments should be considered. Both polymer and metallic debris were linked to activation of
macrophages and giant cells in the peri-implant area, contributing to tissue loss and third-body
accelerated degradation of the implant [160]. In terms of their size distribution, the metallic particles
are typically smaller, more uniformly sized and more abundant compared to the polymer fragments.
Given their size differences, the metallic particles are more mobile and more easily transferred from
the peri-implant space to other tissues and organs, where they can activate host immune cells and

Electronics 2013, 2 21
trigger of an implant-associated inflammatory reaction in the host. Large, irregularly shaped
ultra-high-molecular-weight polymer particles are less mobile, with a tendency to accumulate in
tissues close to the implant site. The reactivity of different types of particles is also affected by their
size to volume ratio, with metallic debris undergoing faster corrosion as a result of exposure to
biological fluids. A comparative in vitro investigation of the inflammatory response of macrophage
cell to metallic and ceramic titanium based particles demonstrated the relative inertness of the
ceramic TiO2 versus Ti particles [161]. While notably less reactive than metallic degradation
fragments, ceramic particles, e.g., alumina, have been demonstrated to instigate an end-stage inflammatory
response [162–164].
As the case with coatings used to control bacterial adhesion and proliferation, encapsulation can be
used to enhance the biocompatibility of the implant surface, improved integration with host tissues,
and limit the degradation particles from entering the tissue-biomaterial interface. Recently, a novel
biocompatible packaging process for implantable electronic systems is described, combining excellent
biocompatibility and hermeticity with extreme miniaturization [165]. In addition to trying to find ways
to enhance the stability and biocompatibility of electronic devices when functioning in vivo, a recent
trend concerns with the development of fully resorbable electronic systems. These systems are
specifically designed to remain stable for a pre-defined length of time, during which the implanted
device will perform its functions. Once the task is completed, the implantable device will break down
under the influence of the physiological environment in which the implant resides. The key
competencies this technology aims to achieve are precisely controlled degradation kinetics and
cyto- and tissue-compatibility of the degradation by-products. Resorbable devices are particularly
useful where the goal of the implantation is not to permanently replace a function, but to provide
temporary physical framework and stimulation to enable tissue restoration, medical diagnostics,
accurate spatio-temporal delivery of drugs and other molecules. Recently, Hwang et al. reported on the
development and in vitro study involving transient electronic devices based on silicon with a silk
substrate [166]. The developed devices had tunable electrical properties and degradation kinetics.
Devices comprising of Mg-based inductive coils, resistive doped Si NMs microheaters, and silk based
substrate and packaging were suggested as a bioresorbable tool for non-antibiotic thermal therapy to
control surgical site infection (Figure 3). It is important to understand, however, that biodegradation
behavior of resorbable material is largely affected by its environment. By means of different
physico-chemical parameters (e.g., pH, ion concentrations, oxygen), the biological environment
directly affects the properties and the behavior of the implant material. Concomitantly, the implant, as
an introduced foreign body, incites an immunological response and influences the surrounding tissues
due to the direct and intimate contact. For example, Mg wires undergo extensive biocorrosion when
placed in the rat arterial wall, whereas little corrosion was observed for those Mg wires exposed to
blood in the arterial lumen for 3 weeks [167]. Therefore, further investigation regarding the
cytocompatibility and biocompatibility of the materials used in the bioresorbable electronic devices,
and the potential toxicity of their degradation by-products are required in order to bring these exciting
technologies to clinical applications.

Electronics 2013, 2 22
Figure 3. In vivo evaluations and example of a transient bioresorbable device for thermal
therapy. (A) Images of an implanted and sutured transient electronics platform located in
the subdermal dorsal region of a BALB/c mouse. (B) Implant site and histological section
of tissue at the implant site, excised after 3 weeks, showing a partially resorbed region of
the silk film. (C) Resonant responses of an implanted transient RF metamaterial structure
before and after placement in a silk package, immediately after implantation and at several
time intervals thereafter. (D) Measured and calculated Q factor for the metamaterial, where
the results indicate transience dominated by the diffusion of biofluids through the silk
package. (E) Transient wireless device for thermal therapy, consisting of two resistors (red
outline) connected to a first wireless coil (70 MHz; outer coil) and a second resistor (blue
outline) connected to a second, independently addressable, wireless coil (140 MHz; inner
coil). The inset shows a thermal image of this device coupled with a primary coil operating
at two frequencies, to drive both the inner and outer coils simultaneously. (F) Primary coil
next to a sutured implant site for a transient thermal therapy device, with the inset showing
an image of a device. (G) Thermal image collected while wirelessly powering the device
through the skin; the results show a hot spot (5 °C above background) at the expected
location, with a magnified view in the inset [166].
7. Conclusions
Since the days of first pacemakers, implantable electronics systems have undergone a major
transformation. The advent of micro-, nano- and molecular scale technologies have brought upon
tremendous miniaturization of all components of the indwelling module, from sensors to actuators and

Electronics 2013, 2 23
electrodes. The very-large-scale integration enabled small yet efficient low-power implantable
microsystems that can support increasingly complex processing. The advancement of battery
technologies has allowed for the development of long-term implantable devices with high reliability,
multiple functions and improved performance. Device powering via short range wireless links has
been used to extend the life-time of the electronic system, with in vivo energy generation and
harvesting being an area of active research. The use of wireless communication technologies has
extended from biomedical research to clinical health care, by enabling remote monitoring and control.
Significant progress has been made with regard to the stability and biocompatibility of the packaging
and encapsulation used to shield the indwelling electronics from the aggressive physiological
environment. Additional functionalities, such as ability to retard bacterial attachment or encourage
tissue growth, have been imparted onto these encapsulants. Together, these technologies have
contributed significantly to the quality of life of the patient, preventing critical incidents and
decreasing patient mortality.
Given the ageing population, increased longevity and an ever increasing number of patients
admitted into hospital care every year, the technologies that support individualized out-of-clinic
automated monitoring and patient status-responsive treatment will continue to be an area of great
interest and concentrated research effort. Further miniaturization of the sensing and stimulating devices
will enable on-organ monitoring and highly-specific treatment delivery, without compromising normal
functioning of surrounding organs and tissues. The advancement of closed loop systems will facilitate
simultaneous stimulation and high-resolution sensing of both natural and evoked activity, with utility
in intricate surgical procedures and neuromodulation. In addition to more sophisticated
neuroprosthetics and artificial organs that will improve patient survival and quality of life, further
developments in brain-computer interfacing will enhance our ability to investigate and alter cognitive
or sensory-motor functions in humans.
Acknowledgments
The authors acknowledge the JCU Collaboration across boundary grant and other internal grants.
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