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MEMS: Enabled Drug Delivery Systems
Angelica Cobo, Roya Sheybani, and Ellis Meng*a micropump for controlled insulindelivery in diabetic patients.[6] The goalof micropump therapy was control of glucose levels without having to rely on frequent needle injections.
Oral and intravenous injection are systemic drug delivery methods that haveenabled significant medical advances,
however, large doses are required toachieve the desired therapeutic drug concentration at the target site within the
body which can produce harmful side
effects.[7] For example, systemically administered drugs such as anti-inflammatorysteroids, anti-cancer and anti-fertility are associated with severe
unintended side effects.[8] The effectiveness of the therapeutic
agent depends on the method of administration, therefore,
treatments can be optimized by improving drug delivery systems. Localized delivery is particularly important when using
pharmaceuticals that have short half-lives in vivo, such as proteins and peptides, or when administering drugs that are highly
toxic. Site specific drug delivery can achieve therapeutic drug
levels at the target site and limit exposure to healthy tissues
to reduce the occurrence of side effects. The effectiveness of adrug therapy is also dependent on the timing of administration
as drug function is often tied to periodic biological fluctuations
such as circadian rhythms.[9]MEMS is a rapidly growing field that allows the batch production of small devices by utilizing fabrication techniquesborrowed from the semiconductor industry. In the last few decades, MEMS technologies have enabled advances in infusion
devices that can pump, sense, mix, and control fluid volumes.
Such devices include biocapsules, microreservoirs, microneedles, microparticles, and implantable pumps. These devices
provide new approaches to drug delivery not possible withconventional methods and improved access to additional drug
administration routes (Figure 1).[10] The majority of MEMS
drug delivery systems consist of three components: drugchamber, drug release mechanism, and packaging.[11] Drugis transferred from the drug chamber to a specific location in
the body using a variety of actuation mechanisms that afford
accuracy, precision, and reliability. Additionally, drug delivery
systems can be integrated with microvalves and microsensors
to further regulate drug flow and achieve feedback control,
respectively.
This Review examines both non-powered and powered(hereby referred as micropumps) MEMS drug delivery devices.
Important features, operation mechanisms, and examples ofdevices under development are described. The integration of
dosing control systems, examples of commercially available
MEMS-enabled drug delivery devices, current challenges, andfuture outlooks are also discussed.

Drug delivery systems play a crucial role in the treatment and managementof medical conditions. Microelectromechanical systems (MEMS) technologies have allowed the development of advanced miniaturized devices for
medical and biological applications. This Review presents the use of MEMS
technologies to produce drug delivery devices detailing the delivery mechanisms, device formats employed, and various biomedical applications. The
integration of dosing control systems, examples of commercially available
microtechnology-enabled drug delivery devices, remaining challenges, andfuture outlook are also discussed.

1. Introduction
Efficient and effective drug administration is essential to
improve the clinical outcomes in the treatment and management of medical conditions. Drug delivery technologies haveexisted since ancient times; Egyptian physicians employedoral tablets and ointments over 4000 years ago.[1] Intravenous drug administration technologies evolved after Harvey’s
description of the circulatory system in 1657. Access to thecirculation was achieved with the development of a cannulaand gravity flow device (hollow needle) by Frances Rynd in
1845.[2,3] Controlled drug delivery dates back to the 1960s
when Judah Folkman from Harvard University used a Silastic(silicone rubber) anterio-venous shunt to circulate rabbit
blood and noticed that the rabbit would fall asleep when the
shunt was exposed to anesthetic gases (e.g., nitrous oxide,
halothane, and cyclopropane).[4] He demonstrated that the
diffusion rate of these gases could be controlled by the wall
thickness of the tubing. Folkman and Long also reported
controlled release of therapeutic agents encased in a silicone
rubber capsule for local or systemic drug delivery in vivo.[5]Pumping technologies were introduced in the 20th century. Thomas and Bessman were among the first to develop

A. Cobo, R. Sheybani, Dr. E. Meng
Department of Biomedical Engineering
Viterbi School of Engineering
University of Southern California
1042 Downey Way
DRB-140, Los Angeles, CA 90089-1111, USA
E-mail: ellis.meng@usc.edu
Dr. E. Meng
Department of Electrical Engineering
Viterbi School of Engineering
University of Southern California
3651 Watt Way
VHE-602, Los Angeles, CA 90089-0241, USA

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2. Non-powered MEMS Drug Delivery Devices
MEMS-based drug delivery devices need not require electrical
power to operate; drug payload can be delivered by osmotic or
diffusive transport, or in response to an environmental stimulus.[12] These systems are generally simple in format andtherefore easy to fabricate. Since an external power source is
not required, it is possible to for such devices to have minimal
footprint.[13] However, the drug release mechanisms often limit
performance to low release rates and slow response.[14] The drugdelivery rate is pre-determined by the selected materials, fabrication methods, or drug formulation, and is highly dependent
on the properties of the fluid transported and the delivery site’senvironmental properties (e.g., temperature, pH, saccharide
concentration, and antigen concentration) that fluctuate overthe course of treatment.[13] Usually this rate of delivery cannot
be modified or stopped after the device is placed, implanted, or
injected into the targeted area.[15]Angelica M. Cobo received
her B.S. degree inBiomedical Engineeringfrom Florida International
University in 2008. She wasa research assistant in the
Bioengineering Laboratory
at the Diabetes ResearchInstitute where she focusedon the development of abioactive scaffold material
for use in cellular replacement therapies to treat
Type 1 Diabetes Mellitus. In 2012, she received her M.S.
in Biomedical Engineering from University of Southern
California (USC). She is currently a PhD student in
Biomedical Engineering at USC where she is working on thedevelopment of implantable MEMS drug delivery devices for
management of chronic diseases in preclinical research.

2.1. Diffusion-Based
Diffusion-based systems operate by passive diffusion of drug
through its polymer reservoir encapsulation as shown in
Figure 2. The rate of diffusion is controlled by the size of the
drug molecule, the membrane structure, and the pore sizeor space between the polymer chains.[16] Although generally
simple in design and small in size, these systems suffer fromlack of precise control over drug release rate[13] and difficulty inachieving reproducibility.[17]This technology has been used for oral and transdermaldelivery systems, as well as for ocular, vascular, and oncology
implants.[11,16,18,19] Diffusion-based drug delivery systems are
created from either non-biodegradable solids such as polyvinyl
alcohol (PVA), ethylene-vinyl acetate (EVA), and polysulfone capillary fiber (PCF), or from biodegradable materials such as polylactic acid (PLA), polyglycolide acid (PGA), polylactic-co-glycolicacid (PLGA), polycaprolactone (PCL), and polyanhydride.[20] Foreach specific application, the properties of the materials, as wellas the intended drug, should be carefully considered.
Wang et al. studied controlled release of ethacrynic acidfrom PLGA thin films (3 mm × 3 mm × 0.1 mm) for glaucoma
treatment.[21] Less than 10% of the initial drug concentration
remained within the remaining polymer film on day seven.[20]Desai et al. microfabricated silicon-based biocapsules for encapsulation of rat neonatal pancreatic islets.[22,23] The biocapsule’s
membrane provided immunoisolation to the transplanted cells
while allowing them to release insulin into the surrounding
environment.

2.2. Osmotic
In osmotic devices, flow is generated when differing soluteconcentrations across a semi-permeable membrane lead to thedevelopment of a hydrostatic pressure (Figure 3). This pressure difference is used to displace drug.[24] The device footprint
can be minimized by using the drug as the solute that will
dissolve when in contact with bodily fluids. Alternatively, the2wileyonlinelibrary.com

Roya Sheybani received her
B.S. (2008) and M.S. (2009)degrees in biomedical engineering from the Universityof Southern California (USC),
Los Angeles, where she iscurrently a PhD candidate inbiomedical engineering. She
is developing closed-loop
implantable wireless MEMS
drug delivery devices for management of chronic diseases.
Ellis Meng received her B.S.
degree in engineering andapplied science and her M.S.
and Ph.D. degrees in electricalengineering from the California
Institute of Technology(Caltech), Pasadena, in 1997,
1998, and 2003, respectively.
She is a Professor of Biomedical
Engineering and ElectricalEngineering at the Universityof Southern California, Los
Angeles where she has been
since 2004. She directs the
Biomedical Microsystems Laboratory which specializes in biomicroelectromechanical systems, implantable biomedical microdevices, neural interfaces, and microfluidics. She is a recipientof the NSF CAREER, the Wallace H. Coulter Foundation Early
Career Translational Research, and the ASEE Curtis W. McGraw
Research Awards. In 2009, she was recognized as one of theTR35 Technology Review Young Innovators under 35. E. M. has asignificant financial interest in Fluid Synchrony LLC (CA, USA).
drug can be placed in an adjacent reservoir separated from thesolute by a piston or flexible membrane. This would allow for© 2015 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim

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Figure 3. Schematic diagram of an osmotic drug delivery device.

Figure 1. MEMS enabled technologies provide new approaches for sitespecific drug delivery.

independent control over the osmotic pressure, and hence thedrug delivery rate.[25] Osmotic devices do not require external
power and are attractive due to their simplicity, robustness, andsmall footprint, however, they suffer from low flow rate andslow response with long delay; the rate of delivery cannot bealtered or stopped once started.[26] Also, in some instances, thesemi-permeable membrane may detach from the rest of thesystem after a certain period of time, which could halt drugdelivery or lead to immediate release of the entire remaining
drug payload.[27]Figure 2. Schematic diagram of a diffusion-based drug delivery device.

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LiRIS is a small, flexible osmotic system originally introduced by Lee and Cima.[28] The device is based on a doublelumen silicone tube that operates as a semi-permeable
membrane releasing lidocaine when placed in the bladder(Figure 4).[24] The device, recently acquired by Allergan fromTARIS Biomedical, is currently in phase 2 clinical development
for the localized treatment of interstitial cystitis/bladder pain
syndrome.[29]A biodegradable osmotic micropump developed byresearchers at Stanford University allows for the controlled
release of basic fibroblast growth factor to aid bone repair(Figure 5). The planar device fabricated from PLGA and polyethylene glycol (PEG) is capable of delivering 40 ng/day for
four weeks.[30]Herrlich et al. developed BuccalDose, a disposable intraoral
drug delivery cartridge for the self-medicated treatment of
Parkinson’s disease. The disposable cartridge is magneticallyattached into the receptacle of a partially removable dental prosthesis, allowing for a constant release of dopamine agonists
to the buccal mucosa and subsequently to the bloodstream
through osmosis. Precise release rates over 97% of the system’s storage capacity with a rate deviation of only 1.1% can beachieved.[31] however, device performance may be affected by
the degree of saliva secretion.[24]Figure 4. Photograph of LiRIS: a small flexible osmotic system for drugdelivery to the bladder. Reprinted with permission.[24] Copyright 2012,
Elsevier.

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Figure 5. Photograph of osmotic device and the zoomed-in view of a
micro-channel developed by Ryu et al. Reprinted with permission.[30]Copyright 2007, Elsevier.

Figure 7. Schematic diagram of a microneedle drug delivery device.

2.4. Microneedles
Other examples of osmotic drug delivery devices can befound in the literature.[24,32,33]

2.3. Responsive Hydrogels
Responsive hydrogels are biocompatible and can be specifically engineered to respond a variety environmental stimuli andmarker molecules within the blood such as temperature, pH,saccharide concentration, and antigen concentration. During
the course of treatment, these factors fluctuate and result in thealtered response of the hydrogel and therefore the delivery rateof drug as shown in Figure 6.[13] Hydrogel-based systems may
suffer from non-continuous and decreasing release rates. Thismay be partially mitigated by limiting the quantity of water toless than required by the hydrogel for swelling to equilibrium.[24]Eddington and Beebe developed a disposable infusion system
for the delivery of protein therapeutics.[34] Under specificconditions, the pH-responsive hydrogel expands and the drug
is expelled from an adjacent reservoir at a rate of 2 µL/h for
12 h.[34] Chang et al., developed a self-healable chitosan (CS)/
polyvinyl alcohol (PVA) hydrogel for anti-tumor therapy.[35] Thissystem achieved continuous and controllable drug release at pH
5.0. Temperature responsive hydrogels where used to constructgrippers composed of poly(N-isopropylacrylamide-co-acrylicacid) and poly(propylene fumarate) for sustained drug releasein the gastrointestinal tract.[36] Drug release from the grippers
into the tissue occurred at body temperatures above 32 °C.

Microneedles, fashioned into syringes and patches, allow fortransdermal and intradermal drug and vaccine delivery without
pain.[16] Using microfabrication techniques, needles can be precisely engineered to penetrate the stratum corneum without
reaching the underlying nerve cells. Drugs are delivered into
the dermis, from which the drug can access the circulatory
system relatively quickly.[26] In general, microneedles can be
categorized as solid microneedles for tissue pretreatment,
drug-coated microneedles, dissolving microneedles, and hollow
microneedles (as shown in Figure 7. Solid microneedles arefabricated from silicon, metal (e.g., titanium, stainless steel,etc.), and select polymers (e.g., polycarbonate (PC), polymethyl
methacrylate (PMMA), etc.). These solid microneedles can beused to increase skin permeability prior to topical applicationof a drug. Solid microneedles can be also coated with watersoluble drug that would dissolve under the skin prior to theremoval of the microneedle. Alternatively, polymer and sugarbased microneedles have been developed to completely dissolvein the skin after insertion. Typically drugs are encapsulated
inside the microneedle and are released after placement in
the skin. The fourth class of microneedles are hollow and provide a defined conduit for drug delivery into the skin or othertissue.[37] Transdermal drug delivery reduces the need for alarge reservoir size, as a surface mounted device can be easily
replaced or refilled. However, penetration with microneedles
remains relatively shallow within the skin,[38] and will not allow
for site-specific drug delivery.
A number of microneedle-based devices have been developed for medical and cosmetic use.[26,37] For instance, bothcoated[39] and dissolving[40] microneedles (Figure 8a) were used
to deliver influenza vaccine. Hollow microneedles were used todeliver insulin in subjects with type 1 diabetes at 1 mm depth.[41]Microneedles have also been used to deliver drugs into other biological tissues such as the eye[42] and the nasal mucosa.[43] They
can also be incorporated onto vascular stents (Figure 8b).[44]2.5. Fluorocarbon Propellant-Driven

Figure 6. Schematic diagram of a responsive hydrogel drug delivery
device.

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In fluorocarbon propellant-driven pumps, the inner reservoirchamber contains the drug while an outer chamber contains
a fluorocarbon liquid that once vaporized can exert a vapor
pressure well above atmospheric pressure at body temperature as shown in Figure 9. Flow rate is modified by altering the© 2015 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim

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length of the delivery catheter that acts as a flow regulator, orby adjusting the drug concentration. These pumps are sensitive to changes in ambient pressure and temperature.[45] Whilea macro-sized implantable insulin delivery pump was developed[46] by Blackshear et al., the ban of fluorocarbon propellants[47] halted research in this area.

3. Powered MEMS Drug Delivery Devices
Powered MEMS drug delivery devices can be classified into
two categories: non-mechanical and mechanical.[48] Nonmechanical micropumps transform non-mechanical energy
into kinetic momentum to drive fluid out of the reservoir. Thisphenomenon is well suited for implementation in pumps at
the microscale.[49] Non-mechanical micropumps do not require
moving parts, resulting in simpler structures and fabrication
techniques. However, these actuation mechanisms are not suitable or have not been successfully used in drug delivery devices
because their driving effect and performance (e.g., flow rate,
response time, and pressure generation) are inferior whencompared to mechanical actuation.[14,49]
Mechanical micropumps utilize moving parts to generate oscillatory or rotational pressure forces on the working
fluid to displace it. Three movement mechanisms have been
employed by mechanical micropumps: reciprocating, rotatory,

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Figure 8. a) Dissolving polymer microneedles developed to deliver influenza
vaccine. Reprinted with permission.[40] Copyright 2010, Nature PublishingGroup. b) silicon microneedles developed for intravascular drug delivery.
Reprinted with permission.[44] Copyright 2000, John Wiley and Sons.

and peristaltic.[50] The majority of micropumps reported utilize reciprocating motion. This type of micropump requires apumping chamber coupled to a physical actuator and a moving
surface (diaphragm), and check valves to control fluid flow
during the supply and pumping cycles. During operation, theactuator mechanism acts on the diaphragm resulting in an
overpressure on the drug that displaces it from the pumping
chamber. Figure 10 depicts the structure and operation of ageneric mechanical micropump. The majority of mechanicalmicropumps have fast response time, large actuation force,
good biocompatibility, but are limited by high driving voltages
and complex fabrication processes. Popular mechanical micropump actuation mechanisms are discussed here.

3.1. Electromagnetic
A magnetically actuated micropump consists of a chamberwith a permanent magnet embedded into a flexible membrane
and a set of drive coils (as shown in Figure 11).[51] Current flow
through the coils generate a magnetic field that induces repulsive or attractive magnetic force between the micro coils andthe permanent magnet resulting in movement of the membrane. The force generated by this actuator is dependent on thenumber of turns in the coils and the applied electrical current.
Electromagnetic actuation requires low operating voltages.[49]However, these actuators consume significant power, generate
heat, and can be difficult to miniaturize.
An electromagnetically actuated device was developed tocontrol on-demand delivery of an antiproliferation drug, docetaxel (DTX), for the treatment of diabetic retinopathy. The
device consists of a polydimethylsiloxane (PDMS) microreservoir sealed by a magnetic membrane with a small aperture.
The membrane is a composite of iron oxide nanoparticles ina PDMS matrix (Figure 12). An external 255 mT magnetic
field deformed the magnetic PDMS membrane resulting in a
release rate of 171 ± 16.7 ng per actuation.[52] Kwon et al. utilized electromagnetic actuation to propel a microrobot intended
for site specific drug delivery in diseased blood vessels. The
electromagnetic systems consists of a Helmholtz and Maxwell
electric coils for manipulation of the microrobot movement.
This microactuator is capable of delivering flow rates up to
1.98 mL/s.[53]3.2. Piezoelectric

Figure 9. Schematic diagram of a fluorocarbon propellant driven drugdelivery device.

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Piezoelectric materials can undergo deformation with anapplied electrical current and when attached to a membrane toprovide usable displacement for expelling fluid from a pumpchamber.[14,54] The micropump diagram is shown in Figure 13.
Such actuators possess simple structure. However, actuators are limited to available piezoelectric materials and highdriving voltage of up to 200 V is required to generate a useful
deformation.
Liu et al. developed a disposable piezoelectric micropump for
insulin delivery and close-loop monitoring of glucose concentration (Figure 14).[55] This pump consisted of four chambers
in serial connection. The pump was made out of biocompatible© 2015 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim

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1.7 g. The device was capable of delivering
sixteen 0.25 µL doses and each dosage only
required 3.375 Ws of electrical power.[58]Figure 10. Reciprocating micropump operation. a) Configuration and major components ofa typical reciprocating diagram pump. b) The drug chamber is emptied during the pumping
mode and c) filled during the supply mode.

materials and contained a volume of drug reservoir of 3.2 mL.
A maximum pressure of 22 kPa can be obtained with an applied
voltage of 36 V and a driving frequency of 200 Hz. Junwu et al.
developed a high-frequency piezoelectric cantilever-valve micropump for site specific drug delivery. The micropump achieved amaximum flow rate of 3.5 mL/min at back pressure of 27 kPa.
The micropump’s flow rate performance was dependent on the
cantilever valve dimensions.[56]3.3. Thermal/Shape Memory Alloy (SMA)
SMAs are characterized by pseudoelasticity and can return totheir original shape after a heating/cooling cycle.[14] A commonly used SMA is titanium/nickel alloy (TiNi), which has
high recoverable strain and is capable of withstanding large
pumping rates and high operating pressures. The micropump
diagram is shown in Figure 15. Advantages include high stress(>200 MPa), high force-to-volume ratio, chemical resistance,
biocompatibility, and long operating cycles. These are balancedby the limited selection of SMA materials, high power consumption, thin-film shape memory training, and uncontrolled
deformation due to temperature sensitivity.[51]Xu et al. reported a micropump actuated by a NiTi/Si composite diaphragm. The micropump contained two silicon flap
valves and its outer dimension were 6 mm × 6 mm × 1.5 mm.
The maximum flow rate was 340 µL/min at a frequency of
50–60 Hz under a back pressure of 1 hPa. The pump achieved
more than 4 × 107 working cycles.[57] Spieth et al. presented a
T-shaped SMA actuated micropump for microinfusions into thecentral nervous system of freely moving animals. The device
was operated by thermally expandable microspheres embedded
in a PDMS matrix (Figure 16). The total device weight was

3.4. Electrostatic

Reciprocating membrane deflection can beproduced by controlling the electrostatic
attraction between closely spaced parallel plates through application of periodically switching voltages.[48] In this manner,fluid in the reservoir can be displaced as shown in Figure 17.
Electrostatic actuation features low power consumption, andsimple fabrication. The main limitations are high applied voltages, small stroke (membrane deflection), and pumping only of
non-conductive fluids.
Zengerle et al. developed an electrostatiacally actuated bidirectional micropump for miniaturized chemical analysis systems.
The silicon pump incorporated two passive check valves. The
outer dimension of the pump was 7 mm × 7 mm × 2 mm. The
actuation frequencies determined the direction of the flow in thismicropump (e.g., 2–6 kHz results in reverse pump operation).
The maximum flow rate in the forward direction was 850 µL/min
at 200 V applied voltage. In the reverse direction the maximum
flow rate was 400 µL/min. The maximum back pressure in the forward and reverse direction was 31 kPa and 7 kPa, respectively.[59]Bourouina et al. designed an electrostatic micropump for drugdelivery applications where a very small flow rate was required.
The overall device size was 5 mm × 5 mm and was capable of flow
rates in the 10–100 nL/min with a working voltage of 10 V.[60]3.5. Thermopneumatic
Thermopneumatic actuation, shown in Figure 18 consists of athermally expandable medium (either a gas or liquid) enclosed
in a sealed cavity that can be heated or cooled down to induce
a pressure change in the cavity.[49,61] This pressure change isused to deform a diaphragm. Resistive heating can be utilized
to achieve the temperature change. This actuation mechanism
is limited by slow response time and low efficiency.
Zimmermann et al. developed a thermopneumatically actuated planar micropump with two in-plane flap valves for highpressure and flow rate applications such as cryogenic and drugdelivery systems. A maximum flow rate of 9 µL/min and pressure of 16 kPa was achieved with a power consumption of
180 mW.[62] A thermopneumatic micropump was also reported
by Mousoulis et al. for transdermal drug delivery applications(Figure 19). The device consisted of PDMS layers on a silicon
substrate. This substrate was thermally conductive and utilized body heat to actuate the pump. A perfluoro compound(FC-3284) was selected as the working fluid due to its low
boiling point and large vapor pressure values. The maximum
volumetric flow rate and back pressure of the pump were
28.8 µL/min and 28.9 kPa, respectively.[63]3.6. Bimetallic

Figure 11. Schematic diagram of an electromagnetic drug delivery device.

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can produce displacement in response to a temperature change(as shown in Figure 20).[48] This actuation requires relatively
low voltages, however, it has low frequency response.
Zhan et al. developed a biometallic micropump operated
by an aluminum-silicon diaphragm. The overall device sizeis 6 mm × 6 mm × 1 mm. The pump is capable of delivering
45 µL/min and maximum back pressure of 12 kPa at a frequency of 0.5 Hz and driving voltage of 5.5 V.[65] A silicon micropump operated by bimetallic and thermopneumatic action wasreported by Zou et al.[66] The bimetallic actuation consisted ofan aluminum and silicon membrane. Finite element moldeling(FEM) simulations showed that this micropump was capable ofa flow rate of 5.6 µL/s when the open pressure of the valve was
0.5 kPa.

3.7. Ionic Conductive Polymer Film (ICPF)
ICPF consists of two metal electrodes with a core layer of aperfluorosulfonic acid polymer or Nafion/silica as shown in
Figure 21.[48] One end of the electroactive polymer diaphragm isfixed allowing a bidirectional bending motion of the film whenan alternating voltage is applied across the electrodes. The
advantages of ICPF are low driving voltage, and ability to workin aqueous environments. A major drawback is lack of repeatability in batch fabrication processes due to the complex fabrication of the ICPF actuator.

Figure 13. Schematic diagram of a piezoelectric drug delivery device.

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Figure 12. Conceptual drawing of electromagnetic drug delivery device by Pirmoradi et al.
Reprinted with permission. [52] Copyright 2011, Royal Society of Chemistry.

An ICPF micropump with two one-way
valves was proposed as a servo actuator for
active wire guiding in a micro active catheter system. Simulation experiments showedthat this active catheter system was suitablefor intracavity operations. The overall pumpsize was 23 mm in length and 13 mm indiameter. The maximum flow rate of the
micropump was 37.8 µL/min at a frequencyof 15 Hz and with an applied voltage of
1.5 V.[67,68] Hiraoka et al. developed an ICPF
actuated micropump for genotyping applications (Figure 22). The actuator consistedof stacked conductive polymer layers glued
together with epoxy dots. The average flowrate for this micropump was approximately
1.5 µL/min.[69]3.8. Electrochemical
An applied electrical current to a pair of electrodes in a water
filled chamber causes electrolysis of the water into hydrogenand oxygen gases as shown in Figure 23. This gas generationinduces an increase in pressure providing the driving force todispense the fluid[49,54] Once the current is turned off, gases
recombine to water in the presence of a catalyst such as platinum.[71] These actuators have a relatively simple structure, andare easily integrated with other microfluidic devices. However, a
low rate of bubble generation might result in partially collapsed
and recombined bubbles which can affect drug release.[14]Meng et al. developed site specific drug delivery micropumps
operated by an electrolysis actuator (Figure 24). These micropumps were intended for chronic drug treatments (e.g., cancer
and ocular diseases) in small laboratory animals.[72,73] The anticancer micropump could deliver a variety of working fluids
within a large range of flow rates (0.33–141.9 µL/min) by controlling the amount of applied current (1–10 mA).[71,74]

4. MEMS Dosing Control Systems
The need for patient tailored drug delivery therapies has
increased demand for more advanced and precise drug delivery
systems over the past 20 years.[75] Real time monitoring ofpump performance can be achieved with the incorporation ofclosed-loop feedback systems. Physical sensors can be utilized
to provide information on pressure, flow rate, dose size andstate of the pump. Medtronic’s FDA-approved external artificial
pancreas (MiniMed 530G with Enlite) combines insulin pumptherapy with continuous glucose monitoring. In this system,
glucose sensors suspend insulin delivery when the glucose
levels fall below a preset value.[76] The Medallion Therapeutics,
Inc. (Minneapolis-St. Paul, MN) is developing the only implantable drug delivery system that incorporates pressure sensors
to monitor delivery. The pump is currently in clinical trials toassess the pressure sensor.[77]In most of the current drug delivery devices, failure is determined only when the patient has shown physical side effects.

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complex fabrication techniques and theiroperation was dependent on the pumping
mechanism.[84]Thermal flow measurements are moresuitable for low volume liquid flow and gas
flow monitoring. Thermal flow sensors relyon heat transfer principle to determine fluid
flow. This type of sensors are characterized byhigh sensitivity, simple structure and implementation, and low power consumption,[81]however,they are limited by sensor drift andFigure 14. a) 3-D assembly schematic of micropump prototype and b) photograph of in vivo
animal test disposable piezoelectric micropump for insulin delivery by Liu et al. Reprinted with dependence on chemical properties of thepermission.[55] Copyright 2010, Elsevier.
fluid.[85] Meng et al. developed a thermal flow
sensing array that allows measurements ofseveral flow parameters and was constructed of biocompatible
This late indication of failure could lead to serious health commaterials. This array was able to measure flow rates as low asplications or even death.[78,79] Current methods for monitoring
0.5 µL/min with minimal heating of the working fluid.[86]dosage levels such as blood measurements, nuclear imaging
or direct observation lack resolution, lack accuracy and do not
Electrochemical impedance sensing is a promising method
offer real-time monitoring.[80]for flow measurements. It is capable of tracking blockage, ordose volume in real-time. EI sensors are highly sensitive, easy
to fabricate and incorporate into microfluidic systems, andrequire low power. Bohm et al. utilized EI measurements as
4.1. Sensors
a closed-loop feedback mechanism for microfluidic applications.[87] This system was able to track the volume of a gas/
Over the last few decades, MEMS technology has allowed theintegration of miniaturized sensors into microfluidic systems
liquid fraction contained in an electrolysis reservoir. Only voland micropumps for flow measurements. These sensors areumes lower than 1.5 µL were tracked, therefore, it is not suitcrucial for monitoring gas or fluid flow.[81] Flow sensors employ
able for implantable drug delivery devices where the dosevolume is larger. One challenge of EI sensors is drift, which cana variety of physical sensing techniques such as electrical, radiaresult in errors in flow and volume measurements. This can betion, magnetic, mechanical, thermal, or biochemical.[82,83] Incoraddressed by adding a third electrode to provide a more stableporation of traditional flow sensors is challenging due to complex
reference and minimize drift.[88]fabrication and integration methods, size, biocompatibility,
external calibration, and power consumption.[83,84] Dose volume,
flow rate, or catheter occlusion measurements through microfluidic devices can be accomplished with thermal, pressure or elec4.2. Valvestrochemical impedance (EI) sensors (among other techniques).
Pressure sensors integrated into a microfluidic device can
Advanced drug delivery devices incorporate valves for flow
be used to indirectly measure flow rate by measuring the difregulation, sealing, and on/off switching.[89] Valves are crucial
ferential pressure inside a reservoir. Li et al. incorporated piefor accurate dosing, and prevention of backflow of biologicalzoresistive sensors in a drug delivery system to measure flow
fluids into the device. Microvalves are classified as active (powrate (Figure 25). These sensors can also be used to calculateered) or passive (unpowered) valves. They can both be operated
residual reservoir volume and catheter blockage. While the senwith mechanical or non-mechanical moving parts.[54] Passive
sors achieved high sensitivity (≈698 ppm/kPa), they required
valves are operated by a pressure difference between the inlet

Figure 15. Schematic diagram of a thermal/SMA drug delivery device.

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Figure 16. Photograph of SMA drug delivery device by Spieth et al.
Reprinted with permission.[58] Copyright 2012, Springer.

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Figure 19. Photograph of thermopneumatic drug delivery device.
Reprinted with permission.[64] Copyright 2012, Royal Society of Chemistry.
Figure 17. Schematic diagram of an electrostatic drug delivery device.

and outlet of the valve. The actuation mechanism for active
valves can be solenoid plunger, piezoelectric, electrostatic,
thermopneumatic, electromagnetic, among others.[90] Several
parameters should be taken into account when designing or
selecting a microvalve such as dead volume, leakage due to an
applied reverse pressure, resistance to flow, power consumption,size, and response time. Active valves offer improved performance, but are limited by their complexity and fabrication cost.
Piezoelectric microvalves were used to control fluid flow rate ina drug delivery pump.[91] This dual valve system was employed
to control and mix drug flows from two separate pressurized
reservoirs. Flow rate through the microvalves was controlled by
piezoresistive pressure sensors embedded in the MEMS valves
resulting in a flow rate range of 0.51 to 2.30 mL/h. Sim et al.
reported the development of a phase-change type micropumpwith two aluminum flap passive check valves for fluid rectification.[92] A maximum flow rate of 6.1 µL/min was achieved witha maximum back pressure at zero flow rate of 69 kPa.

of diabetic macular edema. The cylindrical implant (3.5 mm
long and 0.37 mm in diameter) consists of a polyimide tube
reservoir containing 190 µg fluocinolone in a PVA matrix
capped with rate controlling membranes. It is inserted into thevitreous of the eye with a 25 G needle during an outpatient procedure[16] and is intended for 36 months of continuous delivery.
As the implant is non-bioerodable, it remains in the vitreous
cavity even after the termination of drug release and patientsrequiring repeated injections may have multiple devices
trapped in the vitreous base for an indefinite period of time.[93]Iluvien was approved by US Food and Drug Administration(FDA) on September 26, 2014.

5.2. Duros Pump and Durin Implant

Iluvien is a diffusion-based, non-bioerodable implant developedby Alimera Sciences (Alpharetta, GA) for long-term treatment

The Duros pump is a mini osmotic pump (3.8 mm in diameterand 44 mm long) developed by Durect Corporation (Cupertino,
California). The pump is meant for subcutaneous systematic
drug delivery. It can deliver up to 1000 mg of concentrated drug
for up to a year at a constant rate (±10%), ultimately releasing
greater than 95% of its drug content. The device was FDA
approved in 2000 for one year subcutaneous delivery of prostate cancer therapy. However, the product was discontinued andclinical trials for new indications have been suspended pending
redesign of the delivery system in order to address performanceissues caused by premature shutdown.[16,94] Durect Corporationis also developing Durin, an injectable biodegradable implant,
in which drug release is controlled by drug content, polymer

Figure 18. Schematic diagram of a thermopneumatic drug delivery
device.

Figure 20. Schematic diagram of a bimetallic drug delivery device.

5. Some Examples of Commercial Drug Delivery
Pumps
5.1. Iluvien

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composition and molecular weight, device geometry, and the
manufacturing process.[95]5.5. OmniPod

REVIEWS
Figure 21. Schematic diagram of an ICPF drug delivery device.

drug payload is initiated when thermal actuation removes themembrane seal from the reservoir (Figure 26). As a resultdelivered dose can be precisely controlled in single reservoir
increments and dosing can be terminated without the needfor device extraction.[99,100] In 2012, MicroCHIPS (Lexington,MA) reported successful human clinical trials of subcutaneousdelivery of an anabolic agent for the treatment of osteoporosis.
Eight osteoporotic postmenopausal women were implantedwith the device for 4 months. Human parathyroid hormone
fragment (hPTH) doses were delivered from the device once
daily for up to 20 days. Dosing with the device produced similar
pharmacokinetics to multiple injections and had lower coefficients of variation.[101]5.3. JewelPump
Debiotech, a Swiss-based company, has developed a wearable
insulin delivery system. The device is driven by a battery-powered piezoelectric actuator that delivers insulin from a disposable reservoir worn as a patch on the skin for seven days.[96] The
system can be wirelessly programmed and monitored using
a cell phone.[97] While this system is in final stages of development in Europe, Debiotech is also in the process of further
miniaturizing the device so that it can be completely implanted
inside the body.[98]5.4. MicroCHIPS
MicroCHIPS is an implantable drug delivery device in which
100 individually controlled microreservoirs can be wirelessly activated for drug release.[16] Diffusion of each 300 nL

OmniPod is an SMA-actuated wearable mini-pump developed
by Insulet Corp (Billerica, MA). The device allows for subcutaneous delivery of insulin via a small cannula (Figure 27). A total
insulin volume of 2000 µL is stored in the disposable reservoir
and can be delivered in 0.5 µL boluses continuously for about
72 hours. Pump activation is achieved wirelessly using a wireless handheld device.[102] The second generation of the device
was approved by the FDA in 2013.

5.6. Prometra
The Prometra implantable pump system, developed by Flowonix, Medical (Mt. Olive, NJ), utilizes a positive pressure gas
expansion actuation design with battery powered valves for flow
regulation (Figure 28).[77] The device is intended for chronic
pain management and delivers morphine into intrathecal spacewith a flow rate of up to 28 mL per day[103] (overall 97.5% dose
accuracy). The 20 mL fixed-volume reservoir is refillable. The

Figure 22. ICPF microactuator by Hiraoka et al.: a) schematics of the packaged actuator; b) packaged actuator glued to a micro-fluidic system;c) photograph of the packaged actuator mounted on a fluidic structure. Reprinted with permission.[70] Copyright 2012, Elsevier.

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REVIEW

Figure 23. Schematic diagram of an electrochemical drug delivery device.

system (diameter of 71 mm, 20 mm thick, weighing 150 g), is
comprised of mostly immobile parts that allow for usage formore than 10 years.[104] The system was FDA approved in 2012.

6. Current Challenges
The medical device ecosystem is particularly complex in partdue to the involvement of many stakeholders with differing
motivations and requirements. As large medical device companies, pressured by the need to realize near-term payoffs, arebecoming more focused on introducing products that provide
incremental improvement over previous iterations, innovationin the form of new and disruptive technologies largely originates in start-ups, academic laboratories, and clinical research
environments which have the freedom to explore new highrisk concepts.[106] Successful commercialization requires thatdevices navigate lengthy and costly regulatory pathways.
Despite the numerous advantages of MEMS-based drugdelivery systems, there are several technical challenges thatremain. Appropriate medical packaging of drug delivery systems often require custom design and extensive engineering tosafely house the drug prior to its dispensation into the body. The
trend towards personalized drug therapy will likely dictate thatcomplex dosing regimens be available and made possible withintegrated electronics and physiological sensors for feedback

Figure 24. Photograph of electrochemical drug delivery micropump
developed by Meng et al.

Adv. Healthcare Mater. 2015,
DOI: 10.1002/adhm.201400772

Figure 25. Conceptual drawing of drug delivery device having piezoresistive pressure sensors developed by Li et al. Reprinted with permission.
Copyright 2012, Elsevier.

control of drug delivery.[11] Sensing technology capable of stablechronic interfaces with the body is a significant technical challenge; for example, many have attempted to develop implantable sensors for continuous glucose monitoring in the insulin
management of diabetes with limited success.[107] Advanced
systems may incorporate wireless electronics to remotely control device operation and allow monitoring of system performance by the patient, caregiver and healthcare provider. However, incorporation of wireless electronics adds additional packaging complexity and interference testing requirements, andposes device security risks.

6.1. US Food and Drug Administration Regulations
In the US, which is the largest medical device market in theworld, medical devices are regulated by the FDA based on thedefinition laid out in section 201(h) of the Federal Food Drug &Cosmetic (FD&C) Act. This definition divides medical devices
based on function and degree of risk into three categories. A
device’s class designation determines the appropriate regulatory pathway to gain federal clearance for the device to bemarketed in the US. Class III devices typically provide life-sustaining function and therefore pose the highest risk to healthin the event of failure (out of Class I, II, and III with Class I
devices posing the least risk). Implantable drug delivery devices
will largely be classified as Class III. As a result, such devices
typically take the premarket approval (PMA) route, are subjected to the most stringent controls and regulations, and must
be shown to be both safe for use and effective in their intended
clinical utility. New devices may receive a Class II designation
if shown to be similar to an existing approved device (predicate)within the class (premarket notification or 510(k) clearance) orby going through the de novo classification process introducedby the Food and Drug Administration Safety and Innovation
Act (FDASIA) introduced on July 9, 2012.[108] There is also ahumanitarian use device (HUD) regulatory pathway for devices
addressing rare and orphan conditions that affect or manifest© 2015 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim

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Figure 26. Photograph of MicroCHIPS micro-reservoir device. Reprinted with permission.[99] Copyright 2006, Nature Publishing Group.

in fewer than 4000 patients per year.[109] For these, instead ofrequiring both safety and efficacy, the FDA requires demonstration of safety and probable clinical benefit.
To acquire sufficient evidence for safety and efficacy, the FDA
will require non-clinical data (related to biocompatibility, toxicology, immunology, stress, wear, etc.) as well as preclinical andclinical studies. Successful completion of preclinical studies
will enable clinical trials in humans. At the conclusion of thetypically multiple clinical trial phases, a final regulatory review
occurs with the goal of achieving regulatory approval.[106] For a
more thorough discussion of medical device development, thereader is referred to the literature.[110]6.2. Funding and Translation from Academia to Market
Academic environments allow for the pursuit of high risk,
high reward projects, as well as, promoting close interaction
between clinicians and engineering faculty with know-how that
could lead to new solutions to urgent unmet medical needs.
However, many challenges such as sustained funding, multiinvestigator collaborations, and lack of knowledge and experience on successfully translating early stage inventions from the

Figure 27. Photograph of OmniPod insulin delivery device. Reprinted
with permission.[102] Copyright 2012, Elsevier.

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lab to the marketplace, must be overcome.[106] Uncertainty inthe regulatory approval environment for new devices, the focus
on short term pay offs, and changing health care laws (theMedical Device Tax Act took effect January 1, 2013 and charges
a 2.3% tax on revenues for sales of medical devices) have led
to a decline in early stage investment in medical devices traditionally sought by start-ups[111] (40% since 2007 according toPricewaterhouseCoopers and National Venture Capital Association).[112] Substantial investments now occur after clinical validation and regulatory approvals assuming that the clinical needand the market size are substantial.[113] New federal (National
Institutes of Health’s (NIH) National Center for Advancing
Translational Sciences (NCATS) program, and National Science
Foundation’s (NSF) Innovation Corps (I-Corps)) and foundation(The Wallace H. Coulter Foundation) funding programs seek toprevent lapses in funding by supporting promising new technologies and encouraging the creation of start-ups.

7. Future Work
Recent developments in drug delivery systems highlight theongoing technical challenges driven by the need to achieve
greater functionality and usability by both patients and caregivers.
MEMS-enabled systems allow for miniaturization, as well as integration of multiple functionalities leading to greater efficacy andperformance (including precision, automation, and personalization) while featuring less invasive and painful administration and

Figure 28. The Prometra Programmable Pump and catheter. Reprinted
with permission.[105] Copyright 2009, John Wiley and Sons.

© 2015 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim

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Received: December 3, 2014
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REVIEWfewer side effects. To achieve optimized and personalized patienttailored therapy, drug delivery systems of the future will likely
combine monitoring and therapy in closed loop systems that are
informed by the patient’s needs and can appropriately respond tothem to reach the desired therapeutic effect.
Digital and wireless health solutions can play an importantrole in achieving closed-loop therapy by enabling programmability of drug delivery systems to monitor device status and performance and send commands to adjust the system’s operation,
as well as data transmission between the system and an internet
based network. Remote monitoring of clinical events and symptoms reduces the frequency of routine follow-up visits. This inturn reduces staff time and costs while improving the patient’s
quality of life.[114]Another emerging trend in drug delivery system development is the emphasis on affordable technologies. Driven by costpressures from health economic considerations for reimbursement, there is a strong demand for systems that are external or
require only minimally invasive surgical procedures and offer
shorter and less costly patient recovery. With the increasingly
difficult US market environment for medical device innovation,
there is now increased attention on commercial development ofsystems intended for global markets where pricing pressures
dictate technology adoption and market penetration.[115]Despite challenges in development, funding, regulatory pathways, and market penetration, continued interest in the development of MEMS-based drug delivery systems is expected, driven
by the increasing demand in the ever increasing age of the world
population and emphasis on personalized medicine that can be
enabled by responsive, closed-loop therapeutic devices.[106]13

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© 2015 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim

Adv. Healthcare Mater. 2015,
DOI: 10.1002/adhm.201400772