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A reﬁllable microfabricated drug delivery device for treatment of oculardiseases†
Ronalee Lo,a Po-Ying Li,b Saloomeh Saati,c Rajat Agrawal,c Mark S. Humayuna,c and Ellis Meng*aReceived 19th March 2008, Accepted 19th May 2008
First published as an Advance Article on the web 29th May 2008
DOI: 10.1039/b804690e

An implantable manually-actuated microfabricated drugdelivery device was demonstrated as a new approach fordelivering therapeutic compounds to ocular tissue in acutein vitro, ex vivo, and in vivo studies.
Delivering therapeutic compounds into the body in a precisely
controlled and targeted manner is an ongoing challenge. Ocular drug delivery is necessary to slow disease progression
but is especially difﬁcult due to physiological barriers, space
limitations in and surrounding the eye, and trauma to theeye resulting from invasive therapies. If left untreated, chronic
retinal diseases such as glaucoma, age-related macular degeneration, diabetic retinopathy, and retinitis pigmentosa canlead to irreversible blindness.1 Currently available treatment
methods include topically and orally administered medications,intraocular injections, surgical intervention, and biodegradable
implants. However, these treatments are limited by minimal drug
diffusion to the target tissues, tissue trauma, and side effects.
We present the ﬁrst microfabricated, all-polymer oculardrug delivery device.2 This prototype features a reﬁllable drug
reservoir from which targeted delivery is provided to intraocular
tissues via a valved ﬂexible cannula (Fig. 1). The reﬁllable reservoir offers key advantages over existing ocular drug delivery systems: (1) only a single surgical intervention is required; (2) functional lifetime of the device is extended without increasing theoverall size; (3) in contrast to sustained delivery counterparts
which need to be removed and replaced when the infused drug
has been depleted, the reﬁllable drug delivery device can bereplenished without signiﬁcant surgical intervention; and (4) thedrug regimen can be altered on an as needed basis, including
replacing the drug with new, advanced pharmaceutical solutions.
The drug delivery device was formed by assembling threelayers of molded polydimethylsiloxane, or PDMS, (Sylgard184, Dow Corning, Midland, MI, USA). Assembled devices
consisted of a reﬁllable reservoir, cannula, normally-closed
check valve, support posts, and suture tabs (see electronic
supplementary information, ESI, for device dimensions).† The

aDepartment of Biomedical Engineering, Viterbi School of Engineering,
University of Southern California, *1042 Downey Way- DRB-140, Los
Angeles, CA, USA. E-mail: ellis.meng@usc.edu; Fax: +1-213-821-3897;
Tel: +1-213-740-6952bMing Hsieh Department of Electrical Engineering, Viterbi School ofEngineering, University of Southern California, Los Angeles, CA, USAc
Department of Ophthalmology, Doheny Eye Institute at the Keck Schoolof Medicine, University of Southern California, Los Angeles, CA, USA† Electronic supplementary information (ESI) available: Description ofdevice fabrication and delivery. See DOI: 10.1039/b804690e

This journal is © The Royal Society of Chemistry 2008

Fig. 1 (A) Device illustration and (B) assembled device photo. The
main components (reﬁllable reservoir, cannula, normally-closed checkvalve, support posts, and suture tabs) are indicated.

drug reservoir was placed beneath the conjuntiva and sutured tothe sclera between the rectus muscles. The transscleral cannula
was directed intraocularly via a surgical tunnel in the sclera(3 mm posterior to the limbus) and the cannula tip was placed
into either the anterior or posterior segment of the eye in close
proximity to the target treatment site associated with the disease
being managed (Fig. 2). Sutures were used to close and seal thescleral tunnel opening around the cannula.

Fig. 2 Illustration of device placement on the eye; temporal placement
provides easier access for reﬁll and improves patient comfort. Here,the anterior segment conﬁguration is shown (i.e. glaucoma treatment).
Image modiﬁed using image courtesy of National Eye Institute, National
Institutes of Health.

Drug solution was dispensed through the valved cannula
following manual depression of the reservoir; the internal
reservoir pressure must increase beyond the check valve cracking
pressure to permit drug ﬂow through the valve oriﬁce. Flowceased after the driving pressure was removed and the valve
closed preventing bodily ﬂuids from backﬂowing. Support posts
were positioned along the length of the cannula and arrayed in
the reservoir as a precaution to prevent the top and bottom walls
from collapsing due to stiction between these surfaces.3
The PDMS components were molded using a soft-lithography
process with silicon and acrylic masters. 4 silicon wafers were
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etched to create masters for the bottom and middle layer moldsthat form the base layers of the drug delivery platform. The
top layer, or reservoir cap, was formed from a conventionallymachined acrylic master. An illustration of the fabrication
process can be found in the ESI.†Silicon masters were fabricated using deep reactive ion etching(PlasmaTherm SLR-770B, Unaxis Corporation, St. Petersburg,FL, USA) to create 100 lm and 250 lm deep features in thebottom and middle molds, respectively. PDMS mixed in a 10 : 1
base to curing agent ratio was poured into the silicon masters,degassed under vacuum, and cured (1 h at 70 ◦ C). Molded
PDMS sheets were carefully removed from the master andindividual replicas were separated using a ﬁne-tipped blade. Thecheck valve was formed by cutting a hole, 305 lm in diameter,in the middle layer at the end of the cannula. This hole wasaligned over the last support post of the bottom layer to createa normally-closed check valve.
The acrylic master for the top layer was made by securing
Plexiglas squares onto a glass slide using epoxy (Loctite
Professional Epoxy, Henkel Consumer Adhesives, Inc., Avon,OH, USA). The interior volume of the reservoir was deﬁned
by conventionally-machined Plexiglas squares; a target internal
volume of 50 lL was obtained using 6 mm × 6 mm squares cut
from a 1.59 mm thick Plexiglas sheet (actual volume 57.15 lL).
PDMS was poured over the acrylic masters, degassed, andpartially-cured for 30 min at 70 ◦ C. Reservoirs were individually
separated from the molded PDMS sheet.
The bottom and middle PDMS layers were aligned andirreversibly bonded by oxygen plasma surface modiﬁcation. The
partially-cured top layer was then assembled and the entire
device was allowed to fully cure at room temperature (12 to24 h). A layer of uncured PDMS was applied to the device
perimeter to enhance bond strength between device layers andproduce a mechanically robust structure able to survive surgicalhandling.4,5 This reinforcing layer was cured (1 h at 70 ◦ C) andexcess PDMS was carefully removed.
The ocular drug delivery device was ﬁlled and reﬁlled bypiercing the reservoir membrane with a syringe needle. The
device lifetime is dependent on the ability to repeatedly reﬁll
the device and on the integrity of the punctured membrane. A
smaller needle diameter prolongs the lifetime of the reservoir;
however, the needle must be stiff enough to puncture both theconjunctiva covering the device and reservoir membrane withease. Given these tradeoffs, 30 G needles (305 lm outer diameter)
were selected and two needle tip types (coring and non-coring)were investigated.
PDMS membranes were punctured in vitro with both needle
types. The needle insertion points as well as the needle track

cross-sections were compared. SEM images of both needle tips
were taken to show the needle tip shape (Fig. 3A). The noncoring needle displaced material as it penetrated the membrane,
whereas the coring needle removed a cylindrical plug from themembrane. After needle removal, the displaced material aroundthe non-coring needle was restored to its original location,effectively re-sealing the tear, whereas a hole was left by thecoring needle (Fig. 3B,C). Thus, the 30 G non-coring needle wasselected as the reﬁll needle.

Fig. 3 Comparison of 30 G coring and non-coring needles. (A) SEM
images of the needle tip, (B) top view of the needle puncture site in a
PDMS slab and (C) cross sectional view of the needle track.

The puncture location on the reﬁllable reservoir must beable to withstand a pressure gradient without leaking. Average
intraocular pressure (IOP) is 15.5 ± 2.6 mmHg (mean ± SD),
however, glaucoma patients can have elevated IOPs (typically
21 to 35 mmHg, higher values are possible due to temporal
ﬂuctuations or inﬂammation, such as glaucomacyclitic crisis).6,7
The mechanical integrity of punctured PDMS was characterizedby piercing PDMS membranes (250 lm and 670 lm thick)with a 30 G non-coring needle 8, 12, and 24 times. The
punctured membrane was placed in a custom jig with pressurizedwater applied to one side of the membrane and a 50 lL
calibrated pipette (Clay Adams, Parsippany, NJ, USA) attachedto the other side. Pressure was increased in 0.1 psi (0.69 kPa,5.17 mmHg) increments every 10 min until water leakage wasobserved by monitoring the pipette. A wide range of testingpressures well above normal and abnormal IOP values wereapplied.
The membranes were punctured repeatedly in the same location and evaluated. Puncture events through the same site represented the worst-case scenario where each subsequent puncture
can cause further damage to the puncture location on themembrane. A summary of leakage pressures for both membrane

Table 1 Summary of leakage pressures for PDMS membranes punctured with a 30 G non-coring needle (mean ± SE, N = 4)
Number of punctures

250 lm membrane leakage pressure

670 lm membrane leakage pressure

8

1.01 ± 0.1 psi(6.96 ± 0.7 kPa, 52.23 ± 5.2 mmHg)
0.93 ± 0.2 psi(6.41 ± 1.4 kPa, 48.09 ± 10.3 mmHg)
0.64 ± 0.1 psi(4.41 ± 0.7 kPa, 33.10 ± 5.2 mmHg)

15.40 ± 1.5 psi(106.18 ± 10.3 kPa, 796.41 ± 77.6 mmHg)
4.24 ± 0.6 psi(29.23 ± 4.1 kPa, 219.27 ± 31.0 mmHg)
4.45 ± 1.4 psi(30.68 ± 9.7 kPa, 230.13 ± 72.4 mmHg)

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This journal is © The Royal Society of Chemistry 2008

thicknesses is presented in Table 1. The leakage pressure in both
membranes decreased with additional punctures. The leakagepressure associated with the thicker membrane leveled out after
12 punctures, suggesting subsequent punctures did not induce
further damage. As expected, higher pressures were necessary
to induce leakage across the thicker membranes, however, theincrease in leakage pressure was not linearly proportional tothe increase in membrane thickness, suggesting a substantial
improvement in mechanical integrity with additional thickness.
In all but one setup, punctured membranes had leakage pressures higher than normal or abnormal IOPs. This experiment
was conducted under laboratory conditions and the results may
differ when the material is exposed to biological conditions.
However, these results provide a baseline for estimating thenecessary reﬁll membrane dimensions to withstand IOP ranges.
The ability to reﬁll the device signiﬁcantly increases the device
lifetime. Clinicians have suggested a practical reﬁll frequency of
no more than once every 3 months, with an associated 50 to500 nL daily dosage volume to accommodate a wide range ofpossible drug regimens. A 50 lL reservoir can deliver 1000 dosesof 50 nL or 100 doses of 500 nL, resulting in a reﬁll rate of once
every 33 or 3 months, respectively. A 3 month reﬁll frequency
with a maximum of 24 possible reﬁlls results in a device lifetimeof 6 years.
In vitro, ex vivo and in vivo studies were conducted to demonstrate the feasibility of drug delivery into the eye. All animal
procedures were approved by the Institutional Animal Careand Use Committee at the University of Southern California,
Los Angeles and performed in conformance with the Guide forCare and Use of Laboratory Animals (National Institutes ofHealth).
Reﬁll and device function were ﬁrst veriﬁed in vitro. Then,
enucleated porcine eyes were used to develop the appropriate
surgical procedure for securing the device to the eye andplacement of the cannula into the anterior chamber. Following
the surgical procedure, device functionality was demonstrated
ex vivo through repeated dispensation trials of dyed water
followed by reservoir reﬁll via a 30 G non-coring needle.
Following several successful ex vivo trials, in vivo directed
delivery of dyed water into the anterior chamber of male Dutch
Belted pigmented rabbit eyes was conﬁrmed.
Acute drug delivery was investigated using 10% phenylephrine
solution; phenylephrine delivery enables real-time physiologicalveriﬁcation of drug delivery by inducing rapid pupil dilation.
Surgical sham devices were also fabricated and tested to aid in
rapid prototyping of the device to incoporate improvements indevice size, shape, or component placement. Identical 22.5 lL

volumes were delivered by conventional intraocular injection(30 G needle and control eye) and our drug delivery device, andthen pupil dilation for each case was compared.
The drug delivery device and surgical sham were secured to thesclera and the cannula was directed into the anterior chamber
through a limbal incision or scleral tunnel, respectively (seeESI for description of surgical technique).† Baseline pupillary
diameters (vertical and horizontal) measurements were taken
prior to device dispensation. The phenylephrine was delivered
into the eye by depressing the reservoir with surgical forceps
or cotton swabs (see ESI for description of the experimental
setup).† After dispensation, pupil diameter measurements weretaken and compared to baseline values. A summary of resulting
pupillary diameters is found in Table 2.
The control eye exhibited a greater pupillary diameter change
compared to the other delivery mechanisms even when asmaller drug volume was used. Minimal leaking occurred during
intraocular injection; it is well known that punctures <1 mmdiameter are self-sealed. For both the device and the surgicalsham, drug contact with the iris was diminshed due to theexistence of leakage paths around the cannula. Sutures areunable to completely seal the limbal incision around the cannulain the device. A scleral tunnel reduces leakage around thecannula, however, in the unvalved surgical sham, drugs canbackﬂow into the sham.
This manually-controlled drug delivery device enables targeted delivery of pharmaceuticals to intraocular sites. Here,successful delivery of phenylephrine to the anterior segment wasdemonstrated; anterior segment delivery is of particular interestin the management of glaucoma. However, precise control ofdelivered drug volume in this delivery platform is limited byvariations in the duration and force applied during manual
dispensation. The next generation device will utilize the samedelivery paradigm but replace manual actuation with electrolysis
actuation to precisely pump the desired dosage volume.8–10
Automated delivery at the expense of a more complicated device
may be preferred to further improve patient compliance.
In summary, a prototype, manually-actuated, reﬁllable oculardrug delivery device capable of targeted delivery was designed,fabricated, and tested. Device functional lifetime was extended
compared to conventional dissolving ocular drug delivery implants by incorporating a reﬁllable drug reservoir accessible witha 30 G non-coring needle. The device feasibility and functionality was demonstrated in vitro, ex vivo, and in vivo. Real-time physiological evidence of drug delivery was obtained in acute studiesof phenylephrine delivery. Electronically-controlled pumping for
precise dosage control will be integrated into the next generation

Table 2 Summary of in vivo phenylephrine delivery studies in rabbits including vertical and horizontal pupillary diameter changes

VolumeDelivery mode
Baseline
Following delivery
Total change

Control (syringe)

Drug delivery device

Surgical sham22.5 lL
Injected through limbus
VerticalHorizontal
5 mm
4 mm
9.5 mm
9 mm
4.5 mm (90%)
5 mm (125%)∼22.5 lL
Cannula (limbal incision)
VerticalHorizontal
6.5 mm
6 mm
8 mm
7 mm
1.5 mm (23%)
1 mm (16%)∼70 lL
Cannula (scleral tunnel)
VerticalHorizontal
4.5 mm
4.5 mm
7.5 mm
7.5 mm
3 mm (66%)
3 mm (66%)

This journal is © The Royal Society of Chemistry 2008

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device for chronic treatment of glaucoma and other chronicocular diseases. Chronic in vivo studies are being conducted todetermine long-term biocompatibility and functionality, as well
as determining any biofouling that may occur during long-term
implantation. These microfabricated drug delivery devices offer
many advantages over existing ocular drug delivery therapies
and enable a new mode of drug delivery not possible withconventional therapies.

Acknowledgements
This work was funded in part by the Engineering ResearchCenters Program of the NSF (EEC-0310723) and Bausch andLomb. The authors would like to thank Mr Trevor Roper,
Mr Merrill Roragen, Dr Yu-Chong Tai, Mr Tun Min Soe, andthe members of the USC Biomedical Microsystems Laboratory
for their assistance with this project.

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Notes and references
1 D. H. Geroski and H. F. Edelhauser, Invest. Ophthalmol. Visual Sci.,
2000, 41, 961–964.
2 R. Lo, K. Kuwahara, P. Y. Li, R. Agrawal, M. S. Humayun and E.
Meng, MMB 2006, Okinawa, Japan, 2006.
3 T. J. Yao, X. Yang and Y. C. Tai, Sens. Actuators, A, 2002, 97–8,771–775.
4 M. A. Unger, H. P. Chou, T. Thorsen, A. Scherer and S. R. Quake,Science, 2000, 288, 113–116.
5 H. K. Wu, B. Huang and R. N. Zare, Lab Chip, 2005, 5, 1393–1398.
6 Glaucoma: Science and Practice, Thieme Medical Publishing, NewYork, 2003.
7 R. Ritch, M. B. Shields and T. Krupin, The Glaucomas, Mosby,
St. Louis, 1989.
8 P.-Y. Li, J. Shih, R. Lo, B. Adams, R. Agrawal, S. Saati, M. S.
Humayun, Y.-C. Tai and E. Meng, MEMS 2007 Proceedings, Kobe,Japan, 2007.
9 P.-Y. Li, J. Shih, R. Lo, R. Agrawal, S. Saati, M. S. Humayun, Y.-C.
Tai and E. Meng, Sens. Actuators, A, 2008, 43, 41–48.
10 E. Meng, J. Shih, P.-Y. Li, R. Lo, M. S. Humayun and Y.-C. Tai,lTAS 2006, Tokyo, Japan, 2006.

This journal is © The Royal Society of Chemistry 2008