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Reversible thermoresponsive materials for temporary closure of ocular trauma
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Reversible thermoresponsive materials for temporary closure of ocular trauma
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Content
REVERSIBLE THERMORESPONSIVE MATERIALS FOR TEMPORARY CLOSURE OF
OCULAR TRAUMA
by
Niki Bayat
A Dissertation Presented to the
FACULTY OF THE GRADUATE SCHOOL
UNIVERSITY OF SOUTHERN CALIFORNIA
In Partial Fulfilment of the
Requirements for the Degree
DOCTOR OF PHILOSOPHY
(CHEMICAL ENGINEERING)
August 2018
Copyright 2018 Niki Bayat
ii
ACKNOWLEDGMENTS
There are so many people to thank for helping me during my doctoral work. This thesis would not
have been possible without the support of many individuals, to only some of whom it is possible
to give particular mention here.
Above all, I would like to thank my parents who have always been a source of
encouragement and inspiration to me throughout my life. Thank you so much for the trust and
confidence you always had in me. Thank you for believing in me and for always being there for
me. Mom and dad, I cannot thank you enough for giving me endless support and confidence to
make this happen. You are the best parents in the world and I owe my success to you.
Next, I would like to express the deepest appreciation to my mentor, dissertation advisor,
Professor Mark E. Thompson for his valuable guidance, encouragement and persistent help in
different aspects of my research. It was an honor and pleasure to work under his supervision and I
have learned a lot from him. Thank you so much for encouraging me to look at my research and
my work in different ways. Your support was imperative to my completion of this degree and
essential to my success.
In addition, I would like to sincerely thank Prof. Katherine Shing and Prof. Barry
Thompson for serving on my committee and providing indispensable advice, information and
guidance over the last year and a half.
I would especially like to thank, Prof. Mark S. Humayun and Dr. Jack J. Whalen, our
wonderful collaborators at University of Southern California Roski Eye Institute, who provided
clinical and biomedical engineering guidance on the overall strategy for this project. I am most
iii
grateful to Dr. Yi Zhang and Dr. Paulo Falabella for helping me with the in vivo study design,
setup, and execution. Many of my experimental work would have not have been completed without
their assistance.
The persistent assistance, great patience and cooperation of my amazing fellow graduate
student, Roby Menefee, was necessary for the success of this project. His work was essential in
providing me with the data that I needed to choose the right direction and successfully complete
my doctoral dissertation. I am extremely lucky and profoundly thankful for his presence and
support throughout my project.
To my dear friends, Becky, Manely, Qiwen, Narcisse, Piyumie, Rasha, Hossein, Manuel,
Nafiseh, Nasim, John, Alireza, Majid, Ehsan, Andrew, Bita, Keyvan, Nazzy, Mehdi and Judy, who
have stuck by me through so many ups and downs. I want you to know that I will always be there
for you and I look forward to sharing many more of life’s lovely moments with you.
I would like to thank my lovely brothers, Amirmasoud and Shahin, for their unconditional support
and love. I want to thank you both for taking care of our parents when I was not here. You are
absolutely amazing and wonderful.
Last but certainly not least, I am so thankful to my wonderful and genius husband, Amjad.
Without your patience, constant encouragement and understanding, it would have not been
possible for me to achieve so much in such a short time. Thank you for loving me despite my
shortcomings and my imperfections. I just want you to know how thankful I am that I have a
husband as amazing as you.
iv
Table of Contents
List of Figures ................................................................................................................................ vi
List of Tables ................................................................................................................................ xii
CHAPTER 1 ................................................................................................................................... 1
1.1 Introduction ........................................................................................................................ 1
1.2 Study design ....................................................................................................................... 7
CHAPTER 2 ................................................................................................................................... 9
2.1 Engineered injectable TRS and mode of function .............................................................. 9
2.2 Synthesis and characterization of the N85NT15 and N95BA5 copolymers ......................... 10
2.3 Optimizing rheological properties of shape-persistent, moldable TRS ........................... 16
2.4 Physical characterizations of the TRS .............................................................................. 22
2.5 Materials and methods ...................................................................................................... 27
CHAPTER 3 ................................................................................................................................. 31
3.1 Ex vivo testing of N95BA5 for ocular trauma ................................................................... 31
3.2 In vitro adhesion testing on TRS hydrogel ....................................................................... 34
3.3 In situ gelation mechanism of TRS in the eye .................................................................. 35
3.4 Thermo-induced volume change of molded TRS ............................................................ 40
3.5 Materials and methods ...................................................................................................... 43
CHAPTER 4 ................................................................................................................................. 45
4.1 Design of a custom injection tool ..................................................................................... 45
4.2 Effective TRS deployment in an in vivo model of ocular trauma .................................... 49
4.3 N95BA5 biocompatibility up to 1 month after ocular implantation .................................. 56
4.4 Materials and methods ...................................................................................................... 62
v
CHAPTER 5 ................................................................................................................................. 67
5.1 User feedback workshop with ophthalmologists and technicians .................................... 67
5.2 Clinical user workshops to collect design feedback ......................................................... 70
CHAPTER 6 ................................................................................................................................. 73
6.1 HERA design for ocular trauma ....................................................................................... 73
6.2 Body temperature activated HERA with superior mechanical properties ....................... 78
6.3 The nanostructure of crosslinked HERA network at body temperature .......................... 89
6.4 Materials and methods .................................................................................................... 104
CHAPTER 7 ............................................................................................................................... 106
7.1 In vitro adhesion function of HERA hydrogels .............................................................. 106
7.2 Ex vivo testing of HERA as an ocular adhesive ............................................................ 118
7.3 Materials and methods .................................................................................................... 121
REFERENCES AND NOTES .................................................................................................... 124
vi
List of Figures
Fig. 2.1: Design of a thermo-responsive hydrogel to seal scleral perforation. ............................... 9
Fig. 2.2: The sealing hydrogel transitions from hydrophilic coils to hydrophobic globules at its
lower critical solution temperature (LCST), markedly changing physical properties. ................. 10
Fig. 2.3: Schematic synthesis routes of the N85NT15 and N95BA5 copolymers. (A) poly(NIPAM-
co-N-tert-butylacrylamide), (B) poly(NIPAM-co-butylacrylate). ................................................ 12
Fig. 2.4: Molecular structures of formulations of poly(N-isopropylacrylamide) (PNIPAM) with
butylacrylate (BA) or N-tert-butylacrylamide (NT) and a table of resulting alterations of
molecular properties and LCST values. ........................................................................................ 13
Fig. 2.5:
1
H NMR spectrum for N85NT15 and N95BA5 copolymers in CDCl3. ............................. 14
Fig. 2.6: Scattering intensity spectra of N95BA5 as a function of temperature. ............................ 15
Fig. 2.7: Normalized scattering intensity as a function of temperature for PNIPAM,
poly(NIPAM-co-NT) (N85NT15), and poly(NIPAM-co-BA) (N95BA5). A.U., arbitrary units. .... 16
Fig. 2.8: Strain amplitude for N95BA5 and N85NT15 at T = 6°C. .................................................. 17
Fig. 2.9: Strain amplitude of N85NT15 at its LCST. ...................................................................... 18
Fig. 2.10: Storage and loss moduli (G′ and G″, respectively) over strain for N95BA5. ................ 19
Fig. 2.11: G″ (B) and G′ (C) modulus representations of viscoelastic behavior for co-BA and co-
NT as a function of angular frequency at fixed strain. ................................................................. 19
Fig. 2.12: Table of storage moduli for co-NT and co-BA compositions at different angular
frequencies. Storage and loss moduli (G′ and G″, respectively) over frequency for co-NT and co-
BA compositions. .......................................................................................................................... 20
Fig. 2.13: Measurement of complex viscosity for N95BA5 according to temperature. ................. 21
Fig. 2.14: Complex viscosity of N95BA5 as a function of temperature and concentration. .......... 22
Fig. 2.15: Compressive stress-strain characterization and compressive modulus of the N95BA5
hydrogel. ....................................................................................................................................... 23
Fig. 2.16. Tensile stress-strain characterization and tensile modulus of the N95BA5 hydrogel. ... 25
have elastic modulus of 117 kPa, representing improved stretching capacity compared to 60 kPa
(25%) and 45 kPa (20%). .............................................................................................................. 26
vii
Fig. 2.17: (A to C) Images of square and round gel molds formed by heating the hydrogel
solutions to 32°C. Scale bars, 1 cm. (D to F) Images of solid N95BA5 hydrogel demonstrating
horizontal resilience (D), vertical resilience (E), and strength of form on contact (F). Scale bars,
1 cm. .............................................................................................................................................. 26
Fig. 3.1: Schematic (top) and image (bottom) depicting ex vivo procedures carried out in a
pressure-controlled explanted cadaveric pig eye. ......................................................................... 31
Fig. 3.2: Images (top) and schematic depiction (bottom) of hydrogel injection through scleral
perforation by deployment of a sealant trail through the wound, leaving rivet-like caps
subsequently removed to leave the occlusion flush with the scleral surface. ............................... 32
Fig. 3.3: Comparison of maintained intraocular pressures across a concentration spectrum for
N95BA5 and N85NT15 hydrogels (n = 3 per group) (left). Image of a solid plug removed from a
test eye (right). .............................................................................................................................. 33
Fig. 3.4: Schematic depicting tissue adhesion tests comparing PNIPAM, N95BA5, and
cyanoacrylate adhesion strength to scleral tissue ex vivo. ............................................................ 34
Fig. 3.5: Adhesion force of different concentrations of N95BA5, PNIPAM, and cyanoacrylate to
scleral tissue. ................................................................................................................................. 35
Fig. 3.6: N95BA5 particle size as assessed using dynamic light scattering in solution below LCST
(2°C), showing a 96% scattering intensity for small-radius particles. .......................................... 36
Fig. 3.7: Particle size in the phase transition region (12°C). A split was observed between
particles with hydrodynamic radius of 4.5 and 236.2 nm. ............................................................ 37
Fig. 3.8: Particle size toward the end of the phase transition region (18°C). Ninety-eight percent
of scattering intensity was due to large-radius N95BA5 aggregates. ............................................. 38
Fig. 3.9: Intensity distribution graph of DLS spectra for N95BA5 recorded at different
temperatures. ................................................................................................................................. 39
Fig. 3.10: Intensity distribution graph for N95BA5 as a function of temperature.......................... 39
Fig. 3.11: Hydrodynamic radius size of particles traced through the hydrogel transition to show
higher aggregate populations at higher temperatures (n = 3 per temperature). ............................ 40
Fig. 3.12: Gross images of molded hydrogel samples held at the expected eye temperature
(32°C) for up to 30 days, showing slight volume decrease and good shape persistence and
stability. ......................................................................................................................................... 41
viii
Fig. 3.13: Hydrophobic/hydrophilic nature of the N95BA5 hydrogel above and below the LCST.
....................................................................................................................................................... 42
Fig. 4.1: Design diagrams of a custom injection tool to effectively control hydrogel deployment
and regulate its temperature. ......................................................................................................... 46
Fig. 4.2: Validation of a custom injection tool to effectively control hydrogel deployment and
regulate its temperature. ................................................................................................................ 47
Fig. 4.3: An image of the prototype injector. ................................................................................ 47
Fig. 4.4: Injector tool cooling reaction calibration curves for 2.5 and 12.5 g of ammonium nitrate
to various volumes of added water. 2.5 g (left) and 12.5 g (right) ammonium nitrate. ................ 48
Fig. 4.5: Two-arm study design to assess safety and efficacy of the hydrogel versus the current
standard of care for posterior segment open globe injuries. ......................................................... 49
Fig. 4.6: Images of the surgical procedure in rabbits. ................................................................... 51
Fig. 4.7: Visual evaluation of eye responses to surgical procedure and sealant placement. ........ 52
Fig. 4.8: Representative baseline intraocular pressure (IOP) values showing no statistical
difference between eyes of the same animal or any circadian-induced variations; columns show
6.5 ± 0.2, 6.5 ± 0.2, 5.8 ± 0.2, 6.4 ± 0.3, 8.1 ± 0.4, 8.0 ± 0.4, 7.4 ± 0.5, and 8.6 ± 0.4 mmHg (n =
8 per group). .................................................................................................................................. 53
Fig. 4.9: Wald test comparison of mean IOP values of the treatment group versus no
intervention, after procedure, showed a statistically significant improvement in mean IOP with
sealant placed (*P < 0.05 and **P < 0.001). ................................................................................. 55
Fig. 4.10: Series of histological cross sections prepared for control (left pairs) and treatment
(right pairs) in hematoxylin and eosin (H&E) and Masson’s trichrome stain for each of the study
endpoints (t = 48 hours, 1 week, and 4 week). Scale bars, 800 mm. ............................................ 57
Fig. 4.11: Increased magnification of one of the laceration margins for the treatment group
showing evolution of the tissue-hydrogel interface from acute inflammatory infiltrate to a
mature, compact fibrotic encapsulation layer at 4 weeks. Scale bars, 50 mm. ............................. 59
Fig. 4.12: Gross visualization (top row) and high-magnification (bottom row) evaluation of
treatment group retinas showing no evidence of trauma induced retinal detachment or hydrogel-
induced retinal neurotoxicity. Scale bars, 5 mm (top row) and 100 mm (bottom row). ............... 60
Fig. 5.1: TRS in the hands of professionals: In vitro application at Walter Reed Medical Center.
....................................................................................................................................................... 68
ix
Fig. 6.1: Engineered HERA for ocular trauma and mode of function. HERA application and
temperature-mediated removal from a scleral perforation. ........................................................... 74
Fig. 6.2: Images of adhesive hydrogel temperature transition from translucent gel to opaque
aggregate when crossing lower critical solution temperature. ...................................................... 75
Fig. 6.3: Phase diagram of hydrogel phase behavior demonstrating the role of polymer
composition in the LCST transition between hydrophilic chains and less soluble aggregates. ... 76
Fig. 6.4: Molecular structures of the Arg-Gly-Asp-Ser (RGDS) peptide and 3-
guanidinopropionic acid (GPA) additives selected for similarity of functional groups (red) and
bioactive adhesive potential. ......................................................................................................... 78
Fig. 6.5: Strain-amplitude sweep of GPA-enhanced hydrogel [0.58% (w/w)] performed at fixed
temperature (32°C) isolated the hydrogel linear region for further analysis and demonstrated a
large capacity for deformation. ..................................................................................................... 79
Fig. 6.6: Storage and loss moduli (G′ and G″, respectively) of GPA hydrogel [0.58% (w/w)] were
compared at temperatures spanning the transition region (10°, 20°, 30°C), showing a significant
increase in storage modulus and solid-like behavior at higher temperatures (n = 3). .................. 80
Fig. 6.7: Loss & storage modulus of unmodified TRS as a function of temperature. .................. 80
Fig. 6.8: Loss & storage modulus of RGDS-included hydrogel as a function of temperature. .... 81
Fig. 6.9: Results of fixed-temperature (16°C) oscillatory rheometry confirmed an increase in loss
modulus (G″) and an overall decrease in tan, illustrating improved overall material durability and
a relative increase in solid-like behavior. ..................................................................................... 83
Fig. 6.10: Storage modulus of additive-included hydrogels at 16°C (n = 30). ............................. 83
Fig. 6.11: Fixed-temperature oscillatory rheometry above LCST (32°C) demonstrated dramatic
increases in elastic moduli (G′) of additive-enhanced samples [2.91% (w/w)] with GPA dwarfing
other samples. ............................................................................................................................... 85
Fig. 6. 12: Loss modulus of TRS as a function of GPA and RGDS at 32°C (n = 30). ................. 85
Fig. 6.13: Oscillatory temperature sweep of additive samples followed the same complex
viscosity trend as pure TRS, but exhibited steeper increase with a higher apex even with low
concentration [0.58% GPA (w/w)]. .............................................................................................. 86
Fig. 6.14: Below LCST, complex viscosity values were comparable across all studied samples,
but indicating an additive mediated increase (n = 3). ................................................................... 88
x
Fig. 6.15: Above LCST, complex viscosity values were dramatically elevated for GPA-enhanced
formulations. ................................................................................................................................. 88
Fig. 6.16: Dynamic light scattering comparison of GPA-enhanced [2.91% (w/w)] N95BA5
hydrogel to a sample without additive indicated two particle size populations below LCST, with
an increase in hydrodynamic radius of large particles for the additive-enhanced hydrogel. ........ 90
Fig. 6.17: Temperature increase above transition indicated the transition to one dominant particle
size and a preserved elevation of GPA particle sizes. .................................................................. 91
Fig. 6.18: Comparison of GPA concentration effects on particle size revealed similar particle
sizes below LCST, but a concentration-dependent increase in particle size at higher temperatures
(n = 3). ........................................................................................................................................... 92
Fig. 6.19: Normalized average particle size for GPA hydrogels across temperature transition. .. 94
Fig. 6.20: Comparison of RGDS [2.91% (w/w)] particle dynamics to an unmodified N95BA5
showed similar improved particle size below the transition temperature. .................................... 95
Fig. 6.21: Comparison of particle distributions showed a striking increase in particle size of
GPA-enhanced samples relative to the RGDS-included formulations. ........................................ 96
Fig. 6.22: DLS spectra of the intensity distribution graph of copolymer solution with RGDS at
18°C. ............................................................................................................................................. 97
Fig. 6.23: Concentration-dependent of RGDS effect on particle size distribution within dilute
hydrogel. ....................................................................................................................................... 98
Fig. 6.24: RGDS peptide caused a significant increase in average particle size above LCST (n =
3). ................................................................................................................................................ 100
Fig. 6.25: Normalized average particle size for RGDS-enhanced hydrogels across various RGDS
concentrations as a function of temperature. .............................................................................. 101
Fig. 6.26: Temperature dependence of hydrodynamic radii in additive (GPA/RGDS)-enhanced
copolymer solution...................................................................................................................... 101
Fig. 6.27: Additive concentration independent transition temperature value revealed by
normalized particle size. ............................................................................................................. 103
Fig. 6.28: Comparison of average particle size at body temperature (32°C) demonstrated larger
polymer networks in GPA [2.91% (w/w)] hydrogels compared to RGDS samples (n = 3). ...... 103
Fig. 7.1: Schematic of uniaxial adhesion tests as performed on a variety of biological and non-
biological substrates. ................................................................................................................... 107
xi
Fig. 7.2: Hydrogel adhesion forces on scleral squares with negative (tissue-tissue) and positive
(cyanoacrylate) controls. ............................................................................................................. 108
Fig. 7.3: Contact angle of DI water on surface-modified glass slides (n = 5). ........................... 109
Fig. 7.4: Contact angle of HCl solution (pH = 0) on surface-modified glass slides (n = 5). ...... 110
Fig. 7.5: Contact angle of NaOH solution (pH = 14) on surface-modified glass slides (n = 5). 111
Fig. 7.6: Demonstration of adhesion test raw data with similar maximal force for bond breakage
[GPA sample].............................................................................................................................. 112
Fig. 7.7: Comparison of hydrogel adhesion strength on unmodified sandblasted glass squares. 112
Fig. 7.8: Comparison of hydrogel adhesion strength on hydrophobic (octyl) modified sandblasted
glass squares................................................................................................................................ 113
Fig. 7.9: Comparison of hydrogel adhesion strength on negatively charged (carboxylic acid)
modified sandblasted glass squares. ........................................................................................... 113
Fig. 7.10: Comparison of hydrogel adhesion strength on positively charged (amide) modified
sandblasted glass squares. ........................................................................................................... 114
Fig. 7.11: Comparison of adhesive and cohesive properties of the pure and modified TRS. ..... 115
Fig. 7.12: 0.1 mL hydrogel adhesion to 2.5 cm textured steel square (n = 5). ........................... 116
Fig. 7.13: Comparison of hydrogel adhesion on highly textured steel substrates. ..................... 116
Fig. 7.14: Schematic of intraocular pressure test using an ex vivo cadaveric pig eye ................ 119
Fig. 7.15: Ex vivo adhesive performance using pressure-controlled explanted cadaveric pig eye.
..................................................................................................................................................... 120
xii
List of Tables
Table 4.1: Injector tool design requirements. ............................................................................... 45
Table 4.2: In vivo study tabulated trajectory. ............................................................................... 50
Table 4.3: In vivo statistical analysis of average OD/OS for groups over time. .......................... 55
Table 4.4: Scleral tissue response at the hydrogel-sclera interface. ............................................. 57
Table 5.1: Responses from freehand write-in section of user survey administered during clinical
user workshop. .............................................................................................................................. 69
Table 5.2: Responses from multiple choice section of user survey administered during clinical
user workshop. .............................................................................................................................. 69
Table 6.1: Comparison of additive [2.91% (w/w)] effects on TRS complex viscosity by fixed-
temperature rheology across LCST showed similar values below the transition region, but
additive-mediated increase beyond LCST. ................................................................................... 87
Table 6.2: Temperature-dependent variation of the hydrodynamic radii of GPA-included N95BA5
solution [5% (w/v)]. ...................................................................................................................... 92
Table 6.3: Hydrodynamic radii of GPA-enhanced N95BA5 solutions demonstrates additive-
mediated aggregation. ................................................................................................................... 93
Table 6.4: At gelation point (12°C), additives increased the size of larger particles and the
amount in that larger population. .................................................................................................. 95
Table 6.5: Hydrodynamic radius and intensity distribution profiles of the TRS hydrogel in the
presence of RGDS and GPA [2.91% (w/w)] at 18˚C. .................................................................. 97
Table 6.6: Variation of hydrodynamic radii of the aggregates as a function of temperature for
N95BA5 solutions with different concentrations of RGDS. .......................................................... 99
Table 6.7: RGDS increases temperature-mediated aggregation in N95BA5 solutions [5% (w/v)].
..................................................................................................................................................... 100
1
CHAPTER 1
1.1 Introduction
At least 2.5 million eye injuries occur in the United States each year, and open globe injuries
account for 10% of these injuries (1, 2). Open globe injuries can quickly escalate in complexity
and yield poor visual outcomes if not managed carefully. Although incidence rates are relatively
low, virtually all open globe injury patients see a reduction in visual acuity (VA), and the
probability of VA loss increases with increasing time to intervention beyond 24 hours from the
injury (3). In addition to affecting the quality of life of the patient, lifetime health care costs
associated with visual impairment can approach $500,000, thus also having a major financial and
societal impact (4).
Combat- and mass casualty–related ocular traumas are subsets of ocular injuries in which
time to intervention is often delayed due to circumstances where patients are separated from
medical services or triaged behind other casualties with more critical injuries. This delay in
receiving treatment increases the risk for substantial visual impairment. In the U.S. campaigns in
the Middle East, up to13%ofall casualties presented eye injuries, most attributed to improvised
explosive devices (5), and between 20 and 40% of battlefield ocular injuries included penetrations
to the sclera (6). A study of the Boston Marathon bombing determined that 13% of those injured
required ophthalmology intervention (7). Military studies and clinical observations predict that
treatments closest to the time of injury have the best outcomes (8). In these instances, developing
a strategy to rapidly and temporarily close the globe without further trauma to the tissues is
desirable.
2
Current treatment options for managing open globe injuries include sutures and adhesives.
In general, these approaches require the use of microsurgical instrumentation accompanied by
surgical microscopes to visualize tissue repair. Foreign body sensation resulting from abrasive
material has been associated with eye rubbing, prolonged healing times, infection, and fibrosis (9).
Novel bioadhesives like fibrin matrices have been used (9, 10) but can carry considerable risks
associated with assurance prevention of viral or prion contamination, in addition to challenges with
glue deployment and ease of use. Cyanoacrylates (DERMABOND, TRUFILL, and DYMAX 222)
exhibit some difficulties in dispensing and cannot be used to close globes with missing tissues
(11). Acute inflammatory reactions in vascular tissue have been reported (12). There are currently
no U.S. Food and Drug Administration–approved indications for using medical adhesives for
closure of scleral penetrations.
Here, we propose a system designed to temporarily occlude open globe injuries. The
system leverages the reversible, thermoresponsive properties of poly(N-isopropylacrylamide)
(PNIPAM) to reversibly occlude injuries without causing additional trauma to surrounding tissues
during placement or removal. PNIPAM is a smart biostable polymer investigated for a range of
biomedical, drug screening, biotechnology, and medical diagnostics applications (13-16). Below
about 32° to 33°C [lower critical solution temperature (LCST)], hydrophilic interactions of
PNIPAM with water enable a translucent liquid state; above its LCST, it forms a partially
dehydrated, soft-solid aggregate. In Chapter 2 and 3, I show that, through copolymerization with
other monomers such as N-tert-butylacrylamide (NT) or butylacrylate (BA), thermo-responsive
behavior and mechanical strength of PNIPAM can be tailored to create a hydrogel that shape-fills
upon injection at a wound site, adapting to irregular margins and sealing traumatic injuries. This
engineered thermoresponsive sealant (TRS) demonstrated a significantly lower transition
3
temperature of 25°C, ideally suited for injection through an ocular wound at relatively lower
viscosity and a rapid gelation to form a soft-solid perfect seal. The thermosensitive behavior allows
the TRS to be easily removable by the application of cold water.
We used this copolymer material to develop a reversible approach to temporarily occlude
penetrating injuries to the posterior segment of the eye. In chapter 4, I discuss development of a
custom tool that controls the hydrogel temperature to enable effective deployment in the eye.
Feasibility studies were conducted using ex vivo and in vivo models of ocular trauma in rabbits,
and two user feedback workshops were held where military ophthalmologists tested the prototype
systems for ease of use, general concept, and performance in an ex vivo porcine model of ruptured
globe. These will be discussed in chapter 5.
Strong adhesion to biological tissues is one of the essential properties of tissue adhesives.
Therefore, after successfully developing a thermo-responsive material that physically seals open
globe eye injuries, we aimed to increase the adhesion of the hydrogel to tissue in order to develop
a wound closure adhesive. Wound closure adhesives are an increasingly common alternative to
longstanding and invasive clinical practices. Despite some drawbacks discussed above,
cyanoacrylates, fibrin glues, and modified polyethylene glycol (PEG)-derived bioadhesive
products have all seen increased FDA approvals in the past decade, offering novel approaches to
trauma mitigation for indications from internal tissue attachment to external skin grafts. Increased
off-label use of these adhesives can also be observed in literature—especially on ocular repair—
as patients benefit from replacement of cumbersome suturing techniques with minimally invasive
alternatives (17-19). Notably, the first sealant approved for corneal incisions in the United States,
ReSure®, was validated by a significantly lower leak rate than sutures following intraocular lens
implantation (4.1% vs 34.1%) (20). Another ocular bandage available in Europe, OcuSeal®, also
4
successfully treats clear corneal incisions with improved outcomes to suturing (21). Both new
technologies are designed to allow ocular healing and then undergo a natural degradation,
eliminating need for follow-up surgery. These technologies exemplify an ongoing transition from
conventional standards of care towards responsive materials that can interface with human tissue
or further promote regenerative processes. Despite their popularity elsewhere, no bioadhesives are
approved for closure of scleral wounds in the US (22).
Following our observations of TRS adhesive behavior we anticipated that the reversible
hydrogel could be also engineered to make a topical adhesive, to be applied to the outside of the
eye. PNIPAM or poly(NIPAM-co-BA) copolymers themselves, when applied to the surface of the
sclera, do not have sufficient adhesive strength to seal wounds and maintain ocular pressures above
10 mm Hg. The goal of the work presented in chapters 6 and 7 is to develop thermo-responsive
adhesives that will seal ocular wounds, maintaining normal intra-ocular pressures, at physiological
temperatures and have little or no adhesion to the ocular tissue at low temperatures (i.e. 5°-10°C).
While effective mechanical wound closure by TRS formulation was demonstrated for closure of
scleral tears, further development of this smart material could improve adhesive behavior and merit
application in other realms.
PNIPAM hydrogels have a long history of modification by both co-polymerization and
addition of co-solutes (23, 24). Addition of salts, surfactants, and organic solvents can raise or
lower the LCST, and their effects on copolymer behavior have been well characterized (25). For
instance, addition of many salt additives to PNIPAM increases polymer networking interactions,
both decreasing the LCST and facilitating the formation of larger aggregates. These effects have
been attributed to direct increase in solvent hydrophobicity, hydrogen bonding to the NIPAM
amide, and ionic binding to the polymer structure (26-29). In the case of larger molecular additives,
5
solubility or hydrophobicity were determined less impactful on LCST decrease than hydroxyl
group presence and steric availability (30). Within the field of hydrogel modification, polymer-
additive interactions have arisen as a simple route to assemble adhesive and tunable polymeric
materials without the need for complex synthetic approaches. Several systems have been reported
utilizing additive components to alter the physical properties of PNIPAM and its thermal behavior.
However, the thermo-responsive hydrogels presented to date are limited by low adhesion and poor
mechanics or require challenging, costly and poorly scalable synthesis of molecular components
through protein engineering or complex, multi-step chemical functionalization.
Improved bioactivity and adhesion has been established previously through combination
of the hydrogel with mixtures of extracellular matrix protein (31, 32). This methodology, while
effective, invites biofouling and disease depending on the protein serum source (33).
Functionalization of PNIPAM using the tissue adhesive peptide Arg-Gly-Asp (RGD) has also been
explored, but only by theoretical consideration of the process; no synthesis or characterization of
these compounds has been achieved (31). In tissue engineering applications, other hydrogels such
as PEG have been functionalized with poly-L-lysine or RGD to improve tissue culture
survivability through their cell-adhesive activity (34, 35). Conventional addition of this functional
component requires molecular bonding of the peptide to the molecule (36). Previous studies
suggest, however, that PNIPAM’s unique response to additive molecular structure can be elicited
by simply mixing the additive into the hydrogel. Inspired by these approaches, we selected two
different additives, with potential bioactive or adhesive capability, to explore this method of
hydrogel modification on the previously designed TRS gel.
Using this approach, we developed a highly elastic reversible adhesive (HERA) hydrogel.
HERA is an additive-integrated TRS with enhanced tissue adhesiveness that creates a strong
6
reversible adhesive for use in a range of wet biological environments. Our efforts focused on
mixture of Arg-Gly-Asp-Ser (RGDS) peptide and 3-Guanidinopropionic acid (GPA) with TRS, a
hydrated copolymer of NIPAM and BA. RGDS was selected for its well-documented bioactivity
and cell-adhesive properties. GPA, while unexplored as a tissue adhesive, is a widely available
protein supplement that has been noted to perform as an antihyperglycemic (37, 38). It shares the
guanidinium structural motif with the Arginine amino acid in RGDS, offering a potential
recognition region for cell attachment. Importantly, when initial isolation of the RGDS cell-
adhesion motif was characterized by Pierschbacher and Ruoslahti in 1984, removal of the Arginine
amino acid was shown to lead to elimination its effects (39), arguing for the important role of this
residue in cell adhesion. Both additives were prone to charges in water, offering changes in
aggregation behavior in addition to improved adhesion. In chapters 6 and 7, I present physical
properties and in vitro analysis of adhesive behavior in additive-modified, temperature-responsive
hydrogels and characterization of underlying molecular changes relative to the unmodified TRS.
7
1.2 Study design
The overall objective of this study was to assess the feasibility of developing a temporary ocular
repair technology with physical and mechanical properties suited to improving outcomes for
globes compromised by scleral tears. This objective was segmented into three efforts: a polymer
chemistry investigation of hydrogel properties correlated to synthesis route and composition, a
materials science study of performance characteristics using benchtop models of ocular trauma,
and an in vivo, preclinical study of safety and performance in an animal model of ocular trauma.
Our initial work evaluated the potential of two compositions at different concentrations. These
copolymers incorporated hydrophobic monomers (NT and BA) with NIPAM to form a
temperature responsive hydrogel in water with improved mechanical strength. Preliminary
evaluation consisted of transition temperature’s shift verification by scattering intensity, followed
by examination of storage and loss moduli for relative viscoelastic properties. Identification of
poly(NIPAM-co-BA), N95BA5, as the preferred sealant was followed by compression, tension
analysis, and tissue adhesion tests at eye temperature. Secondary evaluation consisted of in vitro
and ex vivo testing of different concentrations of the N95BA5 hydrogel in appropriate porcine test
models. These tests led to further narrowing of the potential candidate compositions to only the
30% (w/w) N95BA5 for advancement to in vivo performance and safety testing.
The sealant technology safety and performance were assessed in an in vivo model of ocular
trauma. A 2:1 (treatment/control) randomized and unblinded study was designed to evaluate the
hydrogel sealant against the envisioned standard of care (no intervention). All animals (n = 18)
received a 3-mm full-thickness laceration through the sclera, and treatment arm animals (n = 12)
received the hydrogel sealant. All animals had regular IOP measures taken in triplicate at each
time point to calculate a mean IOP for each time point. Animals were followed to one of three
8
study endpoints (48 hours, 1 week, and 4 weeks), after which eyes were enucleated and fixed for
histological preparation and analysis. Tissue segments were examined for evidence of adverse
tissue responses including chronic inflammation, retinal degradation, cytotoxicity, and
neurotoxicity.
In addition, we synthesize, characterize and validate, a highly elastic reversible adhesive
(HERA) for ocular wound adhesive. The enabling technology is a thermo-reversible, PNIPAM-
based, hydrogel which is adhesive to tissues at body temperature and non-adhesive at low
temperature. A series of smart hydrogels were engineered, characterized to improve the hydrogel
adhesion to ocular tissue and validated in vitro by scleral adhesion tests and intraocular pressure
measurements.
We discovered two different additives that can be combined with the hydrogel polymer
chemistry to further enhance the adhesion to tissues. The properties of engineered thermo-
responsive, additive-enhanced hydrogels were characterized by rheology, dynamic light scattering,
uniaxial adhesion test, and intraocular pressure experiments. Using an in vitro model of open globe
trauma, hydrogel functionality was then assessed to demonstrate the maximum intraocular
pressure. The results revealed that very small amounts of Arg-Gly-Asp-Ser peptide or 3-
Guanidinopropionic acid, as an additive, can improve the maximum pressures by a factor of six
compared to unmodified hydrogel. These observations suggest that the inclusion of the additives
can improve the cohesion within polymeric hydrogel network and adhesion strength against tissue.
9
CHAPTER 2
2.1 Engineered injectable TRS and mode of function
Our goal was to develop a biocompatible, tissue repair technology for temporary intervention at
sites of scleral tissue damage or loss. The approach presented here involves injection of a
purposefully nonbiodegradable liquid sealant capable of body heat–induced gelation to create a
size-adaptable solidified occlusion, which can both restore intraocular pressure (IOP) and be easily
removed within a few days for subsequent treatment without concomitant tissue damage. This
technology would afford the patient a larger window of time to complete surgical intervention
without requiring specialty equipment such as surgical microscopes for implementation. The
system consists of two components: (i) a thermoresponsive hydrogel that, by its reversible
transition from liquid to solid, conforms to wound shape and aggregates to mechanically seal
scleral penetrations and (ii) a custom deployment tool to contain the hydrogel at a controlled
temperature and to inject it at the targeted site (Fig. 2.1).
Fig. 2.1: Design of a thermo-responsive hydrogel to seal scleral perforation.
To implement a hydrogel sealing of scleral perforation, a temperature-mediated adhesive that
adapts to wound shape and a tool for its deployment at a perforation site are used here.
10
2.2 Synthesis and characterization of the N85NT15 and N95BA5
copolymers
The physically cross-linked hydrogel with thermoresponsive behavior was designed from
customized copolymer and water. The reversible transition from water-soluble coils to
hydrophobic globules at the LCST changes the physical properties of the gel (Fig. 2.2).
Fig. 2.2: The sealing hydrogel transitions from hydrophilic coils to hydrophobic globules at its
lower critical solution temperature (LCST), markedly changing physical properties.
11
PNIPAM and its copolymers have been extensively explored, but the combination of
NIPAM and BA to target a biologically relevant transition temperature and mechanical properties
pursues a unique trajectory. Previous studies have imbued custom copolymers with the
temperature sensitivity of PNIPAM, achieving pH-resistant graft copolymers, altered LCSTs, and
temperature-dependent drug diffusion (13, 40). Responsive properties of PNIPAM have also been
applied to acylated polymer bound parylene C for temperature-mediated tissue binding (41). With
these dynamic modifications in mind, PNIPAM was an obvious foundation for an impermanent,
biological sealant. To be used in the eye at 32°C, however, the LCST needed to be lowered.
Copolymerization of NIPAM with hydrophobic NT and BA not only decreases the LCST
of PNIPAM but also improves the polymer’s mechanical properties (42-45). By altering the
compositional ratios of these copolymers, a range of formulations were synthesized via free radical
polymerization and then characterized to optimize molecular weight, LCST, aqueous solution
concentration, and viscoelastic properties (Fig. 2.3). Previous studies have used both comonomers
12
Fig. 2.3: Schematic synthesis routes of the N85NT15 and N95BA5 copolymers. (A) poly(NIPAM-
co-N-tert-butylacrylamide), (B) poly(NIPAM-co-butylacrylate).
Copolymers of NIPAM and N-tert-butylacrylamide (N85NT15), and copolymer of NIPAM and
butylacrylate (N95BA5) were synthesized by free radical polymerization using 2,
2’azobisisobutyronitrile (AIBN) as initiator, under positive nitrogen at T = 50°C for 24 hours.
to create cell culture supports and other thermoresponsive polymers but not with synthetic
uniformity or properties that met our needs (43-45). As a result, we followed a different synthetic
scheme from those found in the literature and performed the necessary measurement of LCST and
molecular weight for all formulations: homopolymers of NIPAM, copolymer of NIPAM with NT,
and copolymer of NIPAM with BA, each using NIPAM as the main formulation component (Fig.
2.4). Composition ratios for each sample were verified by
1
H nuclear magnetic resonance (NMR)
(Fig. 2.5).
13
Fig. 2.4: Molecular structures of formulations of poly(N-isopropylacrylamide) (PNIPAM) with
butylacrylate (BA) or N-tert-butylacrylamide (NT) and a table of resulting alterations of molecular
properties and LCST values.
14
Fig. 2.5:
1
H NMR spectrum for N85NT15 and N95BA5 copolymers in CDCl3.
Proton NMR spectra for poly(NIPAM-co-NT) and poly(NIPAM-co-BA) in CDCl3 verified the
N95BA5 and N85NT15 molar conformations of the polymers. In both NMR spectra of hydrogels, a
clear peak around 1.10 ppm (a) signifies the shielded isopropyl methyl groups of the NIPAM
monomer. For the NT included copolymer, shielded tert-butyl protons appear at 1.30 ppm (e).
Comparison of isopropyl methyl peak (a) to tert-butyl peak (e) confirmed the feed ratios with high
accuracy. In regard to the N95BA5 co-polymer, a very shielded 0.89 ppm peak (e) with low
integration represents the methyl protons at the end of the carbon chain. The isopropyl methyl peak
(a) was not well resolved in this NMR, lacking a baseline and hinting at overlap with other peaks.
This required the comparison of the downfield peak at 3.96 ppm (b+h), representing the protons
deshielded by electronegative atoms, to the butylacrylate methyl peak (e) for a metric of feed ratio
accuracy. Calculated compositions for the copolymer from 1H NMR spectrum are close to the
stoichiometric ones, demonstrating relatively accurate feed ratio.
We deliberately engineered the LCST of the polymers studied here to fall below that of
PNIPAM to achieve a transition temperature for the hydrogel that is well below the eye’s
physiological temperature. The contributions of the NT and BA monomers to the phase transition
were examined through scattering intensity measurements of aqueous PNIPAM, poly(NIPAM-co-
NT) (N85NT15), and poly(NIPAM-co-BA) (N95BA5) solutions over temperature ranges that
included their phase transition temperature (Fig. 2.6). Higher scattering intensity was observed at
higher temperatures (above phase transition temperature), which was attributed to transformation
from a more soluble coil conformation below the LCST to a largely insoluble compact
conformation. For instance, the scattering intensity values of N95BA5 exhibited sharp increases
around 16°C, indicative of a gelation point.
15
Fig. 2.6: Scattering intensity spectra of N95BA5 as a function of temperature.
Measurements of scattering intensity over temperature ranges spanning conformation LCSTs
allowed preliminary characterization of the gelation process. While the increase in scatter is
attributed to particle size, the decreasing region was attributed to the opaque surface of the sample
gradually preventing beam entry beyond the surface of the sample. The distributions for N95BA5
shown above were indicative of the desired shift in PMIPAM LCST. For the sake of clarity, the
scattering intensity distribution graphs in the whole range of temperature not included here.
N95BA5 [5% (w/w)].
Comparison of temperature-dependent scattering intensity distributions of three different
hydrogels confirmed a PNIPAM gelation point around 32°C. By inclusion of only 5% BA or 15%
NT, the gelation points shifted to 16° and 22°C, respectively (Fig. 2.7). We successfully engineered
smart hydrogels, N85NT15 and N95BA5, with appropriate transition temperatures for human eye
application.
16
Fig. 2.7: Normalized
scattering intensity as a
function of temperature for
PNIPAM, poly(NIPAM-co-
NT) (N85NT15), and
poly(NIPAM-co-BA)
(N95BA5). A.U., arbitrary
units.
2.3 Optimizing rheological properties of shape-persistent, moldable
TRS
Selection of the ideal copolymer conformation required examination of temperature-dependent
aggregation behavior. Previous studies of PNIPAM copolymer behavior have used viscosity and
DLS measurements to characterize particle size and shape (46). The viscoelastic properties of the
hydrogels were determined by rheological analysis. Conventional elastic responses at LCST were
demonstrated for both N95BA5 and N85NT15 samples (47). The loss (G″) and storage (G′) moduli—
representations of the viscous and elastic behavior, respectively—were measured by strain
amplitude sweeps across a range of temperatures (Figs. 2.8 and 2.9). Without a high-enough
storage modulus above LCST, the hydrogel would not be sufficiently elastic to resist intraocular
pressures while maintaining an effective occlusion. Without an optimized loss modulus below
LCST, the hydrogel might be too runny or thick, making it difficult to apply.
17
Fig. 2.8: Strain amplitude for N95BA5 and N85NT15 at T = 6°C.
Change of loss modulus (G”) with strain amplitude for N95BA5 and N85NT15 [10% (w/w)] at T =
6°C. At low temperature (below LCST), samples are liquid and there is no storage modulus. The
loss modulus results indicated that the co-BA composition was more viscous, foreshadowing later
viscoelastic superiority.
18
Fig. 2.9: Strain amplitude of N85NT15 at its LCST.
Storage modulus (G’) and loss modulus (G”) values of N85NT15 hydrogel [10% (w/w)] on strain
sweep at its LCST. The data obtained at transition temperature of NT-included hydrogel indicated
the same trend observed for N95BA5 (Fig. 2.10), storage and loss modulus values are almost
constant for strains smaller than 2% (critical strain), following a sharp decrease above that. Despite
strain sweep of N95BA5, we observed higher loss modulus value than storage modulus for
N85NT15.This trend across temperature supported the conclusion that these NT-included hydrogels
had a weaker overall structure than co-BA compositions.
The storage modulus for N95BA5 rapidly decreases above the critical strain region (1%),
indicating gel collapse to a quasi-liquid state (Fig. 2.10). For different temperatures, the respective
G″ (6°, 24°, and 32°C) and G′ (24° and 32°C) values of N95BA5 and N85NT15 were measured as a
function of angular frequency at fixed strain (0.1%) (Fig. 2.11). We observed that hydrogel
dynamic moduli depend on the temperature and that both samples reached their maximum
viscosity and elasticity at 32°C. Although the mechanical strength of both hydrogels showed the
same trend, N95BA5 generated stronger polymeric networks. Replacement of NT comonomer with
BA caused the G″ and G′ values to increase by a factor of up to 30.
Comparison of the viscoelastic profile of 10% (w/w) aqueous solutions of two different
copolymers at eye temperature (32°C) differentiated their aggregation behavior. G′ dominates over
G″ for N95BA5, resulting in a quasi-solid state (tan δ ≡ G″/G′ ≈ 0.4). Under the same condition,
viscosity remained relatively greater than the elasticity for N85NT15, indicating a quasi-liquid state
(tan δ ≡ G″/G′ ≈ 2.3) (Fig. 2.12). N95BA5 was selected for further characterization, owing to its
mechanical strength and desirable phase transition temperature.
19
Fig. 2.10: Storage and loss moduli (G′ and G″, respectively) over strain for N95BA5.
Fig. 2.11: G″ (B) and G′ (C) modulus representations of viscoelastic behavior for co-BA and co-
NT as a function of angular frequency at fixed strain.
20
Fig. 2.12: Table of storage moduli for co-NT and co-BA compositions at different angular
frequencies. Storage and loss moduli (G′ and G″, respectively) over frequency for co-NT and co-
BA compositions.
Viscosity measurements of N95BA5 provided a better understanding of the hydrogel’s
conformational change (Fig. 2.13). Below the gelation point, complex viscosity, η*, was
independent of temperature and constant. At and above the phase transition region, the η* rose
sharply and approached a constant value. This change resulted from copolymer dehydration,
compact globule formation, and resulting polymeric networks between macromolecule chains.
21
Fig. 2.13: Measurement of complex viscosity for N95BA5 according to temperature.
The complex viscosity profiles of N95BA5 solutions [5, 20, and 30% (w/w)] were then
evaluated to compare their concentration dependent strength. Earlier gelation onset accompanied
higher concentration, indicating that more concentrated polymer solutions form gels at
respectively lower temperatures. The η* values also confirmed that higher copolymer
concentrations can improve the mechanical strength of noncovalent hydrogels. Among the three
samples studied, 30% (w/w) had the greatest complex viscosity value of about 10,000 centipoise
(cP) (10 Pa · s), suggesting a strong yet injectable thermoresponsive hydrogel (Fig. 2.14). For
comparison, the viscosities of honey and ketchup, two easily applied yet malleable substances, are
about 3000 and 50,000 cP, respectively. The 30% (w/w) N95BA5 was chosen for having the most
useful viscoelastic properties of the hydrogels prepared herein (Fig. 2.4).
22
Fig. 2.14: Complex viscosity of N95BA5 as a function of temperature and concentration.
Complex viscosity, η*, measurements for three different concentrations of N95BA5 [5, 20, and 30%
(w/w)] indicated a slight decrease in LCST (~1-2°C) with increase in concentration. Initial
viscosity at low temperature confirmed the role of copolymer concentration in hydrogel behavior
in the liquid form. Temperature-mediated transitions showed rapid rises in viscosity following
movement beyond the LCST with rapid gelation beyond the 25°C region. Higher sample
concentration corresponded to an earlier elevation of viscosity from baseline before peak,
indicating that greater amounts of N95BA5 copolymer favor the hydrophobic state.
2.4 Physical characterizations of the TRS
The stiffness and resilience of the 30% (w/w) N95BA5 were evaluated using standard tensile and
compression tests and by casting it into shapes. Compressive stress-strain curves illustrated a
positive correlation between TRS concentration and compressive moduli. The results revealed that
the mechanical properties of the developed hydrogel can be readily tuned by changing the hydrogel
23
concentration, with compressive moduli modified from less than 15 to ≈55 kPa when doubling the
TRS concentration from 15 to 30% (Fig. 2.15).
Fig. 2.15: Compressive stress-strain characterization and compressive modulus of the N95BA5
hydrogel.
24
To determine the effect of copolymer concentration on the mechanical properties of TRS
hydrogels, compression tests were performed across three different concentrations [15, 20, and
30% (w/w)]. Hydrogels transitioned from liquid to solid gel in a glass mold and maintained at
32°C during testing, enabling characterization of the post-application plug properties. The stress-
strain curves for all three hydrogel concentrations demonstrated that increasing copolymer
concentration increased the stiffness at all strains, further confirming that higher concentrations
are more durable. In addition, the compressive modulus of TRS hydrogels increased with
increasing copolymer concentration, 14 kPa (15%) to 53 kPa (30%), suggesting higher copolymer
concentration has the capacity to withstand greater forces without buckling.
The elastic modulus of N95BA5 was characterized as a function of the concentration (20,
25, and 30%). TRS tensile stress-strain curves followed a similar trend, with increasing copolymer
concentration resulting in higher Young’s modulus. The 30% TRS hydrogels were found to have
elastic moduli of 117 kPa, representing an improved stretching capacity compared to 60 kPa (25%)
and 45 kPa (20%) (Fig. 2.16).
The resulting molded objects (Fig. 2.17, A to C) were resilient and sufficiently cohesive at
body temperature to be suspended horizontally (Fig. 2.17, D), vertically (Fig. 2.17, E), and by hand
(Fig. 2.17, F). The noncovalent properties of the hydrogel also allowed it to self-heal when slightly
deformed. The hydrogel satisfied critical requirements for application as an ocular sealant:
moldability, persistence in form, and sufficient toughness to withstand intraocular pressure.
25
Fig. 2.16. Tensile stress-strain characterization and tensile modulus of the N95BA5 hydrogel.
Elastic moduli of N95BA5 were characterized as a function of the concentration [20, 25, and 30%
(w/w)] at T = 32°C. TRS tensile stress-strain curves followed a similar trend with increasing
copolymer concentration resulting in higher Young’s modulus. 30% TRS hydrogel was found to
26
have elastic modulus of 117 kPa, representing improved stretching capacity compared to 60 kPa
(25%) and 45 kPa (20%).
Fig. 2.17: (A to C) Images of square and round gel molds formed by heating the hydrogel solutions
to 32°C. Scale bars, 1 cm. (D to F) Images of solid N95BA5 hydrogel demonstrating horizontal
resilience (D), vertical resilience (E), and strength of form on contact (F). Scale bars, 1 cm.
Our findings verified the concentration-dependent properties of hydrogel aggregation for
this engineered conformation—higher concentration slightly lowered LCST —and reaffirmed
non-Newtonian characteristics observed in studies of additive effects (48, 49). The quasi-solid
state of the 30% N95BA5 hydrogel, in conjunction with its lowered LCST, distinguished it as the
ideal candidate. Potential further improvements to the aggregation process could be explored via
the heating protocol as a sample’s temperature history can affect rheological properties (50).
27
2.5 Materials and methods
2.5.1 Homopolymer and copolymer synthesis
PNIPAM, copolymer of NIPAM and NT (N85NT15), and copolymer of NIPAM and BA (N95BA5)
were synthesized using free radical polymerization (51, 52). For N85NT15, a solution of NIPAM
(4.25 g), NT (0.75 g), and 2,2′-azobisisobutyronitrile (0.021 g) was dissolved in 60 ml of dry
tetrahydrofuran (THF). The magnetically stirred solution was degassed, heated to 50°C for 24
hours under positive nitrogen pressure, and allowed to cool. The reaction mixture was filtered
(0.45-mm Teflon filter), and the filtrated volume was reduced by half. Ether was added with
mixing to precipitate the copolymer. The precipitate was filtered off, washed with ether, and dried
under vacuum to yield dry 4.64 g of copolymer product. For N95BA5, we followed the same
procedure except that we used a different ratio of NIPAM (4.75 g) to BA (0.25 g). Using different
ratios of THF/benzene as a solvent, we were able to synthesize homopolymers and copolymers
with various molecular weights and polydispersities. After differentiation of the hydrogel
compositions and the identification of N95BA5 for further application, the copolymer was then
purchased from Sigma-Aldrich.
2.5.2
1
H NMR
Feed ratio accuracy was confirmed by 1HNMR(VarianVNMRS-600). We prepared 5% (w/w)
solutions of N85NT15 and N95BA5 in CDCL3. Peak integration ratios were compared to theoretical
ratios for compositional verification.
2.5.3 Hydrogel solution preparation
28
The required amount of PNIPAM, N85NT15, and N95BA5 was weighed for various concentrations
of hydrogel aqueous solutions, from 0.8 to 43.2% (w/w). The powder was added directly to the
sterile water. The vial containing the polymer suspension was then processed with a Misonix
Sonicator 3000 using a cup horn high-intensity ultrasonic water bath at the maximum power setting
(10) due to sample viscosity. A circulating temperature control water bath held at 2°C prevented
sample heating due to prolonged horn activity. The sample was sonicated until a transparent clear
hydrogel was obtained. The required sonication time ranged from1 to 30 hours and depended on
the hydrogel concentration and molecular weight of the homopolymers/copolymers.
2.5.4 Scattering intensity analysis
Scattering intensity measurements of 5% (w/w) hydrogels were carried out using a Photon
Technology International Quantamaster fluorescence spectrophotometer. All intensity
measurements were performed at an excitation wavelength (lex) of 450 nm, with detectors
positioned 90° from the light source. Emission spectra were recorded with a slit width of 0.1/0.1
nm. Quartz cuvettes (1 cm × 1 cm × 3 cm) containing the sample were placed in a cell holder,
which was electrothermally controlled at a precise temperature regulated by a Peltier cooler. Each
hydrogel was left under undisturbed conditions for 20 min at each temperature to obtain
thermodynamic equilibrium.
Because the input light source was directed at the sample, the particles redirect the light
away from a direct course through the cuvette with increasing severity as particle size increases.
For small particle sizes, most of the input beam traverses the sample directly, resulting in small
scattering measurements. As particle size increases, more of the light affects the particles, resulting
in sideways diffusion of light that could be focused and then received by the sensor at 90° to the
29
input beam. Scattering values were then normalized. Because of increase in sample opacity above
LCST, most of the input beam was eventually reflected upon initial contact with the sample. As a
result, higher temperatures saw a slow decrease in intensity because less input beam reached the
center of the sample to be scattered toward the receiver.
2.5.5 Rheological analysis
An Anton Paar Modular Compact Rheometer (MCR) was used to measure the rheological
properties of hydrogels. Eight milliliters of 10% (w/w) N95BA5 and N85NT15 hydrogel solutions
was placed into the cylinder with special care to avoid evaporation of water. First, the hydrogels
were investigated with strain-amplitude sweeps below their LCST (6°C), at LCST, and above the
transition temperature (32°C) at constant angular frequency (10 rad s−1). After the critical strain
was found for each hydrogel, oscillation tests were performed to measure the loss (G″) and storage
(G′) moduli at a designated temperature (6°, 24°, and 32°C) using a circulating temperature control
water bath. Strain was fixed at 0.1% (below the critical strain), and moduli were measured as a
response to logarithmic angular frequency ramp from 0.1 to 1000 rad s−1. Temperature-dependent
changes in complex viscosity were performed using a fixed angular frequency of 10 rad s−1, 0.1%
strain, and a heat rate of 0.5°C/ min on 5, 10, 20, and 30% (w/w) N95BA5 samples.
2.5.6 Compression test
To perform compression testing, liquid hydrogels were injected into a 17-mm × 17-mm × 5-mm
glass mold and then heated to 32°C using a temperature-controlled Instron 5567 mechanical tester.
After a 2-min curing period, the glass mold was removed, and the probe was placed in contact with
the top of the molded hydrogel. The samples were then compressed at a rate of 1 mm/min. A stress-
30
strain curve was obtained, and compressive modulus was determined as the slope of the linear
region corresponding to 5 to 15% strain. The number of hydrogel samples was three per group.
2.5.7 Tensile test
Liquid hydrogel samples were shifted to the solid phase (32°C) using a hot plate and were detached
from a metal mold. Samples were then blotted dry and fixed by two clamps of an Instron 5942
mechanical tester. The solid samples were stretched at a constant rate of 1 mm/min under
controlled temperature. Young’s modulus was determined as the slope of the linear region of the
stress-strain corresponding to 0 to 10% strain.
31
CHAPTER 3
3.1 Ex vivo testing of N95BA5 for ocular trauma
Successful completion of the hydrogel design and characterization enabled advancement to
preliminary ex vivo and in vitro performance and safety assessments. Sealant efficacy of the
hydrogel was first assessed in an ex vivo cadaveric porcine eye model of ocular trauma.
Cyanoacrylate was selected over fibrin-, albumin-, and polyethylene glycol–based adhesives as a
positive control because of its well-documented superiority in maintaining IOP and uniaxial
adhesion strength (53, 54). The hydrogel was applied to the
lacerated eye model after an incision procedure (Fig. 3.1).
Fig. 3.1: Schematic (top) and image (bottom) depicting ex
vivo procedures carried out in a pressure-controlled
explanted cadaveric pig eye.
The hydrogel was injected into the posterior
chamber of the eye through the perforation, and the injection
tool was slowly retracted while continuously deploying the
hydrogel, leaving a sealant trail through the wound. At the
exterior surface of the sclera, additional hydrogel was
deposited, creating a gel rivet like cap. The caps were left to
settle for several seconds, allowing them to increase in
temperature and dehydrate before being cut away or
smoothed flat (Fig. 3.2). IOP was then controllably raised by
infusion with warm saline from the cannula inserted in the posterior chamber.
32
Fig. 3.2: Images (top) and schematic depiction (bottom) of hydrogel injection through scleral
perforation by deployment of a sealant trail through the wound, leaving rivet-like caps
subsequently removed to leave the occlusion flush with the scleral surface.
By comparing the IOP values established at eye temperature (32°C) by 15, 20, and 30%
(w/w) aqueous solutions of N95BA5 and N85NT15, N95BA5 was determined to be the superior
hydrogel, maintaining ocular pressures above 70 mmHg for all concentrations. Under the same
conditions, IOP values for 15 and 20% N85NT15 were less effective, containing pressures of 5 and
40 mmHg, respectively (Fig. 3.3). Although high concentrations (30%) of both N95BA5 and
N85NT15 hydrogels were capable of maintaining pressures of up to 72 mmHg, leakage was
observed for N85NT15, indicating a quasi-sealed state. Even so, the formulations identified as
effective could withstand IOPs of up to 77 mmHg without leakage, an environment approximately
five times the physiological IOP range. Both the N95BA5 samples and a cyanoacrylate sealed
positive control held the maximum pressure (78 mmHg) of our experimental setup, demonstrating
33
effective equivalence. In contrast, our negative control without a sealant held no pressure. Figure
3.3 shows a hydrogel plug removed from one of the rabbits at the end of the study.
Fig. 3.3: Comparison of maintained intraocular pressures across a concentration spectrum for
N95BA5 and N85NT15 hydrogels (n = 3 per group) (left). Image of a solid plug removed from a test
eye (right).
Although we designed the hydrogel as an ideal mechanical sealant, in vitro uniaxial
adhesion tests were performed to provide accurate comparison between the adhesion force of
cyanoacrylate, the strongest ocular adhesive, and our TRS. The hydrogel or cyanoacrylate was
sandwiched between two pieces of dissected scleral tissue that were fixated to the base and actuator
arm of a pull tester (Fig. 3.4).
34
3.2 In vitro adhesion testing on TRS hydrogel
Fig. 3.4: Schematic depicting
tissue adhesion tests comparing
PNIPAM, N95BA5, and
cyanoacrylate adhesion strength
to scleral tissue ex vivo.
Apposed tissues were brought down and pressed together using 15g of pressure for 2 min. The
actuator was then pulled until the two tissue samples detached. In addition to the 30% N95BA5,
different aqueous concentrations of PNIPAM ranging from 0.8 to 43% (w/w) were chosen to assess
the impact of concentration. Adhesion results of homopolymers suggested that an increase in the
aqueous concentration of the hydrogel increased the adhesive strength between the hydrogel and
scleral tissue. Adhesion forces between 336 and 560 mN were observed for the 30% (w/w) N95BA5
before adhesion failure at the tissue-hydrogel interface. Under the same uniaxial conditions,
cyanoacrylate’s adhesion force was slightly higher (Fig. 3.5). Wetting, diffusion, and adsorption
theories offer mechanical and molecular explanations for this behavior but were not explored in
this study (55, 56).
35
Fig. 3.5: Adhesion force of different concentrations of N95BA5, PNIPAM, and cyanoacrylate to
scleral tissue.
Columns show 478 ± 71 and 639 ± 110 for N95BA5 and cyanoacrylate, respectively (n = 3 per
group).
3.3 In situ gelation mechanism of TRS in the eye
Ex vivo application indicated that to occlude the incision properly, the hydrogel must change from
transparent liquid to white solid (Fig. 3.2). We used dynamic light scattering (DLS) techniques to
obtain particle size in solution, gel and solution-to-gel transition states, and particle impact on
gelation kinetic and mechanical properties. At 2°C, below LCST, we observed that 96% of the
intensity corresponds to particles with a small hydrodynamic radius (3.6 nm), attributed to single
polymer chains in the coil conformation (Fig. 3.6).
36
Fig. 3.6: N95BA5 particle size as assessed using dynamic light scattering in solution below LCST
(2°C), showing a 96% scattering intensity for small-radius particles.
At the phase transition region, we observed two peaks with comparable intensities. One
peak (236.2 nm) corresponded to N95BA5 molecules that were partially aggregated in the solution,
and the other peak (4.5 nm) corresponded to lower-radius particles (Fig. 3.7). Further temperature
increase to 20°C caused 98% of the copolymer to form large aggregates (439 nm), a consequence
of more favorable polymer-polymer interactions (Fig. 3.8). Amide groups facilitated those
interactions by forming hydrogen bonds between polymer molecules, nesting themselves inside
the globules.
37
Fig. 3.7: Particle size in the phase transition region (12°C). A split was observed between particles
with hydrodynamic radius of 4.5 and 236.2 nm.
Intensity distribution across a range of temperatures (2° to 20°C) allowed further
characterization of the temperature-induced hydrophobic aggregation process (Fig. 3.9). For
brevity, the size distribution graphs of only four temperatures (2°, 8°, 12°, and 20°C) were included
(Fig. 3.10). With temperature increase, the first peak corresponding to smaller particles decreased
in size. Simultaneously, hydrodynamic radius (RH) and intensity of the high aggregation peak
region increased with heat (Fig. 3.10).
38
Fig. 3.8: Particle size toward the end of the phase transition region (18°C). Ninety-eight percent
of scattering intensity was due to large-radius N95BA5 aggregates.
39
Fig. 3.9: Intensity distribution graph of DLS spectra for N95BA5 recorded at different temperatures.
Intensity distribution graph of DLS spectra for N95BA5 hydrogel [5% (w/w)] recorded at different
temperatures including the phase transition temperature. Two clearly evident particle size
distributions indicated the transition from coil to globule and corroborated the trends observed in
scattering intensity and rheological analysis. The hydrogel temperature response was best observed
at 12°C as the intensity peak corresponding to particles of larger radius overtook the peak for
smaller radius geometry. By 16°C, the vast majority of particles were of the larger radius, but
measurement beyond 20°C was prohibited by sample opacity.
Fig. 3.10: Intensity distribution graph for N95BA5 as a function of temperature.
Analysis of RH values for N95BA5 aggregates provided detailed understanding of the gel’s
temperature dependence. At temperatures below the LCST, RH remained nearly constant (<10 nm)
up to 10°C, but increasing the temperature above the 12°C gelation onset condition led to a rapid
increase in the hydrodynamic radius and to sizes markedly larger than the PNIPAM aggregates
(57). Greater aggregate size makes for better ocular sealant because larger particles diffuse less
40
readily into the vitreous gel, as predicted by the Stokes-Einstein equation. The aggregation onset
temperature was confirmed by the results of scattering intensity tests. DLS measurements could
not be performed above 20°C because of the turbidity gain of the hydrogel solution (Fig. 3.11). In
summary, the nanometric copolymer aggregates associate upon exposure to heat, becoming
insoluble in aqueous solution as a consequence of more favorable polymer-polymer interactions.
Fig. 3.11: Hydrodynamic radius size of particles traced through the hydrogel transition to show
higher aggregate populations at higher temperatures (n = 3 per temperature).
3.4 Thermo-induced volume change of molded TRS
As observed previously, heated hydrogel can form shape-persistent objects by in situ gelation. The
thermo-induced volume change of these molded shapes was further characterized to investigate
hydrogel stability in an environment similar to the eye. The swollen hydrogel underwent slight
volume reduction after only 5 min of exposure to 32°C deionized water. Shape persistence and
stability remained for 1 month in an aqueous environment above its LCST (Fig. 3.12). These
formed shapes were also maintained at 10°C, below LCST, resulting in complete structural
41
collapse after a few hours, consistent with the thermoresponsive trends observed (Fig. 3.13) Our
hydrogel challenges the preconception that materials held together by noncovalent forces and
mostly composed of water are weak.
Fig. 3.12: Gross images of molded hydrogel samples held at the expected eye temperature (32°C)
for up to 30 days, showing slight volume decrease and good shape persistence and stability.
42
Fig. 3.13: Hydrophobic/hydrophilic nature of the N95BA5 hydrogel above and below the LCST.
Gel resilience relied on the temperature of the aqueous environment with the mold in 10°C bath
completely deconstructed within hours. In contrast, 32°C bath hydrogel demonstrated shape
permanence and insignificant volume decrease for a month. Environmental degradation would
have disqualified the hydrogel from use in the ocular vitreous humor, but the form persistence of
the molded gel further indicated N95BA5 viability. N95BA5 [30% (w/w)].
Hydrogel plugs maintained IOPs greater than normal physiological IOP, suggesting the
possibility of prolonging the injury-to-operation window, but doing so with a reversible sealant.
In addition, the non-Newtonian state of N95BA5 allowed it to adapt to wound irregularities and
also exhibit self-healing characteristics. This resilience has been noted in other noncovalent
hydrogels, supporting the idea that these moldable materials can exhibit significant mechanical
strength (47). In contrast, the NT comonomer proved less effective at maintaining high IOPs in
the ex vivo test protocol, a result that paralleled reports of n-butyl cyanoacrylate scleral sealants
preventing high IOP leaks in similar models (58).
43
3.5 Materials and methods
3.5.1 Ex vivo IOP measurement
Fresh (harvested within 24 hours) porcine eyes (Sierra for Medical Science) were mounted into a
Styrofoam fixture and immobilized with dissection pins. Partial core vitrectomies were performed
on each eye using the Constellation Vision System vitrectomy console (Alcon Inc.), and IOP was
measured by a digital pressure sensor inserted in the vitreous cavity (posterior segment of the eye)
(Fig. 3.1). A single 3-mm linear incision was created in the sclera, about 3-mm distance radial
from the limbus, with the incision path running tangential to the limbus perimeter. An infusion
cannula was placed in the vitreous cavity on the opposite side from the incision to supply a saline
solution of about 37°C. Saline solution was infused either from a gravity-fed saline drip bag or
from a digitally actuated infusion system. Partial core vitrectomy allowed faster diffusion of the
liquid through the vitreous cavity, and saline ejection from the incision site confirmed incision
success. Once confirmed, the infusion line was clamped to limit leakage during the test. The eye
surface was dried with swabs to enable clean placement of the test substance (Fig. 3.1). After this
process, the IOP was gradually raised from baseline by manually increasing the infusion pump
pressure until leakage was observed or the pressure sensor value no longer increased with
increasing infusion rate (indicative of a nonvisible leak). The maximum pressure held (in
millimeters of mercury) was recorded.
3.5.2 In vitro adhesion test
We prepared a range of 0.8 to 43% (w/w) PNIPAM and 30% (w/w) N95BA5 hydrogels for uniaxial
adhesion testing. Fresh (harvested within 24 hours) porcine eyes (Sierra for Medical Science) were
dissected into a 2-cm × 2-cm square shape. Hydrogel was compressed between two pieces of
44
dissected scleral tissue fixated to the base and actuator arm of the pull tester (Fig. 3.4). Apposed
tissues were put into contact and pressed together using 15g of pressure for 2 min. The actuator
then performed a pulling motion until the two samples separated, compromising the gel adhesion.
3.5.3 Hydrodynamic light scattering analysis
Intensity distribution and RH of 5% (w/w) N95BA5 aggregates were measured by the automated
DynaPro Plate Reader II (Wyatt Technology). The DLS equipment is equipped with a thermostat-
equipped sample chamber to maintain desired temperatures within a range of 4° to 85°C with great
accuracy. A bubble-free sample of about 80 ml was introduced in a square glass cuvette through a
micropipette. A drop of mineral oil (20 ml) was placed on the top sample solution to prevent water
evaporation. Then, the sample cell was placed in the sample chamber of the DLS instrument and
kept at constant temperature for 30 min, a procedure repeated for all desired temperatures to ensure
thermodynamic equilibrium. This instrument measures the movement of particles under Brownian
motion and converts this motion into size by using the Stokes-Einstein equation, as given below,
D = kT / 6πηRH
where k is the Boltzmann’s constant, T is the absolute temperature, η is the viscosity, and D is the
diffusion coefficient. All data were obtained by the instrumental software.
45
CHAPTER 4
4.1 Design of a custom injection tool
Effective deployment of the hydrogel without premature transition required development of a
controlled-environment injector tool to meet optimal use requirements derived from clinical end
users (table 4.1).
Table 4.1: Injector tool design requirements.
Design parameters were created based on inputs from different team stakeholders: polymer
chemists, ophthalmologists, military medical researchers and engineers. The syringe-style design
46
was selected as the starting form factor because it helped satisfy several of these engineering and
performance needs. The ammonium nitrate reaction also satisfied the requirements, providing an
easily activated, inexpensive, and sustained gel cooling process that could be integrated with the
other device form requirements.
The preliminary design consisted of a 1-cm
3
chamber inside a larger 15-cm
3
thermal jacket
(Fig. 4.1) in an easily used form factor. On the basis of the preliminary design and volume of space
available inside the thermal jacket, a series of endothermic chemical reactions of ammonium
nitrate salt and water was performed and optimized to rapidly cool the hydrogel chamber to 0°C
within 60 s of mixing and to maintain the hydrogel’s temperature between 0° and 10°C for up to
10 min (Fig. 4.2).
Fig. 4.1: Design diagrams of a custom injection tool to effectively control hydrogel deployment
and regulate its temperature.
47
Fig. 4.2: Validation of a custom injection tool to effectively control hydrogel deployment and
regulate its temperature.
Fig. 4.3: An image of the prototype injector.
48
A working prototype was fabricated from off-the-shelf parts, including a 1-cm
3
syringe
nested inside a larger 15-cm
3
syringe and a soft removable loading port cap, which allowed
ammonium nitrate to be loaded into the jacket and water to be injected in the jacket when ready
for use (Fig. 4.3). Cooling characteristics were optimized by varying reactant concentrations and
tracking temperature transients to achieve the desired deployment window (Fig. 4.4). A polymer-
based catheter cannula was also used on the injector tip for hydrogel application to reduce thermal
conductivity from the eye, which could have induced temperature transition solidification in the
tool’s lumen during deployment.
Fig. 4.4: Injector tool cooling reaction calibration curves for 2.5 and 12.5 g of ammonium nitrate
to various volumes of added water. 2.5 g (left) and 12.5 g (right) ammonium nitrate.
The cooling (endothermic) reaction relies on the dissolution of the ionic salt, with different ratios
of solute to solvent affecting the rate of dissolution and total endothermic heat absorption.
Thermocouples integrated into the injector tool prototypes were used to identify solute-solvent
mixture ratios that met the desired performance criteria (cooling to T = 0°C in less than 60 seconds,
and maintenance of T < 10°C for at least 10 minutes). Graphs above illustrate temperature-time
curves generated from different mixture ratios with shorter operation time windows (left) and
greater initial cooling effects (right).
49
4.2 Effective TRS deployment in an in vivo model of ocular trauma
An in vivo pilot validation study was performed in a rabbit model of ocular trauma to assess the
ease of use, safety profile, and preliminary efficacy of the N95BA5 hydrogel. A 7-day in vivo
follow-up study was performed to compare the performance of the hydrogel wound closure versus
no intervention (control). A subset of the study animals were followed for an additional 3 weeks
(4-week endpoint) to more carefully examine safety and potential inflammatory responses to
hydrogel placement in the eye (Fig. 4.5 and table 4.2).
Fig. 4.5: Two-arm study design to assess safety and efficacy of the hydrogel versus the current
standard of care for posterior segment open globe injuries.
50
Table 4.2: In vivo study tabulated trajectory.
The in vivo study covered the treatment or trauma progression of 18 rabbits. Days followed and
measurements taken were tabulated to provide data tracking. Variable implant time courses
captured the intended use frame but also extended far beyond. *Animals for which the hydrogel
placement procedure was performed twice because the first attempt was not successful.
For each test run, a 3-mm full-thickness laceration was created in the sclera to simulate a
penetrating injury (Fig. 4.6, a). After trauma creation in a treatment group rabbit, the tool was
inserted into the laceration (Fig. 4.6, b) and the hydrogel was deployed (Fig. 4.6, c). Once
deployed, the hydrogel was then allowed to set for 5 min, enabling the dehydration transition from
translucent (Fig. 4.6, c and d) to an opaque white color (Fig. 4.6, e). Expelled water formed visible
droplets on the gel surface (Fig. 4.6, f). The rivet cap formed on the ocular surface was then cut
away to create a low-profile, flat head (Fig. 4.6, g to i). Ease of use of deployment was evaluated
by comparing the number of successful hydrogel placements into the eye versus the number of
hydrogel placement attempts.
51
Fig. 4.6: Images of the surgical procedure in rabbits.
A 3-mm, full-thickness linear incision was created in the sclera about 3 mm radial from the limbus,
followed by preparation and deployment of the hydrogel through the incision.
Despite limited training with the system, the veterinary surgeon was able to successfully
deploy the hydrogel on the first attempt in greater than 80% of the cases (table 4.2). During follow-
up, trauma sites in the control group were difficult to locate. The conjunctival epithelium formed
a fibrotic layer across the margins in an attempt to repair the breach. Although suggestive of a
natural healing process, the eyes still exhibited hypotony with IOPs of ≤4 mmHg, likely
attributable to a leaky, porous closure.
52
The scleral surface of all study eyes [right eye (OD)] was visually evaluated by an
ophthalmologist at 24 hours, 48 hours, and 1 week to look for signs of scleritis or other vasculitis
of surrounding tissues that might be indicative of adverse tissue responses to the hydrogel material
(Fig. 4.7). Despite some signs of acute inflammation in the first hours (less than 12 hours) after
the original procedure, irritation rapidly subsided. By the 24-hour mark, treatment and control eyes
were barely distinguishable. Evaluation at 48 hours after surgery revealed no hyperemia or
inflammation at the treatment site, which persisted until study termination.
Fig. 4.7: Visual evaluation of eye responses to surgical procedure and sealant placement.
(A-C) Photographic series of study (right) eye adnexa of a treatment group rabbit (R116) taken at:
(A) 6 hours, (B) 24 hours and (C) 48 hours after procedure. No evidence of inflammation or
exudate indicative of a local inflammatory response was seen at the lateral canthus, as might be
expected following the surgical procedure. This was true for all three time points. Note, all animals
wore Elizabethan collars to minimize effects of rubbing. (D and E) Visual inspection of the
53
conjunctiva and scleral surface at the laceration sites (blue arrows) for (D) a representative
treatment group rabbit (R116) and (E) a representative control group rabbit (R117) at 1-week after
procedure showed no distinguishable difference other than the presence of the hydrogel implant in
the treatment group eye.
Hydrogel efficacy in situ was evidenced by IOP restoration to normal ranges and negative
Seidel test, the standard clinical test for identifying ocular leakage/dehiscence. Mean baseline IOP
values showed no statistically significant difference between the study eye (OD) and fellow eye
(OS) of each animal (presurgical OD/OS between groups, P = 0.35), consistent with previous
reports (Fig. 4.8) (59). Immediately after incision creation, both treatment and control groups
showed significant decreases (>80%) in mean-normalized IOP. As restoration of IOP relied on
physiological production of aqueous humor, immediate improvement in IOP was not expected in
both groups.
Fig. 4.8: Representative baseline intraocular pressure (IOP) values showing no statistical
difference between eyes of the same animal or any circadian-induced variations; columns show
54
6.5 ± 0.2, 6.5 ± 0.2, 5.8 ± 0.2, 6.4 ± 0.3, 8.1 ± 0.4, 8.0 ± 0.4, 7.4 ± 0.5, and 8.6 ± 0.4 mmHg (n =
8 per group).
A noticeable IOP increase occurred about 12 to 24 hours after procedure in the treatment
group, and statistically significant improvement in IOP continued over the 72 hours after
placement of the hydrogel relative to the control (Fig. 4.9 and table 4.3). Normalized mean IOP
measurements across all time points beyond 24 hours after procedure showed statistically
significant improvement over control, with an even greater improvement noted beyond 48 hours
(P < 0.05). Normalized to the mean normal clinical IOP (15.5 mmHg), the lower threshold for
normal human IOP pressure (10 mmHg) becomes 65% (60), and unrepaired globes maintained a
low IOP no greater than 30% of normal. In stark contrast, hydrogel closure consequently raised
IOP to 90% of the minimum threshold for normal IOP and sustained IOPs almost twice the
magnitude of no treatment.
55
Fig. 4.9: Wald test comparison of mean IOP values of the treatment group versus no intervention,
after procedure, showed a statistically significant improvement in mean IOP with sealant placed
(*P < 0.05 and **P < 0.001).
Table 4.3: In vivo statistical analysis of average OD/OS for groups over time.
A statistical study of the IOP data was designed and performed to determine whether any measured
differences in IOP between the study and control groups was statistically significant. The study
was a (2:1, treatment/control) randomized and unblinded study (n = 18) to determine the effect of
using the hydrogel versus no intervention in an animal (rabbit) model of open globe injury. For
each animal at each measurement time, a normalized IOP measurement was calculated as
IOPOD/IOPOS. The normalized measurement was compared between treated and untreated
groups over the follow-up time using linear mixed effects models, allowing for correlated data
(arising from repeatedly measured data within animals) and flexible modeling of longitudinally-
collected data. Fixed effects included the treatment group, with inclusion of the pre-injury IOP as
a model covariate; a random effect was specified for an animal-specific intercept (animal-specific
IOP). An omnibus test of treatment effect over post-injury time was tested by the simple test of
the main effect of treatment. Time-specific treatment effects were compared between groups with
the addition of a day-by-treatment interaction term. Mean (Standard Error of the Mean, SEM)
outcome measurements were estimated and plotted for each day and treatment group. Table 4.3.
lists the collected and calculated values from the analysis. Note comparison of pre-surgery IOP
56
versus the post-surgical measures. Fig. 4.9. presents the same data in a graphical format, including
P-values.
4.3 N95BA5 biocompatibility up to 1 month after ocular
implantation
Histological analysis was performed at 48 hours, 1 week, and 4 weeks after hydrogel implantation
to investigate potential adverse reactions. Tissue cross sections with the laceration and hydrogel
placement sites marked (Fig. 4.10, black arrows) show the trauma site and hydrogel placement
location. A comprehensive analysis of tissue reaction over intervention time course was drafted by
a certified pathologist (table 4.4). Despite design intent for acute intervention, elongated exposure
allowed inspection of chronic inflammatory reactions, usually observed at 3 to 4 weeks after an
implantation (61).
57
Fig. 4.10: Series of histological cross sections prepared for control (left pairs) and treatment (right
pairs) in hematoxylin and eosin (H&E) and Masson’s trichrome stain for each of the study
endpoints (t = 48 hours, 1 week, and 4 week). Scale bars, 800 mm.
Table 4.4: Scleral tissue response at the hydrogel-sclera interface.
Histological analysis of eye tissues (n = 10) from the three control and treatment timepoints
confirmed natural healing response at the wound sites from both groups without evidence of
chronic inflammation at later points in the time course. Overall findings suggest minimal to no
toxicity or degenerative effects on the retina. Long term use of the plug resulted in the formation
of a fibrotic tissue encapsulation at the wound site. Any ocular perforation would ideally have been
58
addressed by the 1-week mark, and these long-term results were gathered for biocompatibility
assessment only.
At 48 hours, control group wound margins appeared clean with the presence of some acute
inflammatory cells (Fig. 4.10, left). Partial evulsion of peripheral retina was noted at the margins
(blue arrow) of the control group along with some inward epithelial migration (red arrow). At 1
week, the control eyes exhibited a typical acute inflammatory response, forming a porous fibrotic
layer bridging the lesion margins that matured from a positive Seidel test (confirmed leakage) to a
complete IOP-supportive barrier by week 4. At week 1, there was no evidence of inflammation or
infection in scleral tissues, but epithelial migration into the posterior chamber was observed.
Trichrome staining revealed the mature (week 4) barrier to be dense and primarily composed of
collagen fibers, a natural healing response. Small distributions of inflammatory markers were
found in the newly formed tissues, but there was no evidence of chronic inflammation.
Treatment group eyes exhibited similar clean laceration margins at 48 hours with some
retinal tissue evulsion; however, the opposing margins were noticeably separated (Fig. 4.10, right)
as a result of the hydrogel occlusion. Acute inflammatory cells at wound margins and slight inward
epithelial layer migration were consistent with acute traumatic injury. An immature encapsulation
layer formed over the wound by week 1, with no evidence of lesion closure or epithelial bridging,
providing a quantitative metric for the foreign body reaction (Fig. 4.11). The layer had a moderate
thickness (51.0 ± 15.7 mm) of 10 to 20 cell layers. At 4 weeks, the encapsulation matured into a
compact (13.4 ± 4.9 mm) tissue layer lining the margins of the lesion. Trichrome staining
confirmed this layer to be rich in collagen, indicating evidence of a mature and compact fistula
with a minimally adverse tissue response and no evidence of chronic inflammation forming around
the hydrogel plug. The sclera also created an immature encapsulation layer containing some
59
inflammatory cells, which separated the scleral tissue from the implant. Despite small quantities
of infiltrate observed in the scleral margins at week 1, there was no evidence of sustained or
excessive infiltrate, giant cell formation, or other chronic inflammation indicators at week 4.
Fig. 4.11: Increased magnification of one of the laceration margins for the treatment group
showing evolution of the tissue-hydrogel interface from acute inflammatory infiltrate to a mature,
compact fibrotic encapsulation layer at 4 weeks. Scale bars, 50 mm.
Treatment group retinas showed no evidence of detachment and showed no evidence of
neurotoxicity after 4 weeks of exposure (Fig. 4.12). Different approaches have been used
previously to assess neurotoxicity, including neurophysiological electroretinogram recordings and
histological study of retinal structure (62). Treatment group physical retinal evaluation showed no
signs of photoreceptor outer segment degeneration or other evidence of disorganization in the
laminar tissue structure (Fig. 4.12, bottom row). Organized ganglion cell layers, inner and outer
plexiform layers, and nuclear layers were also present, along with retinal photoreceptor outer
segments, all consistent with normally functioning retina. Any separation between the retina and
choroidal layers was confirmed as artifacts of the histological slide preparation, which is a common
observation (63). Even 30 days after implantation, the treatment group retinal tissues showed no
signs of degradation or detachment.
60
Fig. 4.12: Gross visualization (top row) and high-magnification (bottom row) evaluation of
treatment group retinas showing no evidence of trauma induced retinal detachment or hydrogel-
induced retinal neurotoxicity. Scale bars, 5 mm (top row) and 100 mm (bottom row).
The in vivo study provided preliminary insight into the hydrogel’s efficacy in occluding
open globe injuries. Whereas control groups approached 30% of normal pressure, treatment group
IOP reached as high as 60%, which was corroborated by negative Seidel tests. The hydrogel
formed a barrier capable of improving IOP during a 72-hour wait time for treatment.
Our preliminary biocompatibility assessment of N95BA5 suggests no significant adverse
effects. Most reported studies on PNIPAM biocompatibility for biomedical and cosmetic
applications support our findings (64-67); however, there has been one report of some observed
tissue inflammation (68) and another reporting IOP decrease (51). No such history of testing exists
for the hydrogel studied here, but PINIPAM copolymerized with BA has demonstrated basic cell
61
compatibility (52). Our examinations demonstrated no neurotoxicity, no retinal tissue degradation,
and no significant chronic inflammatory response after sustained exposure (30 days). Study
limitations have now been addressed in this paragraph, and after an exhaustive literature review,
we are not aware of any further material limitations. This preliminary demonstration of safety still
requires a larger study size to rule out effects on retinal function.
This study establishes foundational feasibility, safety, and efficacy of N95BA5 as a
reversible thermoresponsive hydrogel for temporary closure of scleral perforations. This tissue
repair technology sustained IOPs five times greater than the physiological range in bench testing
of maximum adhesion, prevented hypotony for 72 hours after scleral trauma, and demonstrated
sufficient biocompatibility for exposures longer than the intended use period. Furthermore, the
hydrogel’s adaptability to wound shape and unique deployment method make a unique and easy
alternative to conventional methods.
62
4.4 Materials and methods
4.4.1 Deployment tool development and use
To deploy the hydrogel in the eye, a novel tool was developed as follows: an 18-gauge intravenous
catheter tip trimmed to 5 to 8 mm was attached onto the Luer-Lock end of the 1-cm
3
syringe, which
was inserted through the opening of the 15-cm
3
syringe. The space between the external wall of
the 1-cm
3
syringe and the internal wall of the 15-cm
3
syringe was filled with ammonium nitrate
and water. The endothermic dissolution reaction for the cooling mechanism was calibrated by
using a pair of digital thermocouples to track the temperature in the jacketing chamber and inside
the hydrogel chamber when different quantities of water and ammonium nitrate were mixed
together. A working range of about 7 to 15 ml of water mixed with 6 to 8 g of ammonium nitrate
would result in a temperature-lowering profile meeting the performance criteria. For testing and
use, the tool components were first sterilized, and then, a sterile 1-cm
3
syringe loaded with
hydrogel was loaded into the jacket. The jacket was loaded with ammonium nitrate and capped
with a rubber vial cap. When ready for use, a separate syringe filled with sterile water and capped
with an 18-gauge needle was inserted through the rubber vial cap, and water was injected into the
jacket chamber. The tool was shaken a few times vigorously to initiate the reaction. After 60 s, the
gel was ready for application.
4.4.2 Hydrogel preparation for in vivo validation
The 30% N95BA5 hydrogel used for in vivo characterization was sterilized by first measuring a
prescribed quantity (1.0 g) of the dried powder form into a 10-cm
3
glass crimp vial container. Each
open and filled container was placed inside a Tyvek sterilization pouch with the associated stopper
and crimp top. The pouch was then sealed and ethylene oxide– sterilized. Still inside the sealed
63
and sterilized pouch, the stopper and crimp top were manipulated on top of the open container,
and the container was crimped shut inside the pouch. Once sealed, the pouch was opened, and the
jar was withdrawn, leaving a sealed crimp container with 0.5 g of sterile hydrogel powder. The
sterilized hydrogel powders were hydrated by injecting sterile water into the vial and performing
the hydration procedure described previously. Once hydrated, the containers were transferred to a
refrigerator for storage until use.
4.4.3 In vivo study design
An unblinded, two-arm randomized (2:1, treatment/control) study was conducted, in which all
animals received 3-mm full-thickness incisions in the sclera of the right eye (OS) about
3mmposterior to the limbus in the temporal superior quadrant to mimic a traumatic injury. The
fellow eye (OD) in all rabbits was left untouched as a control. Treatment group animals received
the hydrogel intervention. Control group animals received no intervention (current standard of
care). Pigmented New Zealand rabbits were used for this study because the eye dimensions closely
approximate the size of the human eye, thus mimicking approximate conditions for human use of
the system.
For treatment group rabbits, once the open globe injury was created, the hydrogel-loaded
syringe was prepared for use, as described previously. About 0.1 to 0.3 cm
3
of hydrogel was used
for each procedure. Control group rabbits received the same surgical procedure to the right eye but
received no intervention. At the end of each procedure, all rabbits in both groups received
(quantity) subcutaneous injections of buprenorphine with repeat injections at 12-hour intervals for
48 hours.
64
The research team regularly monitored each rabbit for the first 6 hours after the procedure.
They then performed checkups at 6-hour intervals through the first 12 hours, followed by 12-hour
interval checkups through the first week. At each 12-hour follow-up, rabbits were removed from
their cage, and their eyes were visually inspected by the research team for any evidence of adverse
tissue responses, such as swelling, inflammation, or bleeding from the sclera, conjunctiva, or other
surrounding tissues, with instructions to notify the ophthalmologists of the team of any adverse
events.
4.4.4 In vivo IOP measurements
IOP of both eyes of all animals was measured using a magnetically actuated veterinary rebound
tonometer (Tonovet). Baseline IOP values were established by measuring IOP of each eye twice
daily (a.m. and p.m.) for 5 days leading up to the procedure (Fig. 4.8). The Tonovet calculates an
average reading from six tonometric measurements taken in succession. Four successive readings
were taken on each eye; thus, 24 measures contributed to the averaged IOP for each eye. For all
tonometry measures, rabbits were removed from the cage and placed on an evaluation table for 2
min to allow the animal to relax. Stress from handling is known to artificially elevate blood
pressure and IOP values. The average of three IOP measurements was recorded at each time point
for each eye (both OS and OD) at regular intervals after the surgical procedure. Normalized IOP
values were reported by dividing IOPOS by IOPOD to reduce the impact of confounding systemic
effects (stress, infection, and medications). All surgical procedures were performed in the a.m. and
completed before noon, thus allowing IOP measures to resume in the late p.m. of the same day.
65
4.4.5 In vivo study endpoint
Study endpoints for the rabbits were set at 48 hours, 1week, and 4 weeks to evaluate the
progression of the tissue response at the implant site. Rabbits were not followed longer than 4
weeks because the intended use of the hydrogel will be for less than 30 days. Rabbits were
euthanized by first administering a heavy dose of ketamine/xylazine anesthesia, followed by
intravenous injection of a lethal dose of sodium pentobarbital via the auricular vein. Once
euthanized, a surgical procedure was performed to quickly enucleate and fix the study eyes (OS).
4.4.6 In vivo histology analysis
Enucleated eyes were fixed in Davidson’s solution to preserve structure of the total globe. Tissues
were sectioned and stained with either hematoxylin and eosin (H&E) or Masson’s trichrome stain
to evaluate local inflammatory response and characterize fibrosis. Retinal detachments were
assessed by histological evaluation of the posterior segments at study endpoints. Retinas were
analyzed for evidence of photoreceptor outer segment disorganization or complete degeneration,
indicative of retina separation from the choroidal vasculature and nutrient supply. In addition to
evaluation of the overall structure, tissues were sent to an outside pathology laboratory
(Comparative Bioscience Inc.) to be evaluated by a certified veterinary pathologist for evidence of
cytotoxic or neurotoxic effects of the hydrogel on surrounding tissues.
4.4.7 Statistical analysis
For each animal at each time point, a normalized IOP measurement was calculated as IOPOD/IOPOS
(Fig. 4.9). The normalized measurement was statistically compared between treated and untreated
groups using a generalized estimating equations model to account for correlated data arising from
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the repeated measures of IOP within an animal. The repeatedly measured normalized IOP was the
dependent variable; independent variables were treatment group and time (day, a.m./p.m.) of
measurement. The postsurgery treatment effect on IOP was tested over all postsurgical
assessments, as well as at each measurement time. An omnibus test of treatment effect over
postinjury time was tested by a score test of the main effect of treatment. Time-specific treatment
effects were compared between groups with the addition of a day-by-treatment interaction term;
treatment group differences were compared at each time point with a Wald test looking for a
statistically significant difference between the two groups (α = 0.05; two-sided). Mean (Standard
Error of the Mean, SEM) outcome measurements were estimated and plotted for each day and
treatment group. Eighteen animals randomized in a 2:1 fashion provided exceptional power (>95%
for the treatment effect throughout follow-up time) to evaluate the mean group differences
observed.
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CHAPTER 5
5.1 User feedback workshop with ophthalmologists and technicians
To validate the potential implementation of this technology in the clinic, two user feedback
workshops were organized to allow military clinical personnel, who might see ocular trauma
casualties, to use the TRS system in a benchtop model of ruptured globe injury (Fig. 5.1). Two
user feedback workshops that collectively captured end user feedback from 53 clinicians
(ophthalmologists, physicians, medics/ technicians, and researchers) with experience or career
interests in managing ocular trauma in combat casualties were organized. The objective of these
exercises was to assess the clinical relevance of the technology design and to capture any additional
design refinements.
Each participant was given a brief tutorial and then asked to treat an enucleated pig eye
with a full-thickness scleral laceration (0.5 to 2.0 cm in length). Each participant was given two
attempts to close the globe. After sealant placement, warm saline was infused into the eye via a
cannula to test integrity. Forty-three percent of the participants were able to successfully deploy
the hydrogel at the site to effectively occlude the ruptured globe on the first attempt. All of the
participants (100%) were able to successfully deploy the hydrogel at the site of the injury and close
the globe in two attempts.
A questionnaire was administered to the participants with write-in (table 5.1) and multiple
choice (table 5.2) questions. Of the responses collected, 69% of the respondents (n = 22) thought
the idea of the reversible occlusion for open globes was a good idea (table S5), with 59% thinking
it was great. Thirty-one percent found the system easy to use, with 25% noting that the system
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required minimal training. Perhaps most interesting, 94% of the responders (n = 30) said that they
could envision the system being used in the field or in managing combat casualties.
Fig. 5.1: TRS in the hands of professionals: In vitro application at Walter Reed Medical Center.
(A) 5 to 20 mm incisions were created in enucleated porcine eyes. (B) TRS Hydrogel was injected
through the wound site and allowed at least 2 minutes at 32°C to enter the solid state and set. (C)
The eye was then infused with warm saline to pressurize and test for leakage. The sealant was
pulled away from the laceration by its cap to examine the cohesiveness of the material. (D) All
injected hydrogel was removed in one piece without harm to the scleral tissue, indicating that
hydrogel cohesiveness allows reversible placement/removal of the sealant into the eye. (E and F)
The TRS also successfully sealed wounds with missing scleral tissue and stellate incisions, shown
with applied hydrogel before its phase change, to illustrate relative size and nature of traumas
tested.
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Table 5.1: Responses from freehand write-in section of user survey administered during clinical
user workshop.
Aggregation of answers written by military clinicians in response to the open questions (questions
without multiple choice answers to select from). The aggregated responses were reported as
absolute numbers (# Responses) and as a percentage with respect to the number of responders who
filled in this section of the questionnaire (% of Group).
Table 5.2: Responses from multiple choice section of user survey administered during clinical user
workshop.
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Testers were provided a series of multiple choice questions related to the design, function and
performance of the sealant system. Testers were asked to select a response from 1 to 5 where 1 =
strongly disagree, 2 = disagree, 3 = no feeling one way or the other, 4 = agree and 5 = strongly
agree. Answers 1 and 2 were aggregated as Disagree. Answers 4 and 5 were aggregated as Agree.
Responses were reported as a percentage of the total answers recorded for the specific question.
Clinical user feedback is a valuable element in developing new interventions and therapies,
and here, we specifically targeted military clinicians, who have encountered or who will encounter
ocular trauma under conditions where temporary intervention may be preferable over full
intervention. Although full intervention at first admission is preferred (8), some scenarios, such as
mass casualty events, may create scenarios where multiple ocular injuries may delay intervention
(7). The feedback captured from the military clinicians provided some good validation that the
ease of use and the relevance of the technology were both real.
5.2 Clinical user workshops to collect design feedback
Two user feedback workshops were organized to allow military clinical personnel, who might see
ocular trauma casualties, to validate the potential implementation of this technology in the clinic.
The tests were performed on a benchtop model of ruptured globe injury.
The first workshop was a small preliminary user feedback study at the 2017 meeting of the
Association of University Professors of Ophthalmology (AUPO) in January 2017. The second was
an expanded 3-day feedback study conducted in collaboration with the US Army Ophthalmology
Service at the Walter Reed Military Medical Center during the 2017 annual Tri-Service Ocular
Trauma Training Workshop in June. The first study at AUPO was a first introduction of the system
to 9 military ophthalmologists, which provided us valuable preliminary feedback on the system
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design and evaluation workshop approach. This information was used to plan and build out the
expanded study at Walter Reed. At the June workshop, a total of 44 clinicians and researchers
were invited to test their ability to close a model of open globe injury (enucleated porcine eye)
using the TRS system.
Without prior exposure to the technology, each tester was shown a video of how the system
works, then provided a demo injector tool to practice its operation before the test. Each participant
was given two attempts to close an open globe using the system. A full thickness scleral laceration
ranging from 5 to 20 mm was created on each pig eye for testing (Fig. 5.1). Once the hydrogel
adhesive was deployed on the laceration and allowed to set, the eye was infused with warm saline
to test the seal. Following the test, a written feedback questionnaire was provided to capture the
individual’s demographics and work experience; as well as their opinions on the system, its
components, and its envisioned use.
Of the 44 participants in the study, 84% (n = 37) were physicians, 36 of whom were
ophthalmologists. In addition to the physicians, 4 military EMTs (army medics and navy
corpsmen) also participated in the study, as well as 3 scientists/researchers. All participants were
part of the Tri-Service Ocular Trauma workshop, and therefore had some experience or interest in
ocular trauma management. Of the 36 participating ophthalmologists, 78% (n = 28) reported
having firsthand experience with managing open globes. Experience ranged from 1 per year to as
many as 24 open globes per year.
Analysis of the results found that 43% of the participants were able to successfully deploy
the hydrogel at the site to effectively occlude the ruptured globe on the first attempt. All of the
participants (100%) were able to successfully deploy the hydrogel at the site of the injury and close
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the globe in two attempts. Note, even the 7 non-physicians (4 medics/corpsmen and 3 scientists)
successfully occluded the penetrating injuries.
Questionnaire responses (tables 5.1 and 2) provided valuable insight into the existing
design and where additional changes may be helpful. Of the responses collected, 69% of the
respondents (n = 22) thought the idea of the reversible occlusion for open globes was good idea,
with 59% thinking it was great. Thirty-one percent found the system easy to use with 25% noting
that the system required minimal training. Perhaps most interesting, 94% of the responders (n =
30) said they could envision the system being used in the field or in managing combat casualties.
Collectively, the feedback from the clinicians provided good confirmation from a population of
relevant end users that the reversible sealant technology we have developed provides a pragmatic
solution to managing posterior segment globe ruptures. While some minor challenges and design
changes were identified, we believe this exercise provided valuable feedback from potential
clinical end users. These inputs are now being integrated into a final design with plans to begin
transition towards clinical safety evaluation.
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CHAPTER 6
6.1 HERA design for ocular trauma
The goal of the work presented here was to develop a removable, biocompatible adhesive to
perfectly seal ocular wounds, maintaining normal intraocular pressures, at physiological
temperatures. The approach presented here involves application of viscoelastic, liquid hydrogel at
sites of scleral tissue damage or loss, when raised to body temperature, it converts into a soft-solid
adhesive which can restore intraocular pressure. After gelation, the thermo-responsive adhesive
can be repositioned or removed without causing additional trauma, by exposure to cold water, as
a consequence of low adhesion to the ocular tissue at low temperatures (i.e. 5°-10°C). The HERA
would form a strong bond, yet reversible, across apposed wound edges, allowing normal healing
to occur below. It could replace surgical sutures for ocular incisional surgery or laceration repair.
This technology would be easy to use, requires little time, minimizes discomfort for patients and
eliminate the need for suture removal (Fig. 6.1). The HERA is exclusive to the presence of two
main components: (i) a thermo-responsive sealant (TRS) hydrogel that, owing to its mechanical
strength and desirable phase transition temperature, conforms to wound shape and (ii) a
commercially available additive with specific functional groups to achieve strong, rapid adhesion
against tissue.
Initially, the TRS hydrogels were formed by making aqueous solutions of poly (NIPAM-
co-BA) (N95BA5) leverages the unique LCST tunability of NIPAM by the inclusion of a relatively
more hydrophobic comonomer. Extensive characterization previously demonstrated that these
viscoelastic, injectable hydrogels reversibly transition to a soft-solid material when raised to body
temperature, exhibiting a complex viscosity of η* = 11,000 (Pa.s) (23). TRS was chosen as the
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Fig. 6.1: Engineered HERA for ocular trauma and mode of function. HERA application and
temperature-mediated removal from a scleral perforation.
primary formulation for preparation of additive-mediated adhesion hydrogels on account of its
high mechanical strength, functionality and biocompatibility (69). TRS combined with
commercially available additives showed reversible phase transition behavior, similar to the trend
observed for unmodified TRS. At low temperature, below LCST, hydrophilic interactions of
copolymer with water enable a translucent liquid state; as temperature increases above the
transition temperature, it forms white, soft-solid aggregates (Fig. 6.2). This transition also causes
a significant change in the viscoelastic strength and durability of the hydrogel.
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Fig. 6.2: Images of adhesive hydrogel temperature transition from translucent gel to opaque
aggregate when crossing lower critical solution temperature.
Several systems have been reported utilizing bioactive peptides to improve culture
response or adhesion properties of hydrogels. These studies range from PEG to Hyaluronic and
Acrylic Acid and commonly incorporate the peptides RGD, Tyr–Ile–Gly–Ser–Arg (YIGSR), Ile-
Lys-Val-Ala-Val (IKVAV), and AA. Results have included improved culture characteristics,
changed hydrogel material properties, and improved adhesiveness to tissue (70-73). One study
even formed a hydrogel from self-assembling peptide nanofibers containing the RGD motif that
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were then used to create a favorable environment for cell culture and dental tissue engineering
(74). Several of these studies have reported covalently attached peptides to NIPAM copolymers
for tissue culture or biomedical applications. One study created enzyme-cleavable crosslinking in
a temperature responsive tissue culture (75). Another used RGD-functionalized NIPAM
copolymer as a bioactive support structure for bone regeneration, validated in 3D tissue culture
(76). Other material studies have developed composite peptide-NIPAM copolymer gels exhibiting
increased storage modulus, but no change in LCST, mirroring our observations (77). Taken
together, these studies demonstrate the improved mechanical and bioactive properties offered by
peptide inclusion, but not their usefulness in macro-scale biomedical tissue adhesives.
Fig. 6.3: Phase diagram of hydrogel phase behavior demonstrating the role of polymer composition
in the LCST transition between hydrophilic chains and less soluble aggregates.
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To maintain the simplicity of the TRS system, the ideal additive was identified as one
which could increase adhesive strength through mixing with the hydrogel rather than further
chemical modification. The RGDS peptide was initially identified as a likely candidate for its cell-
adhesive and bioactive properties. Based on cell-adhesive capacity of RGDS, we hypothesized that
the same bioactivity that facilitates its interaction with integrin transmembrane receptors could
promote the formation of an adhesive interaction between the tissue and the hydrogel. To
investigate if RGDS peptide elicits a specific adhesive biological response to the tissue or only a
subset of the functional groups is needed to improve adhesion, a smaller molecule known as 3-
Guanidinopropionic acid (GPA) was also identified as a potential additive with similar functional
groups to the RGDS peptide (red) (Fig. 6.4). By altering the compositional ratios of these additives
to TRS, a range of thermo-sensitive hydrogels were prepared then characterized to optimize
transition temperature, adhesion strength, and viscoelastic properties. Three general formulations
include: unmodified TRS, RGDS-enhanced TRS, and GPA-enhanced TRS, each using TRS the
main formulation component.
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Fig. 6.4: Molecular structures of the Arg-Gly-Asp-Ser (RGDS) peptide and 3-guanidinopropionic
acid (GPA) additives selected for similarity of functional groups (red) and bioactive adhesive
potential.
6.2 Body temperature activated HERA with superior mechanical
properties
Additive introduction into the thermo-responsive hydrogel caused a dramatic increase in material
durability, a structural change with behavioral consequences. Viscoelastic behavior of additive-
enhanced hydrogels was characterized by rheological analysis, a well-established method for
determination of polymer chain networking and overall hydrogel rigidity. Thermo-responsive
changes in loss (G’’) modulus—viscosity—and storage (G’) modulus—strength—provided a
quantification of how hydrogels stored or absorbed deformation energy, illustrating increases in
gel networking. For these studies, tan δ, which is the ratio of the loss modulus over the storage
modulus (tan δ = G”/G’), was used as a measure of hydrogel tendency toward fluid-like (δ > 1) or
solid-like (δ < 1) properties. Initial viscoelastic characterization obtained by force-controlled
strain-amplitude sweeps at a range of temperatures established hydrogel linear viscoelastic region
where the gel properties are independent of imposed stress or strain levels. Rheological
experiments must be conducted in this linear region. Strain-dependent oscillatory rheology of the
GPA-enhanced gel [0.58% (w/w)] displayed an extremely broad linear viscoelastic region in
addition to network failure at high strains, indicating a wide processing regime and shear-thinning
behavior (Fig. 6.5). The storage modulus for GPA-enhanced hydrogel rapidly decreased above the
critical strain region (10%).
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Fig. 6.5: Strain-amplitude sweep of GPA-enhanced hydrogel [0.58% (w/w)] performed at fixed
temperature (32°C) isolated the hydrogel linear region for further analysis and demonstrated a
large capacity for deformation.
The optimized pure TRS exhibited high enough viscosity below the transition temperature
that allowed ease of application, but sufficient elasticity at body temperature to mechanically resist
intraocular pressure at the wound interface. Modulus measurements of additive-included
hydrogels obtained by oscillatory temperature ramp at fixed strain (0.5%) provided a better
understanding of additive-mediated changes in hydrogel conformation. For three different
temperatures (10°, 20°, and 30°C), the loss and storage moduli (G” and G’, respectively) of GPA-
included hydrogel [0.58% (w/w)] were measured at fixed strain (Fig. 6.6). Below the gelation
point, the gel viscosity was dominant over the gel elasticity. Above the phase transition region
(30°C), elasticity was relatively greater than viscosity, indicating a solid state. Comparison of
temperature-dependent moduli distributions of GPA and RGDS-enhanced hydrogels confirmed a
similar trend and phase transition to the unmodified TRS (Fig. 6.7 and 8).
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Fig. 6.6: Storage and loss moduli (G′ and G″, respectively) of GPA hydrogel [0.58% (w/w)] were
compared at temperatures spanning the transition region (10°, 20°, 30°C), showing a significant
increase in storage modulus and solid-like behavior at higher temperatures (n = 3).
Fig. 6.7: Loss & storage modulus of unmodified TRS as a function of temperature.
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Rheological measurements of a TRS sample (30% hydrogel) at continuously increasing
temperatures revealed the gradual increase of the storage modulus, indicating the temperature-
dependent formation of hydrophobic networks within the hydrogel. Below the LCST (10°C), a
dominant loss modulus indicated the dominance of fluid-like behavior. As the temperature
increased above the LCST (20°C), an increasing storage modulus begins to approximate the
magnitude of the loss modulus, indicating an increase in overall hydrogel rigidity. At sufficiently
high temperatures where significant dehydration can occur, the storage modulus dominates and
results in solid-like behavior (tan<1).
Fig. 6.8: Loss & storage modulus of RGDS-included hydrogel as a function of temperature.
Rheological characterization of low-concentration RGDS-enhanced hydrogels [0.58% (w/w)]
across continuously increasing temperatures shows parallels in material behavior to the
unmodified TRS. The peptide causes a slight elevation of both viscoelastic moduli at temperatures
lower than LCST (10°C). As temperature increases above LCST (20°C), both moduli increase
relative to TRS, but the same significant gain in storage modulus magnitude relative to loss
modulus magnitude can be observed. At sufficiently high temperatures(30°C), this trend continues
to the point where storage modulus overcomes loss modulus and the RGDS-enhanced hydrogel
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behaves more like a solid than a fluid. This mirroring of unmodified TRS behavior illustrates that
the relative gelation mechanism and gel material properties are preserved despite conformational
changes.
Furthermore, the dynamic nature of the non-covalent interactions of GPA and RGDS with
TRS gel was investigated. Below the transition temperature (16°C), the loss modulus increased,
as predicted, with addition of both GPA and RGDS [2.91% (w/w)] to the unmodified TRS. A
change in the loss modulus (ΔG” ≈ 10 Pa) from a relatively low (42 Pa) to high (52 Pa), indicating
stronger attractive intermolecular forces, is a beneficial property for easier and more controllable
injection. At low temperature, viscous-dominant behavior was verified by large values of tan δ
(G’’/G’ ≥ 5), a consequence of low elasticity (Fig. 6.9). Even blow the transition temperature, the
storage modulus values of pure TRS hydrogel showed an increase by a factor of two, by addition
of GPA [2.91% (w/w)], while the magnitude of increase became smaller as GPA was replaced
with RGDS (Fig. 6.10).
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Fig. 6.9: Results of fixed-temperature (16°C) oscillatory rheometry confirmed an increase in loss
modulus (G″) and an overall decrease in tan, illustrating improved overall material durability and
a relative increase in solid-like behavior.
Fig. 6.10: Storage modulus of additive-included hydrogels at 16°C (n = 30).
The 30% N95BA5 copolymer hydrogel (TRS) samples enhanced with additives (GPA and RGDS)
were compared with an unmodified hydrogel by rheological analysis. Oscillation experiments
performed at a stable temperature of 16°C, below their LCST, revealed that the elastic moduli of
additive hydrogels were increased. A 2.91% (w/w) concentrated GPA-enhanced hydrogel
achieved a 100% increase in the storage modulus, indicating an improved resilience to deformation
and higher cross-linking. Even before temperature-mediated aggregation, both additives promote
the formation of a more cohesive structure.
TRS gels mixed with additives were roughly 1.2 times more viscous at low temperature.
It is important to investigate additive contribution to TRS mechanical properties at body
temperature to predict gel barrier strength and rigidity. We screened the elastic modulus of both
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additive hydrogels above the transition temperature (32°C) for differences from the unmodified
TRS. We observed that incorporation of small amount of GPA into the TRS formulation caused a
dramatic increase in the elastic strength of the gel, indicating the formation of significant polymer-
polymer interactions (Fig. 6.11). The G’ values of the GPA-enhanced hydrogel were also
significantly greater than both the unmodified and RGDS-enhanced hydrogels by factors of more
than 62 and 21 times respectively. Further, above the LCST (32°C), GPA-included samples still
exhibited the greatest storage modulus values, compared to the unmodified and RGDS-enhanced
hydrogels (Fig. 6.12). This relationship demonstrated that both additives strengthen the internal
hydrogel networks, but GPA addition yields a much more cohesive structure. The dynamic moduli
of HERA gels showed that additives, more specifically GPA, are able to serve as non-covalent
crosslinkers between the polymer chains, while the polymer chains may bridge many different
particles, enabling stronger hydrogel formation.
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Fig. 6.11: Fixed-temperature oscillatory rheometry above LCST (32°C) demonstrated dramatic
increases in elastic moduli (G′) of additive-enhanced samples [2.91% (w/w)] with GPA dwarfing
other samples.
Fig. 6. 12: Loss modulus of TRS as a function of GPA and RGDS at 32°C (n = 30).
30% copolymer hydrogel samples were compared by oscillatory rheometry at a fixed temperature
of 32°C, above LCST and after gelation. 2.91% (w/w) GPA hydrogels exhibited over 6000% and
2000% increases in loss modulus relative to non-additive and 2.91% RGDS (w/w) hydrogels
respectively. This increase indicates that more force is required to deform the sample at high
temperatures, but more specifically that the sample has a significant increase in overall viscosity.
This increase can be explained by an increase in mobile non-interacting regions within the
hydrogel that are less likely to store energy than networked polymer chains. (*) Relative to both
samples, GPA was capable of a remarkable increase in hydrogel resilience.
Temperature-dependent viscosity measurements were then performed to investigate the
GPA [0.58% (w/w)] effect on the hydrogel’s conformational change (Fig. 6.13). GPA-enhanced
hydrogel formulation followed a similar trend to unmodified TRS, as reported previously.
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Complex viscosity, η*, remained nearly constant below the gelation point (≈ 10 Pa.s). A sharp
increase of viscosity value was observed at higher temperatures (at and above phase transition
temperature), which was attributed to transformation from a more soluble coil conformation below
the LCST to a compact globule formation. As GPA particles were dissolved into the TRS gel, the
exponential increase in viscosity became steeper and reached a higher apex (Fig. 6.13). These
observations support our hypothesis that gel dynamic moduli would scale directly with addition of
the additive components, bridging the gap between polymer chains.
Fig. 6.13: Oscillatory temperature sweep of additive samples followed the same complex viscosity
trend as pure TRS, but exhibited steeper increase with a higher apex even with low concentration
[0.58% GPA (w/w)].
The complex viscosity profiles of additive-enhanced formulations [2.91% (w/w)] GPA and
RGDS] were then evaluated to compare the strengthening effect of these materials as a function
of temperature (table 6.1). We observed that hydrogel viscosity depends on the temperature and
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all studied samples showed similar viscosity below LCST (10°C). As we further increase the
temperature to 16°C, both RGDS and GPA-enhanced hydrogels generated higher η* values of 3.1
and 3.5 (Pa.s), respectively (Fig. 6.14).
Table 6.1: Comparison of additive [2.91% (w/w)] effects on TRS complex viscosity by fixed-
temperature rheology across LCST showed similar values below the transition region, but additive-
mediated increase beyond LCST.
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Fig. 6.14: Below LCST, complex viscosity values were comparable across all studied samples, but
indicating an additive mediated increase (n = 3).
At body temperature, both TRS hydrogels combined with RGDS and GPA, demonstrated
a large improvement over the unmodified TRS. Although the mechanical strength of all hydrogels
showed the same trend, GPA-enhanced formulation generated stronger polymeric networks (Fig.
6.15). Above the phase transition temperature, replacement of RGDS additive with GPA caused
the η* values to increase by a factor of up to 24 times. We observed that all samples reached their
maximum viscosity at eye temperature (32°C). The η* values also confirmed that GPA can
improve the mechanical strength of non-covalent hydrogels. Among the three samples studied,
GPA had the greatest complex viscosity value of about 419,000 centipoise (cP) (419 Pa·s),
suggesting a strong, yet soft, thermo-responsive hydrogel.
Fig. 6.15: Above LCST, complex viscosity values were dramatically elevated for GPA-enhanced
formulations.
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6.3 The nanostructure of crosslinked HERA network at body
temperature
The dramatic increase in viscoelastic moduli of additive-enhanced samples at body temperature,
indicated an enhancement in overall hydrogel cohesiveness when solidified. To investigate how
the additives impact, at the nano-scale, on the TRS molecular structure [5% (w/w) N95BA5], we
employed dynamic light scattering (DLS) over temperature ranges that included phase transition
temperature. DLS particle size and population distribution measurements offer a better
understanding of macromolecule structure and molecular assembly. The intensity distribution of
TRS in the presence of GPA and RGDS were screened at additive concentrations from 0.58% to
29% (w/w), allowing comparison across additive and temperature gradients. At the sol-to-gel
phase transition region (12°C), we observed two peaks for both unmodified TRS (None) and GPA-
enhanced [2.91% (w/w)] hydrogels (Fig. 6.16).
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Fig. 6.16: Dynamic light scattering comparison of GPA-enhanced [2.91% (w/w)] N95BA5 hydrogel
to a sample without additive indicated two particle size populations below LCST, with an increase
in hydrodynamic radius of large particles for the additive-enhanced hydrogel.
The first peak represented smaller particles (≈ 5 nm) with relatively lower intensity values
(compared to the second peak) for samples with and without the GPA additive. The other
population (the second peak) corresponded to larger particle size and a higher percentage of
intensity. This peak shows that N95BA5 molecules are highly aggregated in the solution. The size
distribution curve for GPA-included hydrogel showed a similar trend to pure TRS solution (Fig.
6.16). However, GPA addition caused a dramatic change in hydrodynamic radius of the second
peak (RH), (ΔRH = 618 nm), which resulted in aggregate size 3.6 times larger than in unmodified
TRS. The resulting cohesiveness was observed in the increased storage modulus and viscosity.
Further temperature increase to 18°C caused 98% of the copolymer to form particles with a larger
hydrodynamic radius (637 nm) in presence of GPA and smaller hydrodynamic radius (428 nm) in
additive-free hydrogel, a consequence of more favorable polymer-polymer interactions (Fig. 6.17).
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Fig. 6.17: Temperature increase above transition indicated the transition to one dominant particle
size and a preserved elevation of GPA particle sizes.
Again, since these measurements were made at fixed temperature, the change in the particle size
arose from the interaction between the polymer and GPA molecules.
The hydrophobic collapse of TRS in the absence and presence of GPA was also monitored
by DLS while heating the sample. An automatic temperature sweep of different concentrations of
GPA [0.58% and 2.91% (w/w)] allowed observation of both the size and scattering intensity as a
function of temperature (Fig. 6.18). As we reported previously, even in the absence of GPA, TRS
is capable of exhibiting heat induced aggregation. At phase transition temperature, TRS showed a
clear phase transition upon heating, going from soluble to partially insoluble in the aqueous
solution. At temperatures below the LCST (≤ 10°C), for all studied samples, we observed
aggregates with small size (≤ 15 nm), which can be attributed to the polymer in coil conformation,
as primary hydrogen bonding to water molecules. It is notable that RH rapidly increased around
12°C, indicating the onset of turbidity and polymer gelation onset (table 6.2).
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Fig. 6.18: Comparison of GPA concentration effects on particle size revealed similar particle sizes
below LCST, but a concentration-dependent increase in particle size at higher temperatures (n =
3).
Table 6.2: Temperature-dependent variation of the hydrodynamic radii of GPA-included N95BA5
solution [5% (w/v)].
Comparison of average hydrodynamic radius values of TRS solution in the absence and presence
of different concentrations of GPA, revealed an increase in aggregate size by addition of GPA. We
noticed that the onset of RH values increment was shifted toward lower temperature in the presence
of all GPA contained samples.
RH values further increased as GPA concentration rose [up to 2.91% (w/w)], increasing
polymer-polymer association, while further inclusion of GPA caused a decrease in aggregate radii
at higher temperature (table 6.3). One can clearly see that, a low concentration of GPA [0.58%
(w/w)] is sufficient to influence the particle size by rupturing the hydrogen bonds between polymer
and water molecules. Even so, the phase transition value did not change with concentration of
GPA, the gelation onset remained nearly constant around 12°C in the most concentrated GPA-
enhanced hydrogel [29% (w/w)] (Fig. 6.19). Comparison of aggregates hydrodynamic radius
among GPA solutions indicates the that higher additive concentrations facilitate aggregation.
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Table 6.3: Hydrodynamic radii of GPA-enhanced N95BA5 solutions demonstrates additive-
mediated aggregation.
Nanometric copolymer aggregates in the solution state start to grow at the temperature of gelation-
onset, per the conventional behavior of the polymer.The onset temperature (12˚C) did not change
as GPA concentration increased, despite the formation of larger aggregates. In the presence of
GPA, N95BA5 copolymer chains start to aggregate forms larger particles, and with increasing
temperature the size of the aggregates further increased due to the hydrophobic interaction of
PNIPAM chains. The RH values of GPA-enhanced samples followed the same trend as RGDS.
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Fig. 6.19: Normalized average particle size for GPA hydrogels across temperature transition.
DLS measurements produced average particle size in addition to population spectra, allowing
direct comparison of LCST across several GPA-enhanced hydrogels. Despite the significant
impact of GPA on the structure of the polymer networks and the overall particle sizes within the
hydrogel, there was no discernible impact on hydrogel LCST. It was observed that even significant
concentrations of GPA up to 29% (w/w) had negligible effect on hydrogel LCST, shifting it by at
most 1°C. That realization indicated that although the additive promoted the formation of larger
particle size, it did not appreciably alter the thermodynamics that govern the copolymer transition
from coil to networked globule.
Next, we investigated the variation of particle size with RGDS concentration as a function
of temperature. RGDS-enhanced hydrogel intensity spectra followed a similar trend, two
populations of aggregates in the phase transition region, the first on the nano range and the second
in the range of 200-500 nm (12°C) (Fig. 6.20). While both the unmodified hydrogel and the RGDS
hydrogel shared the same low-size peak (≈ 4.5 nm), the second fraction (large population) revealed
an increase in size in the presence of RGDS (table 6.4).
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Fig. 6.20: Comparison of RGDS [2.91% (w/w)] particle dynamics to an unmodified N95BA5
showed similar improved particle size below the transition temperature.
Comparison of the particle size distribution of aggregates in two different additive-
enhanced hydrogels [2.91% (w/w) GPA and RGDS] at gelation onset point (12°C) differentiated
their aggregation behavior (Fig. 6.21). We clearly observed that copolymer molecules undergo a
larger intermolecular aggregation, mediated by additives. The quantification of particle size
indicated that the presence of GPA and RGDS in solution increase the size of aggregates by more
than 2-fold (table 6.4). Additive mediated scaling of particle sizes indicated improved molecular
bonding between polymer chains.
Table 6.4: At gelation point (12°C), additives increased the size of larger particles and the amount
in that larger population.
With temperature increase (18°C) of RGDS-enhanced hydrogel, the hydrodynamic radius
and intensity of the high aggregation peak region increased to 541 nm and 97%, respectively (Fig.
6.22). The same monodisperse high-intensity peaks (≈ 98%) of highly aggregated particles were
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observed across all studied formulations (table 6.5). This result confirmed that the intermolecular
cross-linking dominated molecular behavior above the gelation point.
Fig. 6.21: Comparison of particle distributions showed a striking increase in particle size of GPA-
enhanced samples relative to the RGDS-included formulations.
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Fig. 6.22: DLS spectra of the intensity distribution graph of copolymer solution with RGDS at
18°C.
The particle size distributions of unmodified TRS solution and a RGDS-enhanced hydrogel were
obtained by DLS across a range of temperatures (2°C - 20°C). The intensity of small particles
decreases gradually, replaced by the dominant population of larger size. The size distribution for
the [2.91% (w/w)] RGDS-enhances solution shows a similar trend, but exhibits accelerated
aggregation behavior as in pure N95BA5 solution.
Table 6.5: Hydrodynamic radius and intensity distribution profiles of the TRS hydrogel in the
presence of RGDS and GPA [2.91% (w/w)] at 18˚C.
In addition, we studied the variation of the RH of TRS aggregates as function of temperature
and RGDS concentration. For brevity, the size profiles of three concentrations of RGDS [0.58%,
1.16%, and 2.91% (w/w)] were included across a range of increasing temperatures (Fig. 6.23 and
table 6.6). Consistent with the concentration-dependent trend observed in GPA-enhanced
hydrogels, an increase in the concentration of RGDS increased the size of the aggregates,
indicating an additive-induced aggregation phenomenon. During the aggregation process, particle
size was conserved across all samples when approaching the gelation point. As temperature
increased, however, there was a progressive increase in RH values towards a maximum size (384
98
nm) at 20°C, for pure TRS solution. Inclusion of only 2.9% (w/w) of RGDS, shifted that maximum
particle size to 485 nm (Fig. 6.24). Similar to GPA, the largest RGDS effects were observed for
2.91% (w/w) concentrated sample, beyond the gelation point (≥ 16°C) where aggregation scaled
significantly with increase in temperature (table 6.7). Despite these effects, at low concentration
of RGDS, the phase transition temperature values remained nearly unchanged. Higher
concentrations [≥ 4.7% (w/w)] of RGDS did elicit slight decreases of the onset of phase transition
temperature, however (Fig. 6.25). For instance, the value of gelation onset point shifted from 12°C
(additive-free) to 11°C in the presence of 4.7% (w/w) RGDS.
Fig. 6.23: Concentration-dependent of RGDS effect on particle size distribution within dilute
hydrogel.
An overlay of RGDS-enhanced hydrogels [0.58%, 1.16%, and 2.91% (w/w)] with increased
concentration illustrated the significant additive-mediated increase in particle size beyond the
LCST. Temperatures below the gelation onset region (<12°C) demonstrated relative indifference
to increased additive concentration. This difference suggests that the peptide additive can
99
encourage hydrophobic interactions more effectively at higher temperatures, perhaps because
RGDS requires the presence of preliminary hydrophobicity before it can act as a catalyst for it.
Table 6.6: Variation of hydrodynamic radii of the aggregates as a function of temperature for
N95BA5 solutions with different concentrations of RGDS.
RH values of N95BA5 [5% (w/w)] aggregates as a function of temperature and RGDS concentration
revealed an increase in aggregate size by addition of RGDS. At high temperatures (20˚C),
increasing RGDS corresponded to larger particle size. Due to the significant gain in turbidity of
the copolymer solution, DLS measurements could not be performed above 20°C. Below the LCST,
the RH of RGDS samples increased by small amount. Above the transition temperature, 2.91%
(w/w) RGDS increased polymer’s particle size around 100 nm.
100
Fig. 6.24: RGDS peptide caused a significant increase in average particle size above LCST (n =
3).
Table 6.7: RGDS increases temperature-mediated aggregation in N95BA5 solutions [5% (w/v)].
DLS evaluation of a series of RGDS-enhanced solutions [0.58%, 1.16%, 2.91%, 4.74%, 9.49%,
and 23.7% (w/w)] offered clear picture of the additive’s effects on the RH values. At temperatures
below the LCST (2 - 10°C), we observed RH was small (9.5 – 47.6 nm) for various concentrations
of RGDS, which can be
attributed to the polymer’s
coil conformation.
Hydrodynamic radius rapidly
increases around 12 - 18°C.
Moreover, the increase of the
size occurs in a quite wide
temperature range (≈ 5°C) due
to polymer starts globule state
from coil conformation,
revealing the phase transition
of copolymer occurs in this
region.
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Fig. 6.25: Normalized average particle size for RGDS-enhanced hydrogels across various RGDS
concentrations as a function of temperature.
Average particle size was determined by DLS for peptide-enhanced hydrogels of increasing
concentration. While low concentrations caused negligible change in the LCST, normalized
comparison indicated that concentrations of 4.7% (w/w) and above caused a drastic shift of LCST
by at least 2°C. As the transition temperature of the NIPAM copolymers is largely determined by
the thermodynamic favorability of different hydrogen bonding interactions with water, it can be
assumed that the presence of RGDS reaches a critical concentration at which in not only increases
particle size at high temperatures, but also influences improves the hydrophobicity of smaller
particles.
To further evaluate the relative impacts of each additive, we calculated their molar ratios
to copolymer and examined differences in particle behavior. It was verified that values of RH are
almost the same for similar molar ratio (Fig. 6.26). For instance, 1.16% (w/w) GPA (molar ratio
Fig. 6.26: Temperature dependence of hydrodynamic radii in additive (GPA/RGDS)-enhanced
copolymer solution.
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In the presence of RGDS and GPA, RH increase was observed with enhancement of RGDS and
GPA concentration. Interestingly, marked increase of RH values in additive-included samples were
found around phase transition temperature of additive-free hydrogel, over the whole range of the
present concentrations of RGDS and GPA. Moreover, the phase transition temperature values of
additive-included hydrogels are almost as same as additive-free aqueous N95BA5 solution, despite
additive effects on particle size.
of 2.66) shows the same RH value as 2.91% (w/w) RGDS (molar ratio of 2.01) at 20°C. In the
current study for all TRS-based samples (the additive concentration was varied from 0 to 2.91%
(w/w)), the turbidity began to abruptly increase at the same temperature (12°C), indicating that
these concentrations of additives had no influence on the phase transition temperature (Fig. 6.27).
To elucidate further the effect of additives, at fixed temperature (20°C), on the aggregate size of
TRS, the highest concentrations of RGDS and GPA [2.91% (w/w)] were also compared (Fig. 6.28).
The results revealed increased GPA aggregate size relative to RGDS and pure TRS, highlighting
the importance of peptide-inspired GPA as a unique and easy alternative to biological molecules.
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Fig. 6.27: Additive concentration independent transition temperature value revealed by normalized
particle size.
Average particle size was determined across a range of temperatures for all hydrogel DLS
experiments. When normalized for impactful yet low concentrations of additives [≤ 2.91% (w/w)],
negligible differences in LCST can be observed based on concentration or additive type. Despite
their significant impact on particle size, additives at the studied concentrations did not appreciably
alter the thermodynamics of the base copolymer hydrogel. As a result, low concentrations of both
investigated additives showed no impact the phase transition of the dilute TRS solution.
Fig. 6.28: Comparison of average particle size at body temperature (32°C) demonstrated larger
polymer networks in GPA [2.91% (w/w)] hydrogels compared to RGDS samples (n = 3).
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6.4 Materials and methods
6.4.1 Hydrogel solution preparation
The required amount of N95BA5 was weighed for various concentrations of hydrogel aqueous
solutions, from 5 to 30% (w/w). The powder was added directly to the sterile water. The vial
containing the polymer suspension was then processed with a Misonix Sonicator 3000 using a cup
horn high-intensity ultrasonic water bath at the maximum power setting (10) due to sample
viscosity. A circulating temperature control water bath held at 2°C prevented sample heating due
to prolonged horn activity. The sample was sonicated until a transparent clear hydrogel was
obtained. The required sonication time ranged from1 to 30 hours and depended on the hydrogel
concentration and molecular weight of the homopolymers/copolymers.
6.4.2 Rheological Analysis
A TA instruments Discovery Hybrid-2 (DHR-2) Rheometer was used to measure the viscoelastic
properties of the hydrogels. 0.2 ml of 0°C hydrogel was injected onto the lower temperature-
controlled plate at 5°C and the rotational head lowered to a testing distance of 500 µm. A solvent
trap with di water was then placed over the rotational head to prevent solvent evaporation. Samples
were incubated for 5 minutes at the test temperature prior to initiation. Strain-amplitude and
frequency sweeps, were performed at fixed temperatures (6° and 32°C) with a fixed frequency of
15 rad/s and fixed strain of 0.5% respectively. Following isolation of the critical strain regions
both below and above LCST, oscillatory tests were performed with at 15 rad/s and 0.5% strain.
Axial force control at 0 N was also used to monitor for changes in volume and prevent unwanted
loss of contact or sample shearing during the phase transition. A 5 °C/minute temperature ramp
rate was used for 5° to 50°C in an attempt to prevent solvent loss at high temperatures.
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6.4.3 Dynamic Light Scattering
Particle size of 5% (w/w) N95BA5 hydrogels was obtained by automated DynaPro Plate Reader II
(Wyatt Technology). 80 µl was introduced to square glass cuvette through a micropipette. Mineral
oil (20 µl) was also added to prevent water evaporation. The cuvette was then placed in the
temperature-controlled sample chamber and maintained at test temperature for 30 minutes to
ensure thermos dynamic equilibrium. This procedure repeated for all test temperatures (2° to
20°C). DLS measures movement of particles due to Brownian motion and determines particle size
by the Stokes-Einstein equation:
D = kT / 6πηRH
where k is the Boltzmann’s constant, T is the absolute temperature, h is the viscosity, and D is the
diffusion coefficient. All data were obtained and process in Dynamics software.
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CHAPTER 7
7.1 In vitro adhesion function of HERA hydrogels
As observed previously, RGDS-inspired GPA served as functional nanofillers to reinforce the
hydrogel network, resulting in a more elastic hydrogel, which was characterized by rheological
tests. DLS measurements also confirmed that abundant functional groups of GPA, interact with
copolymer network, serving as a cross-linking site to generate a highly entangled network structure
and increase the aggregates size. However, as an on-demand ocular adhesive, the hydrogel should
exhibit not only biocompatibility and strong mechanical properties but also cell/tissue
adhesiveness for cell attachment and tissue integration. The tissue adhesiveness on additive-
mediated hydrogels were investigated in vitro using a uniaxial pull-off adhesion test. The
experiments were performed on surface-modified glass substrate, textured steel and then scleral
tissue to assess the adhesiveness of hydrogel to porcine sclera tissue. Differences in substrate
surface chemistry and texture were selected for initial examination of the hydrogel adhesion
mechanisms. The environmental control and sample conditioning times were chosen to remove
variability between substrate temperatures and to encourage peak bond formation by low
temperature spreading. Maximum separation force served as an indicator of adhesion bond
strength (Fig. 7.1).
In a preliminary examination on porcine sclera tissue, additive-enhanced TRS
demonstrated significantly increased adhesion strength. As we reported previously, cyanoacrylate
was selected over fibrin-, albumin-, and polyethylene glycol–based adhesives as a positive control
because of its well-documented superiority in maintaining IOP and uniaxial adhesion strength. The
improvement scaled with additive concentration and provided the first demonstration of enhanced
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tissue adhesiveness (Fig. 7.2). Both RGDS and GPA elicited this adhesive behavior, but the highest
GPA concentration [2.91% (w/w)] outperformed unmodified TRS by 1.7 N, almost doubling its
bond strength. Even the lowest GPA concentration [0.58% (w/w), 1.33 M/M] matched a hydrogel
with a higher RGDS concentration [2.91% (w/w), 2.01 M/M]. Maximum additive concentrations
tested yielded a 2.8 N and 3.7 N adhesion force for RGDS and GPA respectively. Notably, the
2.91% (w/w) concentration of GPA elicited an average adhesive force within 1.5 N of the
cyanoacrylate positive control.
Fig. 7.1: Schematic of uniaxial adhesion tests as performed on a variety of biological and non-
biological substrates.
Mucoadhesion between an adhesive and tissue surface can best be explained by one or a
combination of four processes: adsorption, diffusion, electronic, and wetting. Adsorption theory
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assumes that primary and secondary bonds form at the interface of the adhesive and the
glycoproteins of the mucin layer, with similar functional groups to RGDS and GPA being bonded
by a cell-surface integrin. Dismissing bonds, diffusion theory assumes that the interpenetration
and entanglement of adhesive polymer chains and mucin chains are responsible for the bond.
While adsorption and diffusion assume chemical adhesive mechanisms, electronic theory
attributes bond strengthening to an attractive charge bilayer formed by electron transfer or charge
attraction—for instance between cationic and anionic substances. Finally, wetting theory uses a
mechanical justification of adhesive bonds, assuming that the bioadhesive enters surface
irregularities and hardens to form a physical anchor (78, 79). Following the initial axial tests on
sclera, any of these explanations could be used to justify the GPA and RGDS-mediated adhesion.
Fig. 7.2: Hydrogel adhesion forces on scleral squares with negative (tissue-tissue) and positive
(cyanoacrylate) controls.
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To further understand the mechanism of enhanced adhesion force, we investigated
adhesion of additive-enhanced hydrogels to non-biological substrates. Highest concentration
additive hydrogels [2.91% (w/w)] were compared on unmodified, hydrophobic, negatively
charged and positively charged glass substrates to examine bonding preferences. These modified
glass substrates were created by surface chemistry through attachment of octyl, carboxylic acid
and amine groups to unmodified glass through a silanization process, verified by contact angle
measurements (Fig. 7.3, 7.4, and 7.5).
Fig. 7.3: Contact angle of DI water on surface-modified glass slides (n = 5).
To verify that the desired octyl, amine, and carboxylic acid groups were successfully bonded by
silanization to the surfaces of the sandblasted glass slides, three phase contact angles were
performed with a series of solutions. In comparison of DI water contact angle, all modified slides
exhibited a significant increase in contact angle relative to unmodified sandblasted glass. This
increase was caused by the alkyl chains attached to the modified surfaces during the silanization
reaction. DI water successfully verified that the octyl slides were hydrophobic as intended, but did
not verify the functional groups of the charged surfaces.
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Fig. 7.4: Contact angle of HCl solution (pH = 0) on surface-modified glass slides (n = 5).
In order to verify the functional groups and charges of the amine and carboxilic acid functionalized
slides, the contact angle of an acidic solution on unmodified slides and the two charged slides was
determined. By applying an acidic solution, the carboxilic acid was gauranteed to be protonated
while the amine formed an alkylammonium. This fundamental difference of a charged or neutral
surface resulted in a lower contact angle for the neutral carboxylic acid than the ammonium slide.
This result was unexpected as more interaction and lower contact angle were expected between
the water molecules and the charged ammonium groups. Nevertheless, a significant difference in
surface chemistry was confirmed.
In contrast to the improved adhesion observed on tissue, no difference in adhesion force
was observed for samples with and without RGDS and GPA additives on the unmodified glass
(Fig. 7.6 and 7.7). A slight increase in adhesion force was observed for both additive-included
hydrogels on the hydrophobic (octyl-immobilized) glass, but only GPA caused adhesion forces
greater than one standard deviation from the mean force of an unmodified hydrogel (Fig. 7.8).
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Fig. 7.5: Contact angle of NaOH solution (pH = 14) on surface-modified glass slides (n = 5).
To further confirm the differences in functionalization of the amine and carboxilic acid slides,
contact angle was measured for a NaOH solution with pH of 10. Under basic conditions, the amine
was prevente from protonation while the carboxilic acid was deprotonated to a carboxylate. This
charged vs neutral disparity was used to further demonstrate the opposing behvaiors of the charged
slides. Under acidic conditions, the neutral slide (carboxylic acid) had the lower contact angle; the
same trend was observed under basic conditions. The decreased contact angle for the amine-
modified slides illustrated the opposing modifications of these two glass slides and provided a
reassuring verification of silanization success.
The slight improvements in adhesion force on the hydrophobic slides may have been
caused by the penetration of the octyl groups into the hydrogel of vice-versa, allowing diffusive
adhesion. Negative and positive charged substrates caused opposite effects on additive-included
hydrogels. GPA-enhanced hydrogel had an average force of 12.6 N on a negatively charged
substrate while unmodified and RGDS-enhance hydrogels had forces of 19.9 N and 21.2 N
respectively (Fig. 7.9).
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Fig. 7.6: Demonstration of adhesion test raw data with similar maximal force for bond breakage
[GPA sample].
Fig. 7.7: Comparison of hydrogel adhesion strength on unmodified sandblasted glass squares.
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Fig. 7.8: Comparison of hydrogel adhesion strength on hydrophobic (octyl) modified sandblasted
glass squares.
Fig. 7.9: Comparison of hydrogel adhesion strength on negatively charged (carboxylic acid)
modified sandblasted glass squares.
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In contrast, RGDS-enhanced hydrogel had an average force of 17.1 N on positively charged
substrate while unmodified and GPA-included hydrogels had adhesion forces of 19.4 N and 20.9
N (Fig. 7.10). Notably, while decreases in adhesion were observed for charged substrates, neither
caused significant increase in adhesion relative to the unmodified substrate.
Fig. 7.10: Comparison of hydrogel adhesion strength on positively charged (amide) modified
sandblasted glass squares.
Cohesive improvements in hydrogel were apparent in a comparison of adhesion to glass
substrates at eye temperature (32°C). Unmodified and additive-enhanced hydrogels were applied
at 10˚C to as a drop and allowed to transition to solid state, cyanoacrylate was then used to fix the
top of the gel to the upper plate of a uniaxial adhesion tester (Fig. 7.11). Upon separation, two
distinct behaviors were observed. The pure TRS hydrogel showed gel-breaking, cohesive failure
(A). In contrast, a TRS gel with GPA [2.91% (w/w)] applied under the same conditions completely
115
separated from the lower plate, demonstrating full adhesive failure (B). This result confirmed that
additive-free hydrogel has weaker polymer networks than GPA-enhanced gel, tearing and
ultimately separating between the two plates.
Fig. 7.11: Comparison of adhesive and cohesive properties of the pure and modified TRS.
Furthermore, in consideration of tissue surface irregularities, a textured steel substrate was
also used to examine whether surface wetting might play a role in the additive-mediated adhesion
mechanism (Fig. 7.12). While the adhesion force of unmodified hydrogel and RGDS-enhanced
hydrogel were both around 28 N, the GPA-enhanced hydrogel reached 40 N (Fig. 7.13). This
significant increase in adhesion force indicated that the GPA-mediated adhesion mechanism relies
in part on physical anchoring in surface irregularities. Moreover, despite the adhesion increase
observed for the RGDS-enhanced hydrogel on sclera, no significant difference from the
unmodified TRS was observed on these substrates. Either the adhesion mechanism of RGDS does
not rely on surface wetting, or it forms a weaker bond during the wetting process. This difference
in RGDS and GPA behavior on this substrate also implicates other forms of interaction on the
tissue surface to explain the improved bonding of RGDS on sclera.
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Fig. 7.12: 0.1 mL hydrogel adhesion to 2.5 cm textured steel square (n = 5).
Fig. 7.13: Comparison of hydrogel adhesion on highly textured steel substrates.
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These results offered some mechanistic insight, providing several justifications for
additive-mediated adhesion. Both RGDS and GPA have functional groups that are reactive and
common in the biological environment. As a result, intermolecular forces between charged groups
are likely to play a role in the increased tissue-hydrogel bond. The adhesive equivalence of
unmodified and additive-included hydrogels on a glass substrate without any organic or charged
elements clearly illustrates this bioactive role, especially when compared to the doubling of
adhesive strength on tissue. The slight improvements in adhesion force on the hydrophobic slides
may have been caused by the penetration of the octyl groups into the hydrogel, allowing diffusive
adhesion. The negative and positive charged substrates also caused opposite effects on additive-
enhanced hydrogels, potentially signaling a preferential interaction with the carboxylate or
ammonia ion. Importantly, neither of the additive-enhanced hydrogels experienced significantly
elevated adhesion on these charged substrates relative to the unmodified substrate. The result
largely ruled out the formation of an electronic bilayer between the hydrogel and the slides,
supporting the conclusion that the electronic theory was not active in adhesion mechanism. Finally,
the drastic increase in GPA adhesion force on a textured steel surface confirmed the role of surface
wetting in the overall tissue adhesive mechanism. Differences in RGDS and GPA behavior on this
substrate also implicated other forms of interaction on the tissue surface to explain the improved
bonding of RGDS on sclera. While undetermined, the additive-mediated adhesion mechanism for
these adhesives caused measurable change in hydrogel behavior. Overall, these results indicated
that the enhanced hydrogels were not universally adhesive and that substrate characteristics have
a significant role in tissue adhesion-mediated. This result proved our hypothesis that engineering
a thermo-responsive hydrogel with an appropriate peptide-inspired material, GPA, results in
improved tissue adhesiveness.
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7.2 Ex vivo testing of HERA as an ocular adhesive
To investigate the potential of HERA gel to effectively seal ocular defects, ex vivo experiments
were performed on pig sclera. Because of the strong, yet non-covalent, mechanical properties,
preferable particle size, and high tissue adhesion of HERA noted during in vitro tests, the ex vivo
experiments were performed using HERA as the ocular adhesive alone, without the use of sutures.
Additive-enhanced TRS formulations with 0.58%, 1.16% and 2.91% (w/w) GPA and RGDS
concentrations were used for the ex vivo experiments.
Ex vivo porcine sclera incision models were used to test the sealing properties of HERA in
the presence of extensive vitreous. A benchtop model of the eye was fabricated as a temperature
and pressure-controlled chamber enclosed at one opening by a replaceable segment of porcine
sclera (Fig. 7.14). Intraocular pressure was measured by an electronic pressure sensor, used to
detect the pressure created by heated saline flowing through the device. Following introduction of
a scleral square (2-cm × 2-cm), eye temperature saline (32°C) was run through the device to
simulate human physiology. A circle incision (≈ 2 mm) was created in the scleral surface, then
hydrogels (5°C) were applied over the point of incision area to study the ex vivo sealing properties
of HERA hydrogels. Following restoration of the ocular chamber, pressure gradually increased via
controlled injector system. Failure of the HERA wound closure system was documented by saline
leakage from the scleral wound and sudden IOP decrease. The adhesive strength of the HERA gel
applied on sclera surface was compared to that of pure TRS sample (Fig. 7.15). Relative to an
unmodified TRS, additive-enhanced hydrogels withstood up to 6-fold higher maximum IOP. This
increase in adhesive efficacy scaled with an increase in additive concentration. Notably, the highest
concentration of GPA hydrogel [2.91% (w/w)] withstood IOP of up to 96 mmHg, more than 5
119
times the bond of pure TRS. Cyanoacrylate positive control exceeded the pressure output of the
flow system.
Fig. 7.14: Schematic of intraocular pressure test using an ex vivo cadaveric pig eye
Increased additive-hydrogel adhesion to the scleral surface in an incision model followed
the same patterns observed in mechanical, particle structure, and uniaxial adhesion analyses.
Addition of the peptide-inspired additives directly increased hydrogel cohesiveness and IOP tests
further confirmed their ability to reinforce tissue adhesion strength. Use over external wound
surfaces further indicated the usefulness of these hydrogels as they would allow scleral recovery
and re-epithelialization beneath the gel layer, followed by easy removal upon cold water exposure.
Notably, the GPA-enhanced samples consistently outperformed RGDS samples of an equivalent
mass percent. Its commercial availability as a nutrition supplement and its greater performance
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Fig. 7.15: Ex vivo adhesive performance using pressure-controlled explanted cadaveric pig eye.
make GPA-enhanced hydrogel the ideal HERA candidate. These data confirm our hypothesis that
the highly elastic HERA hydrogel is capable of sealing severe scleral incision in the absence of
additional staples or sutures in an ex vivo animal model. Further pre-clinical large animal
experiments are warranted to evaluate the long-term fate of HERA implants.
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7.3 Materials and methods
7.3.1 Uniaxial adhesion tests on scleral tissue
Fresh (harvested within 24 hours) porcine eyes (from Sierra for Medical Science) were dissected
into 1-cm squares. These squares were then fixed to the temperature-controlled base plate and
vertical control arm of a TA instruments Discovery Hybrid-2 (DHR-2) rheometer with double
sided mounting tape. Cyanoacrylate was used to improve the bond between the tissue and adhesive
tape. An insulating solvent trap was placed around the plates and the tissue substrates incubated
for 5 minutes at 10°C. 0.02 ml of sample was then applied as a drop to the center of the bottom
substrate and top plate was lowered into contact. A controlled force of 1N was then applied for 5
minutes at 10°C followed by 5 minutes at 32°C. The plates were then separated at a rate of
2mm/minute until full adhesive failure was achieved. Adhesive force was noted as the maximum
force applied by the vertical arm to maintain the separation rate.
7.3.2 Uniaxial adhesion tests on glass substrate
Sandblasted lass slides (untreated, hydrophobic, positive charged and negative charges) were fixed
in 1cm squares to double-sided mounting tape with cyanoacrylate. These squares were then fixed
to the temperature-controlled base plate and vertical control arm of a TA instruments Discovery
Hybrid-2 (DHR-2) rheometer. Substrates were placed in contact and an insulating solvent trap
placed around them while incubated at 10°C for 5 minutes. 0°C sample (0.02 ml) was then applied
as a drop to the bottom substrate and top plate lowered. A controlled force of 1N was then applied
for 5 minutes at 10°C followed by 5 minutes at 32°C. The plates were then separated at 2mm.min
until full adhesive failure. Adhesive force was noted as the maximum force applied by the vertical
arm in order to maintain separation rate.
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7.3.3 Modified glass slide preparation
Glass slides with hydrophobic and charge surfaces were prepared by batch surface
functionalization. Positive slides, with an amine functional layer, were prepared by hydrolyzation
of ozone-treated fully frosted glass slides (Fischer Scientific) in ethanol, water, acetic acid solution
and 3-aminopropyltriethoxy-silane. Slides left in solution for 10 minutes prior to rinsing with
ethanol and drying for 10 minutes in a 120°C oven. Negative slides, with a carboxylic acid
functional layer, were prepared by exposure of amine-functionalized slides to 2 mg/ml solution of
succinic anhydride in DMSO for 16 hours. Slides were then rinsed with ethanol and dried in a
120°C oven. Hydrophobic slides, functionalized with an octyl group, were also prepared by
silanization.
7.3.4 Contact angle tests
One drop (~ 0.05 ml) was applied by syringe onto a glass substrate on the raised platform of a
Tantec CAM F-1 goniometer. The lowest point of the arc was aligned with the measurement vertex
and the ruler arm rotated to determine contact angle value. The half-angle method was used and
repeated 5 times on different substrates to ensure accuracy and account for variability in modified
surfaces.
7.3.5 Cyanoacrylate adhesion tests on glass substrate
Hydrogel samples (0.02 ml) were applied to the center of a glass substrate and allowed 1 minute
to equilibrate at 32°C. Cyanoacrylate (0.01 ml) was then applied directly to the top of the hydrogel
and a top glass substrate lowered to anchor the hydrogel to the top plate with a strong bond.
123
Cyanoacrylate was given 5 minutes to cure at fixed temperature. Plates were then separated at the
normal 1 mm/min rate for uniaxial tests.
7.3.6 Ex vivo IOP measurements
Fresh (harvested within 24 hours) porcine eyes (from Sierra for Medical Science) were dissected
into a 2-cm × 2-cm square shape with an open circle (0.1 cm diameter) at the center. Scleral
segments were then fixed in a specialized testing apparatus such that an internal chamber was
partially enclosed by an exposed scleral opening. An infusion line carried temperature-controlled
saline into the apparatus while a constant force applicator applied variable pressure to the enclosed
fluid. Internal pressure was monitored by a (insert name) pressure sensor with accuracy of (insert
number). Once the temperature of the scleral segment matched the internal temperature (32°C), a
radial puncture or scleral incision was then made to the exposed sclera. The puncture was then
sealed by application of TRS or additive-enhanced hydrogel. Continuous pressure monitoring was
then used to determine the pressure of sealant failure.
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REFERENCES AND NOTES
1. G. McGwin, Jr., T. A. Hall, A. Xie, C. Owsley, Trends in Eye Injury in the United States,
1992-2001. Investigative Ophthalmology & Visual Science 47, 521 (2006).
2. F. Kuhn, D. J. Pieramici, ebrary Inc., D. J. P. Ferenc Kuhn, Ed. (Thieme,, New York,
2002), pp. xxviii, 468 p. ill.
3. D. L. Cruvinel Isaac, V. C. Ghanem, M. A. Nascimento, M. Torigoe, N. Kara-José,
Prognostic factors in open globe injuries. Ophthalmologica 217, 431-435 (2003).
4. A. A. Honeycutt et al., in Using Survey Data to Study Disability: Results from the
National Health Survey on Disability. pp. 207-228.
5. R. J. Blanch, R. A. H. Scott, Military ocular injury: presentation, assessment and
management. Journal of the Royal Army Medical Corps 155, 279-284 (2009).
6. G. A. Byrnes, in Ophthalmic Care of the Combat Casualty, A. B. Thach, Ed. (Office of
the Surgeon General at TMM Publications, Washington, DC, 2003), chap. 13, pp. 511.
7. Y. Yonekawa et al., Ocular blast injuries in mass-casualty incidents: the marathon
bombing in Boston, Massachusetts, and the fertilizer plant explosion in West, Texas.
Ophthalmology 121, 1670-1676.e1671 (2014).
8. F. G. La Piana, T. H. Mader, in Ophthalmic Care of the Combat Casualty, A. B. Thach,
Ed. (The Office of the Surgeon General at TMM Publications, Washington, DC, 2003),
chap. 2, pp. 511.
9. G. Koranyi, S. Seregard, E. D. Kopp, Cut and paste: a no suture, small incision approach
to pterygium surgery. Brit J Ophthalmol 88, 911-914 (2004).
10. R. C. Hall, A. J. Logan, A. P. Wells, Comparison of fibrin glue with sutures for
pterygium excision surgery with conjunctival autografts. Clinical & Experimental
Ophthalmology 37, 584-589 (2009).
11. V. B.J, E. M.J, Cyanoacrylate glue for corneal perforations: a description of a surgical
technique and a review of the literature. Clinical & Experimental Ophthalmology
28, 437-442 (2000).
12. M. Forseth, K. O'Grady, D. M. Toriumi, The current status of cyanoacrylate and fibrin
tissue adhesives. J Long Term Eff Med Implants 2, 221-233 (1992).
13. G. H. Chen, A. S. Hoffman, Graft-Copolymers That Exhibit Temperature-Induced Phase-
Transitions over a Wide-Range of Ph. Nature 373, 49-52 (1995).
14. A. S. Hoffman, Environmentally Sensitive Polymers and Hydrogels - Smart Biomaterials.
Mrs Bull 16, 42-46 (1991).
15. M. A. Cole, N. H. Voelcker, H. Thissen, H. J. Griesser, Stimuli-responsive interfaces and
systems for the control of protein–surface and cell–surface interactions. Biomaterials 30,
1827-1850 (2009).
16. A. Halperin, M. Kröger, F. M. Winnik, Poly(N‐isopropylacrylamide) Phase Diagrams:
Fifty Years of Research. 54, 15342-15367 (2015).
17. J. L. Rushbrook, G. White, L. Kidger, P. Marsh, T. F. O. Taggart, The antibacterial effect
of 2-octyl cyanoacrylate (Dermabond®) skin adhesive. Journal of Infection Prevention
15, 236-239 (2014).
18. H. K. Chenault et al., Sealing and Healing of Clear Corneal Incisions with an Improved
Dextran Aldehyde-PEG Amine Tissue Adhesive. Current Eye Research 36, 997-1004
(2011).
125
19. C. Batman et al., A comparative study of tissue glue and vicryl suture for conjunctival
and scleral closure in conventional 20-gauge vitrectomy. Eye 23, 1382 (2008).
20. . (Ocular Therapeutix, Online., 2018), vol. 2018.
21. H. S. Uy, K. R. Kenyon, Surgical outcomes after application of a liquid adhesive ocular
bandage to clear corneal incisions during cataract surgery. Journal of Cataract &
Refractive Surgery 39, 1668-1674 (2013).
22. W. Bethke, in A look at the ways ophthalmologists use sealants and glues, and where the
newly approved ReSure sealant fits in. (Review of Ophthalmology, 2014).
23. Z. M. O. Rzaev, S. Dinçer, E. Pişkin, Functional copolymers of N-isopropylacrylamide
for bioengineering applications. Progress in Polymer Science 32, 534-595 (2007).
24. Y. Zhang, S. Furyk, D. E. Bergbreiter, P. S. Cremer, Specific ion effects on the water
solubility of macromolecules: PNIPAM and the Hofmeister series. Journal of the
American Chemical Society 127, 14505 (2005).
25. M. C. M. Costa, S. M. C. Silva, F. E. Antunes, Adjusting the low critical solution
temperature of poly(N-isopropyl acrylamide) solutions by salts, ionic surfactants and
solvents: A rheological study. Journal of Molecular Liquids 210, 113-118 (2015).
26. D. Dhara, P. R. Chatterji, Phase Transition in Linear and Cross-Linked Poly(N-
Isopropylacrylamide) in Water: Effect of Various Types of Additives. Journal of
Macromolecular Science, Part C 40, 51-68 (2000).
27. H. Du, X. Qian, Molecular dynamics simulations of PNIPAM-co-PEGMA copolymer
hydrophilic to hydrophobic transition in NaCl solution. Journal of Polymer Science, Part
B: Polymer Physics 49, 1112-1122 (2011).
28. C. M. Burba, S. M. Carter, K. J. Meyer, C. V. Rice, Salt effects on poly(N-
isopropylacrylamide) phase transition thermodynamics from NMR spectroscopy. The
journal of physical chemistry. B 112, 10399 (2008).
29. Z. H. Farooqi, H. U. Khan, S. M. Shah, M. Siddiq, Stability of poly(N-
isopropylacrylamide-co-acrylic acid) polymer microgels under various conditions of
temperature, pH and salt concentration. Arabian Journal of Chemistry 10, 329-335
(2017).
30. C. Hofmann, M. Schönhoff, Do additives shift the LCST of poly (N-
isopropylacrylamide) by solvent quality changes or by direct interactions? Colloid &
Polymer Science 287, 1369-1376 (2009).
31. A. Halperin, M. Kröger, Thermoresponsive cell culture substrates based on PNIPAM
brushes functionalized with adhesion peptides: theoretical considerations of mechanism
and design. Langmuir : the ACS journal of surfaces and colloids 28, 16623-16637
(2012).
32. C. J. Wilson, R. E. Clegg, D. I. Leavesley, M. J. Pearcy, Mediation of biomaterial-cell
interactions by adsorbed proteins: a review. Tissue engineering 11, 1-18 (2005).
33. F. Mannello, G. A. Tonti, Concise Review: No Breakthroughs for Human Mesenchymal
and Embryonic Stem Cell Culture: Conditioned Medium, Feeder Layer, or Feeder‐Free;
Medium with Fetal Calf Serum, Human Serum, or Enriched Plasma; Serum‐Free, Serum
Replacement Nonconditioned Medium, or Ad Hoc Formula? All That Glitters Is Not
Gold. STEM CELLS 25, 1603-1609 (2007).
34. D. S. Hwang, S. B. Sim, H. J. Cha, Cell adhesion biomaterial based on mussel adhesive
protein fused with RGD peptide. Biomaterials 28, 4039-4046 (2007).
126
35. U. Hersel, C. Dahmen, H. Kessler, RGD modified polymers: biomaterials for stimulated
cell adhesion and beyond. Biomaterials 24, 4385-4415 (2003).
36. E. Mauri, A. Sacchetti, F. Rossi, The Synthesis of RGD-functionalized Hydrogels as a
Tool for Therapeutic Applications. Journal of visualized experiments : JoVE, (2016).
37. S. D. Larsen et al., Synthesis and biological activity of analogues of the
antidiabetic/antiobesity agent 3-guanidinopropionic acid: discovery of a novel
aminoguanidinoacetic acid antidiabetic agent. Journal of medicinal chemistry 44, 1217-
1230 (2001).
38. B. L. Baumgarner, A. M. Nagle, M. R. Quinn, A. E. Farmer, S. T. Kinsey, Dietary
supplementation of β-guanidinopropionic acid (βGPA) reduces whole-body and skeletal
muscle growth in young CD-1 mice. Molecular and Cellular Biochemistry 403, 277-285
(2015).
39. E. Ruoslahti, M. D. Pierschbacher, Cell attachment activity of fibronectin can be
duplicated by small synthetic fragments of the molecule. Nature 309, 30-33 (1984).
40. A. Gutowska et al., Heparin release from thermosensitive polymer coatings: in vivo
studies. Journal of Biomedical Materials Research 29, 811-821 (1995).
41. P. N. Wahjudi et al., Improvement of metal and tissue adhesion on surface-modified
parylene C. Journal of biomedical materials research. Part A 89, 206 (2009).
42. S. Shekhar, M. Mukherjee, A. K. Sen, Swelling, thermal and mechanical properties of
NIPAM-based terpolymeric hydrogel. Polym Bull 73, 125-145 (2016).
43. L. T. Allen et al., Interaction of soft condensed materials with living cells:
Phenotype/transcriptome correlations for the hydrophobic effect. P Natl Acad Sci USA
100, 6331-6336 (2003).
44. W. F. Lee, Y. C. Yeh, Studies on preparation and properties of NIPAAm/hydrophobic
monomer copolymeric hydrogels. Eur Polym J 41, 2488-2495 (2005).
45. E. Velzenberger, K. El Kirat, G. Legeay, M. D. Nagel, I. Pezron, Characterization of
biomaterials polar interactions in physiological conditions using liquid-liquid contact
angle measurements Relation to fibronectin adsorption. Colloid Surface B 68, 238-244
(2009).
46. B. Brugger, J. Vermant, W. Richtering, Interfacial layers of stimuli-responsive poly-(N-
isopropylacrylamide-co-methacrylicacid) (PNIPAM-co-MAA) microgels characterized
by interfacial rheology and compression isotherms. Physical chemistry chemical physics :
PCCP 12, 14573-14578 (2010).
47. K. Kinbara et al., High-water-content mouldable hydrogels by mixing clay and a
dendritic molecular binder. Nature 463, 339-343 (2010).
48. A. C. Kumar, H. B. Bohidar, A. K. Mishra, The effect of sodium cholate aggregates on
thermoreversible gelation of PNIPAM. Colloids and Surfaces B: Biointerfaces 70, 60-67
(2009).
49. A. C. Kumar, H. Erothu, H. B. Bohidar, A. K. Mishra, Bile-Salt-Induced Aggregation of
Poly(N-isopropylacrylamide) and Lowering of the Lower Critical Solution Temperature
in Aqueous Solutions. The journal of physical chemistry. B 115, 433 (2011).
50. C. Monteux et al., Shear Surface Rheology of Poly(N-isopropylacrylamide) Adsorbed
Layers at the Air−Water Interface. Macromolecules 39, 3408-3414 (2006).
51. L. Zou et al., Intraocular pressure changes: an important determinant of the
biocompatibility of intravitreous implants. PloS one 6, e28720 (2011).
127
52. N. Y. Becerra, B. L. López, L. M. Restrepo, Thermosensitive behavior in cell culture
media and cytocompatibility of a novel copolymer: poly(N-isopropylacrylamide-co-
butylacrylate). Journal of Materials Science: Materials in Medicine 24, 1043-1052
(2013).
53. M. Banitt, J. B. Malta, H. K. Soong, D. C. Musch, S. I. Mian, Wound integrity of clear
corneal incisions closed with fibrin and N-butyl-2-cyanoacrylate adhesives. Curr Eye Res
34, 706-710 (2009).
54. K. A. Vakalopoulos et al., Mechanical strength and rheological properties of tissue
adhesives with regard to colorectal anastomosis: an ex vivo study. Ann Surg 261, 323-
331 (2015).
55. R. Shaikh, T. R. Raj Singh, M. J. Garland, A. D. Woolfson, R. F. Donnelly,
Mucoadhesive drug delivery systems. J Pharm Bioallied Sci 3, 89-100 (2011).
56. B. Natalia, A. Henry, L. Betty, R. L. Marina, R. Roberto, Probing poly(N-
isopropylacrylamide-co-butylacrylate)/cell interactions by atomic force microscopy. J
Biomed Mater Res A 103, 145-153 (2015).
57. P. Reddy, P. Venkatesu, Ionic Liquid Modifies the Lower Critical Solution Temperature
(LCST) of Poly(N-isopropylacrylamide) in Aqueous Solution. Journal of Physical
Chemistry B 115, 4752-4757 (2011).
58. S. Kaja et al., Evaluation of tensile strength of tissue adhesives and sutures for clear
corneal incisions using porcine and bovine eyes, with a novel standardized testing
platform. Clinical ophthalmology (Auckland, N.Z.) 6, 305-309 (2012).
59. J. W. McLaren, R. F. Brubaker, J. S. FitzSimon, Continuous measurement of intraocular
pressure in rabbits by telemetry. Invest Ophthalmol Vis Sci 37, 966-975 (1996).
60. F. H. Adler, R. A. Moses, W. M. Hart, Adler's physiology of the eye: clinical application.
(Mosby, 1987).
61. J. M. Anderson, A. Rodriguez, D. T. Chang, Foreign body reaction to biomaterials. Semin
Immunol 20, 86-100 (2008).
62. D. W. Herr, in Neurotoxicology: Approaches and Methods, L. W. Chang, W. Slikker,
Eds. (Elsevier Science, San Diego, CA, 1995), chap. 9, pp. 208-221.
63. D. A. Nayagam et al., Techniques for processing eyes implanted with a retinal prosthesis
for localized histopathological analysis. J Vis Exp, (2013).
64. R. Mentens, Comparison of fibrin glue and sutures for conjunctival closure in pars plana
vitrectomy. Am J Ophthalmol 144, 128-131 (2007).
65. M. Rahimi et al., Synthesis and Characterization of Thermo-Sensitive Nanoparticles for
Drug Delivery Applications. Journal of biomedical nanotechnology 4, 482-490 (2008).
66. Z. Cui, B. H. Lee, C. Pauken, B. L. Vernon, Degradation, cytotoxicity, and
biocompatibility of NIPAAm‐based thermosensitive, injectable, and bioresorbable
polymer hydrogels. Journal of Biomedical Materials Research Part A 98A, 159-166
(2011).
67. in International journal of toxicology. (United States, 2005), vol. 24 Suppl 2, pp. 21-50.
68. R. Zhu et al., PNIPAM hydrogel induces skeletal muscle inflammation response. RSC
Adv 5, 2823-2829 (2015).
69. N. Bayat et al., A reversible thermoresponsive sealant for temporary closure of ocular
trauma. Science Translational Medicine 9, (2017).
128
70. S. P. Zustiak, R. Durbal, J. B. Leach, Influence of cell-adhesive peptide ligands on
poly(ethylene glycol) hydrogel physical, mechanical and transport properties. Acta
Biomaterialia 6, 3404-3414 (2010).
71. F. Yang et al., The effect of incorporating RGD adhesive peptide in polyethylene glycol
diacrylate hydrogel on osteogenesis of bone marrow stromal cells. Biomaterials 26,
5991-5998 (2005).
72. C. E. Brubaker, P. B. Messersmith, Enzymatically degradable mussel-inspired adhesive
hydrogel. Biomacromolecules 12, 4326 (2011).
73. Y. Hao, A. B. Zerdoum, A. J. Stuffer, A. K. Rajasekaran, X. Jia, Biomimetic Hydrogels
Incorporating Polymeric Cell-Adhesive Peptide To Promote the 3D Assembly of
Tumoroids. Biomacromolecules 17, 3750 (2016).
74. K. M. Galler, J. D. Hartgerink, A. C. Cavender, G. Schmalz, R. N. Souza, A Customized
Self-Assembling Peptide Hydrogel for Dental Pulp Tissue Engineering. Tissue
Engineering Part A 18, 176-184 (2012).
75. S. Kim, E. H. Chung, M. Gilbert, K. E. Healy, Synthetic MMP‐13 degradable ECMs
based on poly(N‐isopropylacrylamide‐ co ‐acrylic acid) semi‐interpenetrating polymer
networks. I. Degradation and cell migration. Journal of Biomedical Materials Research
Part A 75, 73-88 (2005).
76. S. Morochnik et al., A thermoresponsive, citrate-based macromolecule for bone
regenerative engineering. Journal of biomedical materials research. Part A, (2018).
77. A. Maslovskis et al., Self-assembling peptide/thermoresponsive polymer composite
hydrogels: effect of peptide-polymer interactions on hydrogel properties. Langmuir : the
ACS journal of surfaces and colloids 30, 10471 (2014).
78. R. Shaikh, T. Raj Singh, M. Garland, A. Woolfson, R. Donnelly, Mucoadhesive drug
delivery systems. Journal of Pharmacy and Bioallied Sciences 3, 89-100 (2011).
79. N. Becerra, L. M. Restrepo, B. L. López, Synthesis and Characterization of a
Biocompatible Copolymer to be Used as Cell Culture Support. Macromolecular
Symposia 258, 30-37 (2007).
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Bayat, Niki
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Core Title
Reversible thermoresponsive materials for temporary closure of ocular trauma
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Viterbi School of Engineering
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Doctor of Philosophy
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Chemical Engineering
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07/27/2018
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ocular trauma
polymer
reversible materials
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thermoresponsive
tissue adhesive