Biomimetic Hybrid Hydrogels for Cartilage Tissue Engineering
Anna Maria Ganios
Department of Polymer Science (9871:497)
April 22, 2011
1
INTRODUCTION
Osteoarthritis (OA) is a prevailing degenerative cartilage disease that affects over 20
million individuals in the United States. It is estimated that 67 million Americans will suffer
from osteoarthritis by 2030, and 25 million of these individuals will be subjected to severe OA
(1). Cartilage damage or loss leads to friction between bones, which results in decreased joint
mobility and joint pain (2). Currently a cure does not exist for OA, and the avascular and aneural
nature of cartilage limit the treatments available to treat OA because of its poor regenerative
ability (3). Current interventions concentrate on alleviating pain and improving joint mobility.
Several surgical procedures like microfracture, autologous osteochondral grafts, and autologous
chondrocyte implantation are currently used to treat cartilage injuries or defects. While used
widely, these therapies have significant disadvantages, including the production of fibrous
cartilage that degrades over time.
Contemporary tissue engineering approaches to repair and ultimately replace damaged
tissue by generating functional tissue ex vivo using polymer-based scaffold materials (4). The
tissue engineering field has moved away from cell monolayers to a three-dimensional culture
environment in order to provide an optimal environment for cellular response (5). Biomimetic
scaffolds should resemble native tissue organization and possess mechanical properties similar to
natural materials. The scaffold should provide sufficient porosity for nutrient/waste flux and an
appropriate surface for cell attachment, proliferation, and cellular differentiation to occur (6, 7).
The cellular interaction with the scaffold material influences the rate and quality of tissue
formation. For example, how well cells are able to align and attach, as well as protein
adsorption, are influenced by the surface chemistry of the material. Therefore, hydrophobic
polymers generally do not provide the optimal environment for cell and material interaction (8).
The development of hydrophilic photopolymerized hydrogels, under physiologic conditions, has
provided a biocompatible scaffold for tissue engineering (9, 10).
Hydrogels are three – dimensional, hydrophilic crosslinked polymers that are able to
rapidly convert from a liquid monomer or macromonomer to a polymeric networks through
photopolymerization (10). Poly-(ethylene)-glycol) (PEG) is the most commonly utilized
polymer to create synthetic hydrogels (7). The photopolymerization reaction is driven by
2
photoinitiatiors that produce free radicals when exposed to visible or UV light. A photon from
the light source dissociates the photoinitiator; specifically, the photoinitiatiors are cleaved at the
C-C, C-Cl, C-O, or C-S bonds to form high-energy radicals. These radicals then induce the
polymerization of the monomer or macromer solution (9, 10).
Scheme 1: The polymerization of the PEGDM macromonomer is initiated using 365 nm light which generates
radical formation in the IRGACURE 2959. This particular initiating system is attractive in biological systems
in that both radical products are typically incorporated in the network minimizing cell toxicity.
By varying the molecular mass of the macromonomer or by varying the mass percent of
macromonomer in solutions, the network cross-link density can be easily manipulated (11). The
cross-link density is proportional to the gel modulus and inversely proportional to the swelling
(4). The transport properties and cell behavior are affected by the former and latter respectively.
Hydrogels are able to support cell survival and cell proliferation since they allow for a high
permeability of nutrients and water soluble metabolites.
Cellular interactions with the extra cellular matrix (ECM) are essential for the native
cartilage viability. The ECM is composed of proteoglycans, glycosaminoglycans, glycoproteins,
elastin, and collagen fibres (7). Encapsulated chondrocytes in hydrogels have been shown to
generate native cartilage tissue (11). Although the PEG hydrogel does not contain any chemical
signaling motifs, it can be easily modified to become bioactive. A bioactive scaffold will afford
regulated cell attachment and increased differentiated function. Hyaluronic acid (HA) is a high
molecular weight glycosaminoglycan present in the ECM of all mammalian connective tissue. It
has the ability to form an ionic hydrogel with multiple modification sites through its carboxyl
and hydroxyl groups (12). Collagen, a fibrous protein, is the main component found in
mammalian tissues like bone and cartilage (9). By incorporating ECM components that provide
biochemical signaling, the synthetic hydrogel will provide more natural conditions for matrix
and tissue formation upon cell encapsulation.
3
Osteochondral integration is a key component since the synthetic scaffold should
facilitate cartilage tissue formation, as well as bone. The tissue must be functionally integrated
with the host tissue including the surrounding cartilage and subchondral bone (13). This requires
an interface with a gradient of mechanical properties progressing from soft tissue to bone.
Cellular communication is required at the interface in order for biological fixation to occur,
which can be accomplished by producing a functional gradient that possesses mechanical
properties similar to the native tissue-to-tissue interface (14). By incorporating adhesive peptides
into the hydrogel, like arginine–glycine–aspartate (RGD), cellular adhesion and spreading is
facilitated (8). Integrins are cell surface receptors that mediate adhesion between ECM and cells
by binding to ligands that possess the RGD sequence. These receptors stimulate intracellular
signaling and gene expression (15).
Only few studies have investigated the in vitro response of chondrocytes and osteoblasts
into these scaffolds (7). Our goal was to create the optimal environment for cellular proliferation
and differentiation for chondrocytes and osteoblasts.
EXPERIMENTAL SECTION
Materials. Poly(ethylene glycol) (PEG, FW ≈ 8000 g/mol (8k)), methacrylic anhydride (MA),
hydroquinone, ethyl ether, and triethylamine (TEA) were purchased from Sigma-Aldrich and
used as received. Dichloromethane was purchased from Sigma-Aldrich and dried over activated
molecular sieves (3 Å) prior to use. Photoinitiator IRGACURE 2959 (I2959) was obtained from
Ciba Specialty Chemicals and used as received.
Synthesis of PEGDM. PEGDM (poly (ethylene glycol) dimethacrylate) was prepared from the
reaction of PEG and MA. The synthesis of a 4.6k PEGDM is as follows: PEG (5 g, ~ 0.001
mol), 2.2 equiv of methacrylic anhydride (MA) (0.34 g, 0.0022 mol), triethylamine (TEA) (0.2
mL), and hydroquinone (250 mg) were reacted in ~15 mL of dichloromethane over freshly
activated molecular sieves (~ 3 g) for 4 d at room temperature. The solution was filtered and
precipitated into ethyl ether. The product was dried in a vacuum prior to use. The product was
analyzed using 1H NMR (Varian Mercury 300) and MALDI-TOF MS (Bruker). Deuterated
chloroform was used as a solvent, and the polymer concentration was 2.5 % by mass fraction for
4
the NMR analysis. For MALDI analysis, PEG was dissolved in 1 mL of THF, and sodium was
used as the cationizing reagent in a 1:1 by volume ratio THF: PEG.
Hydrogel preparation for bovine and human chondrocytes. 10 % PEGDM (~8000 g/mol) OptiMEM I reduced-serum medium solution containing 0.1% IRGACURE 2959 with 1% Col I and
0.5% or 1% HA was photopolymerized using ~2.3 mJ/cm2 UVA light for 5 min. Samples (~1 cm
diameter, 0.3 cm thick) encapsulating either 1 million bovine (2-3 week old) or 2 million human
(knee arthroscopy) chondrocytes per 100 µL of 10% PEGDM Opti-MEM solution yielding
samples (~1 cm diameter, 0.3 cm thick) which were cultured up to 6 weeks in Opti-MEM
containing 50 µg/mL ascorbate and 100ug/mL primocin. The primary chondrocytes were
cultured in growth medium composed of Dulbecco's modified Eagle medium, 10% fetal bovine
serum, 1% minimum essential medium (GIBCO, Invitrogen Corp), 50 μg/mL l-ascorbic acid 2phosphate (Sigma), and 1% antibiotics (penicillin/streptomycin) (Mediatech, Inc.). The media
was changed 3 times a week.
Histology & Biochemistry. Proteoglycan staining was conducted on 10% formalin fixed whole
samples with 0.01% thionin. For biochemical testing, samples were desiccated overnight in a
vacuum centrifuge or freeze dryer. Sulfated gylcosaminoglycans (sGAGs) were quantified with
dimethylmethlene blue and collagen content was quantified using dimethylaminobenzaldehyde
to observe chloramines T-oxidized hydroxyproline. All biochemistry was normalized against
samples without cells cultured in media for the same duration and then normalized to the
PEGDM samples.
Solid-Phase Peptide Synthesis. The RGD peptide was synthesized by standard solid-phase
synthesis using Fmoc chemistry. The synthesis consisted of the following steps: swelling of 1.0
g of 1.0 mmol/g resin in DMF, deprotection through Fmoc cleavage in piperidine (0.25 vol
fraction in DMF), rinse cycles, and coupling at 5X molar excess amino acid (0.50 mol/L amino
acid concentrations). After the completion of the synthesis, the peptide is cleaved from the resin
using a solution. The retained product was lyophilized and verified using matrix-assisted laser
desorption/ionization (MALDI).
5
Fabrication of Hydrogel Modulus Gradients. RGD peptide (0.440 M) was added to the 5%
PEGDM along with pre-osteoblast cell line, MC3T3EI. 5%, 15%, and 50% PEGDM (~8000
g/mol) Dulbecco’s Modified Eagle Medium (DMEM) (Invitrogen) medium solutions containing
0.1% Irgacure 2959 were combined using a computer driven gradient maker. Samples (~ 10 cm
in length, 1cm in width) encapsulating 3.85 X 10^6 cells/mL (MC3T3E1) were polymerized
using ~2.3 mJ/cm2 UVA light for 5 min. Samples were cultured up to one week in MEM Alpha
Medium (1X) (Invitrogen) with 0.1% ascorbic acid (Sigma) and 1% beta glycerol phosphate
(Sigma). Media was changed three times a week.
MC3T3E1 Gradient Hydrogel Analysis. Gradient hydrogel analysis was completed at 1, 3, 5, and
7 day time points. Wst-1 assay was used to measure cellular metabolic dehydrogenase activity in
non-homogenized gels. Gel sections were incubated for 2 h in 24-well plates at 37°C with (0.5
mL) Wst-1 Solution containing (22 mM Wst-1 stock solution and 32 mM 1-methoxy PMS stock
solution). Aliquots (0.2mL) from the reacted solutions were transferred to a 96-well plate and
absorbance values were measured at 450 nm by a microplate reader (BioTek). The DNA Sigma
Fluorescence Kit was used to quantify DNA. BioRad DC Protein Assay Kit was utilized for
total protein content. Anaspec Sensolyte pNPP Alkaline Phosphatase Assay Kit was utilized to
detect alkaline phosphate activity in biological samples. Manufacturer’s protocol was followed
for each assay.
RESULTS
PEGDM Synthesis
PEGDM was prepared for the formation of photo-crosslinkable hydrogels. The PEG
hydroxyl groups react with methacrylic anhydride to form PEGDM.
Scheme 2: The synthesis of PEGDM using triethylamine and methacrylic anhydride.
6
The combination of proton NMR and MALDI-TOF MS provide insight on the degree of
methacrylate conversion as well as product purity. Figure 1 shows the 1H NMR of PEGDM.
The main signal is observed at a δ of 3.65. The chemical shift of methylene protons on the
methacrylate groups respectively are δ = 6.35 and 5.77. Any unreacted methacrylic anhydride
and triethylamine have been removed since the spectra exhibits the expected peaks. Molecular
mass, molecular mass distribution, and endgroup functionalities can be determined using
MALDI-TOF MS. As a result, this tool can be used to determine the amount of PEGDM versus
the amount of unreacted PEG or PEG that has reacted one hydroxyl group. The MALDI-TOF
MS spectrum (Figure 2) illustrates the degree of methacrylate conversion and polydispersity.
The molecular mass can be determined from the MALDI, as shown in Figure 2.
Figure 1: 1H NMR spectrum of PEGDM
Figure 2: MALDI-TOF mass spectrum of 4.6k PEGDM
7
Hydrogel Characterization
The hydrogel’s properties play an important role on ECM formation, since properties
determine if the scaffold will provide an optimal environment for cellular response. Primary
bovine and human chondrocytes were encapsulated in 10% PEGDM solution as well a s ECM
components, HA and collagen type I. Figure 3 compiles the change in the swelling ratio and
Swelling Ratio
figure 4 compiles the mesh size of the four different sample groups.
18
16
14
12
10
8
6
4
2
0
10%pedm
10% Pegdm
+1%collagen
10% Pegdm 10%Pegdm +1%
+1% collagen collagen +1%
+0.5%
hyaluronic acid
hyaluronic acid
Figure 3: Swelling Ratio of PEGDM hydrogels with various concentrations of ECM additives.
The greatest swelling ratio is observed when 10% PEGDM is combined with 1%
collagen and 0.5% hyaluronic acid when compared to the rest of the varying conditions (Figure
3). The mesh size displayed similar characteristics as the swelling ratio, since the 1% collagen
and 0.5% hyaluronic acid exhibited the greatest mesh size as well (Figure 4). Although the other
groups were not significantly different, the samples that had ECM components had a slightly
smaller mesh size than the synthetic polymer.
8
Mesh Size (A)
140
135
130
125
120
115
110
105
100
10%pedm
10% Pegdm
+1%collagen
10% Pegdm
10%Pegdm
+1% collagen +1% collagen
+0.5%
+1% hyaluronic
hyaluronic acid
acid
Figure 4: Mesh Size of PEGDM hydrogels with various concentrations of ECM additives.
The following table shows the results of the characterization studies. The greatest
swelling ratio and mesh size was observed when 1% Col I and 0.5% HA were added to the
synthetic polymer. It was found that incorporating ECM components did not lower the storage
modulus in every case. Porosity and mesh size are important for nutrient and waste transport;
therefore, these factors need to be finely controlled in order for cell attachment, proliferation, and
cellular differentiation to occur.
Sample
Swelling Ratio
Mesh Size
(Å)
Storage Modulus
(PA)
Loss Modulus
(PA)
10% PEGDM
13.5 ± 0.935
120.2 ± 7.01
1578 ± 60
793 ± 78
10% PEGDM + 1%
Col I
13.2 ± 1.11
114.4 ± 6.20
2214 ± 189*
545 ± 103*
10% PEGDM + 1%
Col I + 0.5% HA
15.9 ± 0.362**
132.67 ±
1.83*
1275 ± 51**
852 ± 109
10% PEGDM + 1%
Col I + 1% HA
13.3 ± 0.957
112.25 ±6.46
1238 ± 34**
865 ± 6.2
Table 1: Summary of characterization results
** indicates p-value <0.01 and * indicates p-value <0.05 compared to 10% PEGDM.
9
Cellular Response of Bovine and Human Chondrocytes
Bovine and human primary chondrocytes were encapsulated in 10 % PEGDM with various
concentrations of the ECM components, collagen I and HA. The four different groups are 10%
PEGDM, 10% PEGDM + 1% Col I, 10% PEGDM + 1% Col I +0.5% HA, and 10% PEGDM +
1% Col I + 1% HA. The hydrogels were cultured up to 6 weeks in vitro, after which histological
and biochemical analysis were conducted. Proteoglycan, chondroitin sulfate, and
hydroxyproline content were analyzed for the four different groups. Proteoglycans play an
important role in cell and ECM adhesion, and they are composed of glycoproteins containing
covalently linked glycosaminoglycans (GAGs). Hydrogels encapsulating bovine and human
chondrocytes after 6 weeks in vitro show increased proteoglycan staining compared to synthetic
hydrogels (Figures 5 and 6). Chondroitin sulfate and hydroxyproline content is greatest in both
human and bovine chondrocytes when 10% PEGDM + 1% Col + 0.5% HA are used to construct
the hydrogel as compared to the other sample groups (Figures 7-10).
A
B
A
B
C
D
C
D
Figure 5: Proteoglycan staining of bovine
chondrocyte hydrogels 6 weeks after in vitro
culture. A) 10% PEGDM, B) 10% PEGDM
+1% Col I, C) 10% PEGDM + 1% Col I + 0.5%
HA, D) 10% PEGDM + 1% Col I + 1% HA.
Scale bar = 100 µM
Figure 6: Proteoglycan staining of human
chondrocyte hydrogels 6 weeks after in vitro
culture. A) 10% PEGDM, B) 10% PEGDM +1%
Col I, C) 10% PEGDM + 1% Col I + 0.5% HA,D)
10% PEGDM + 1% Col I + 1% HA. Scale bar =
100 µM
10
Relative Chondroitin Sulfate Content
*
4
*
3.5
3
2.5
2
1.5
1
0.5
0
10% PEGDM
10% PEGDM + 10% PEGDM + 10% PEGDM +
1% Col I
1% Col I + 0.5% 1% Col I + 1%
HA
HA
Figure 7: Relative Chondroitin sulfate content of bovine chondrocytes encapsulated in
hydrogels 6 weeks in vitro. Hydrogels with 1% Col I and 0.5% HA were found to
contain significantly more chondroitin sulfate content than samples without ECM
additives. * indicates p-value < 0.05.
*
4
*
3.5
3
.
2.5
2
1.5
1
0.5
0
10% PEGDM
10% PEGDM + 10% PEGDM + 10% PEGDM +
1% Col I
1% Col I + 0.5% 1% Col I + 1%
HA
HA
Figure 8: Relative hydroxyproline content of bovine chondrocytes. Hydrogels
encapsulating bovine chondrocytes with 1% Col I and 0.5% HA were found to contain
significantly more hydroxyproline content than samples without ECM additives.
*indicates p-value < 0.05
11
Relative Chondroitin Sulfate Content
2.5
2
1.5
1
0.5
0
10% PEGDM
10% PEGDM 10%PEGDM + 10% PEGDM +
+1% Col I
1% Col I + 0.5% 1% Col I + 1%
HA
HA
Relative Hydroxyproline Content
Figure 9: Relative Chondroitin sulfate content of human chondrocytes encapsulated in
hydrogels 6 weeks in vitro. Hydrogels with 1% Col I and 0.5% HA were found to contain
significantly more sGAGs than samples without ECM additives.
2.5
2
1.5
1
0.5
0
10% PEGDM
10% PEGDM 10%PEGDM + 10% PEGDM +
+1% Col I
1% Col I + 0.5% 1% Col I + 1%
HA
HA
Figure 10: Relative hydroxyproline content of human chondrocytes.
Hydrogels encapsulating bovine chondrocytes with 1% Col I and 0.5% HA were found to
contain significantly more hydroxyproline content.
Gradient Hydrogel Characterization
Gradient hydrogels were created using a computer driven gradient maker with 5%, 15%, and
50% PEGDM concentrations. As a result, stiffness was varied by modulating the concentration
of PEGDM. Gradient hydrogels were cast into molds (5 cm x 1 cm x 1cm) and photoinitiated for
5 min. Figure 11 compiles the storage modulus of the gradient hydrogels at each position (5 mm
each position length) along the hydrogel (Figure 11). The storage modulus decreases the farther
the position on the hydrogel, signifying the highest concentration of PEGDM is in the first
12
position and lower concentrations are toward the end. Mesh size (Figure 12) and swelling ratios
(Figure 13) were also calculated according to position. The further the position is on the
Swelling Ratio
gradient, the higher the swelling ratio and mesh size.
18
16
14
12
10
8
6
4
2
0
Figure 13: Swelling ratio of gradient hydrogels slightly increases as
position distance increases from the start of the hydrogel.
Cellular Response of Functionalized Gradient Hydrogels
The quantification of DNA and ALP (alkaline phosphatasez) provide insight to the
cellular response of the pre-osteoblast cell line, MC3T3E1. A change in alkaline phosphate
activity is involved in bone development, thus permitting differentiation. DNA quantification
provides insight to cellular proliferation and as DNA content increases this signifies cellular
proliferation within the hydrogels (Figure 14). ALP activity increased as well; thus, pre-
DNA (ng)
osteoblast differentiation occurred throughout the time points (Figure 15-17).
100
0.7
80
0.6
60
Day 1
40
Day 3
20
Day 5
0
Day 7
1 2 3 4 5 6 7 8 9 10
Gradient Position
Figure 14: DNA quantification of positions along the
gradient at various time points.
0.5
0.4
0.3
0.2
0.1
0
1
2
3
4
5
6
7
8
9
10
Figure 15: ALP content of gradient positions on day
13
one.
0.5
0.45
0.4
0.35
0.3
0.25
0.2
0.15
0.1
0.05
0
0.4
0.3
0.2
0.1
0
1
2
3
4
5
6
7
8
9
10
Figure 16: ALP content of gradient positions on
day three.
day 7
day5
1 2 3 4 5 6 7 8 9 10
Figure 17: ALP content of gradient positions on day five
and seven.
DISCUSSION
Tissue engineering scaffolds should provide an optimal environment for both cartilage
and bone tissue formation. Bone is a tissue that is composed of a dense mineralized matrix that
can endure significant compressive loads; therefore, it requires a scaffold that can provide a
foundation for mechanical stability as well as a suitable porosity for cellular transport (5).
Cartilage provides protection of the subchondral bone in order for bone to withstand compressive
forces as well as smooth movement (15).
Cartilage has a poor healing capacity since it is isolated from systemic regulation (6).
When cartilage is damaged, chondrocytes cannot migrate from healthy tissue to the site of injury
like other tissues. Chondrocytes are surrounded by the ECM and lack a blood and nerve supply;
therefore, they are secluded (14). Both cartilage and bone arise from mesenchymal stem cells
and exist concurrently at articular surfaces of synovial joints. Since bone and cartilage are
typically organized into three dimensional structures in the body, the scaffold must provide
adequate space for cellular growth (13). If a scaffold could synchronously support the growth of
both cartilage and bone tissue, osteochondral integration would be facilitated (5). By utilizing a
scaffold whose mechanical properties and porosity are easily controlled, synthetic materials will
provide more natural conditions for matrix and tissue formation.
Hydrogels provide flexibility in design, since they can easily be tailored to one’s needs.
PEGDM hydrogels were modified to become bioactive with concentrations of ECM components,
hyaluronic acid and collagen type I. Bovine and human chondrocytes were encapsulated into the
14
hydrogels to observe the effects. Both bovine and human chondrocytes were found to contain a
greater content of proteoglycans with the ECM additives. The main ECM components in
cartilage are collagen, glycosaminoglycans, and proteoglycans. Proteoglycans are protein
molecules that have covalently linked glycosaminoglycans. HA is one of the major
proteoglycans in cartilage and has a role in maintaining the integrity of cartilage when it is under
compressive forces. The negatively charged glycosaminoglycans and proteoglycans allow for
the perpetual hydration of the tissue. As a result, cartilage possesses high water content and this
necessitates diffusion within the tissue (15). Chondroitin sulfate (glycosaminoglycan) and
hydroxyproline (major collagen component) content were found to be the greatest in the 10%
PEGDM + 1% Col I + 0.5% HA sample. The mechanical properties were also analyzed of the
four sample groups, and the 1% Col I and 0.5% HA hydrogels exhibited a greater swelling ratio
and mesh size. An increase in porosity and swelling ratio facilitate the cellular, nutrient, and
waste transport. This ultimately promotes cell proliferation and differentiation. Although one
would expect the synthetic scaffold to have the greatest storage modulus, it did not. The group
that had the highest storage modulus was the addition of 1% Col I to the PEGDM. This is
probably due to collagen’s fibrous nature. Therefore, the inclusion of low concentrations of Col
I and HA promotes ECM production.
To examine the response of bone, the pre-osteoblast cell line, MC3T3E1, was
encapsulated in functionalized gradient hydrogels. Since hydrophilic polymers may not provide
an optimal environment for osteoblast adhesion, a biomimetic, functionalized hydrogel that
promotes cellular adhesion is required (16). Osteoblasts require adhesion in order for
development and ECM production. By incorporating the peptide, RGD, cells will be able to
interact with their environment and modify their microenvironments. Cellular interactions have
improved with incorporated peptides since the peptide’s mobility is increased and steric
hindrance is decreased (17). A computer driven gradient pump was utilized to analyze the
mechanical properties of the gradient hydrogels produced. By synthesizing a scaffold with a
gradation of mechanical and structural properties, complex tissue-to- tissue interfaces may be
regenerated. As a result, gradient hydrogels will closely mimic the natural transition in
composition (18). The stiffness varied within the hydrogel between 15% - 50% PEGDM
concentrations. Storage modulus decreased the further the position from the start of the gradient.
15
This signifies that the higher concentrations of PEGDM are located at the beginning positions of
the hydrogel, and lower concentrations of PEGDM are located at the end positions. The
composition of PEG and the cross – linking density are directly related. Therefore, an increase
in polymer composition is accompanied with a reduction in the distance between cross-linking
points. The decrease in polymer-rich domains affects the transport of molecules within the
hydrogel (19). Swelling ratio and mesh size increased as position increased. Since the lower
concentrations of PEGDM have greater swelling ratios, indicating the softer a gel, the more
water it will absorb (20). This is the reason stiff hydrogels have smaller mesh sizes.
Cellular response was examined using DNA quantification and alkaline phosphatase
(ALP) content. ALP represents an early marker for osteoblast differentiation. DNA content
generally increased at further time points, specifically by day 5 and 7. This signifies cellular
proliferation (21). After five days the gradient hydrogels had greater DNA quantities for the
beginning positions in the hydrogel, while day 7 exhibited greater DNA content for the positions
toward the end of the hydrogel. ALP content increased greatly from day 1 to day 3 signifying
osteoblast differentiation. When comparing day 5 and day 7, there is a significant increase in the
differentiation marker.
The functionalized gradient hydrogel is able to provide an environment where osteoblasts
can proliferate and differentiate differently within one hydrogel. Therapies generated from tissue
engineering would improve the quality of life for thousands of Americans suffering from
osteoarthritis by providing them with numerous alternatives that will not have serious
drawbacks.
CONCLUSION
Results obtained from this study indicate that biomimetic and functionalized gradient
hydrogels can be engineered to become an interface between bone and cartilage by manipulating
their mechanical properties and surface properties. Future work will examine the effect of these
ECM on gene expression and examine the pathway’s activations triggering the increase in ECM
production. The effect of gradient hydrogels on human chondrocytes will also be further
explored. Ultimately, the goal is to design a single functional gradient hydrogel where both
tissues will exhibit their natural cellular activity.
16
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APPENDIX 1
SAFETY CONSIDERATIONS
Training and approval in biosafety was required to be completed before this project
commenced. This is primarily due to the exposure of human biological samples upon
chondrocyte isolation. Eating, drinking, inserting/removing contact lenses, and applying
cosmetics were prohibited. Safety glasses, gloves, and closed shoes were required at all times.
Needles and scalpels used when dealing with human specimen needed to be recapped and placed
in the biohazardous sharps waste. When finished with laboratory work, hands needed to be
thoroughly lathered with soap, rinsed in clean water, and dried with clean paper towels.
If exposure to biohazardous material did occur, the area was to be flushed with water for
15 min and then washed with soap and water. Exposure includes a spilled solution on an open
cut in the skin, needle prick (used for potentially biohazardous material), splashed cell culture
waste in eyes, nose, or mouth, and ingestion of biohazardous material. The supervisor and UA
Health Unit needed to be contacted, and a sample of the biohazardous material needed to be kept
for testing.
Biohazardous waste disposal required certain procedures to be followed for proper
removal. Liquid waste, like cell culture waste, needed to be decontaminated with bleach (1:10
dilution of bleach to water) then poured down the sink. Sharps waste needed to be placed in
biohazard sharps waste containers, and solid waste needed to be disposed in waste bags/boxes.
In case of a spill, the area needed to be decontaminated with bleach and then wiped with paper
towels. The paper towels would then enter the biohazard waste container. If the area is beyond
one’s capacity to clean, UA emergency needs to be notified.
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Finally, chemical safety was also required when synthesizing PEGDM. Hydroquinone,
methacrylic anhydride, triethylamine, and dichloromethane were all utilized in the reaction. Skin
contact, eye contact, ingestion, and inhalation all needed to be avoided with the compounds. If
the substance was exposed to any mucous openings, the area needed to be flushed with water for
15 min. Medical attention needed to be sought if irritation occurred. Safety glasses, lab coat,
and gloves were required at all times when in the laboratory. Shorts and open-toed shoes were
prohibited, and hair and loose clothing needed to be secured. Once chemicals were utilized, they
needed to be stored by compatibility. Disposable of chemicals needed to be in the proper
container: non-halogenous vs halogenous waste containers. If a spill occurred, it needed to be
absorbed with an inert dry material and placed in the appropriate waste disposal container.
Finally, an important rule for safety everywhere is to be aware of your surroundings.
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