Characterization of Novel Photocrosslinked Carboxymethylcellulose Hydrogels for Encapsulation of...

7/25/2019 Characterization of Novel Photocrosslinked Carboxymethylcellulose Hydrogels for Encapsulation of Nucleus Pulpos

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Characterization of novel photocrosslinked carboxymethylcellulose hydrogels

for encapsulation of nucleus pulposus cells

Anna T. Reza, Steven B. Nicoll *

Department of Bioengineering, University of Pennsylvania, Room 240 Skirkanich Hall, 210 S. 33rd Street, Philadelphia, PA 19104, USA

a r t i c l e i n f o

Article history:Received 18 January 2009

Received in revised form 26 May 2009

Accepted 1 June 2009

Available online 6 June 2009

Keywords:

Carboxymethylcellulose

Hydrogels

Nucleus pulposus

Photopolymerization

Mechanical properties

a b s t r a c t

Back pain is a significant clinical concern often associated with degeneration of the intervertebral disc(IVD). Tissue engineering strategies may provide a viable IVD replacement therapy; however, an ideal

biomaterial scaffold has yet to be identified. One candidate material is carboxymethylcellulose (CMC),

a water-soluble derivative of cellulose. In this study, 90 and 250 kDa CMC polymers were modified with

functional methacrylate groups and photocrosslinked to produce hydrogels at different macromer con-

centrations. At 7 days, bovine nucleus pulposus (NP) cells encapsulated in these hydrogels were viable,

with values for the elastic modulus ranging from 1.07 0.06 to 4.29 1.25 kPa. Three specific formula-

tions were chosen for further study based on cell viability and mechanical integrity assessments: 4%

90 kDa, 2% 250 kDa and 3% 250 kDa CMC. The equilibrium weight swelling ratio of these formulations

remained steady throughout the 2 week study (46.45 3.14, 48.55 2.91 and 42.41 3.06, respectively).

The equilibrium Youngs modulus of all cell-laden and cell-free control samples decreased over time, with

the exception of cell-laden 3% 250 kDa CMC constructs, indicating an interplay between limited hydro-

lysis of interchain crosslinks and the elaboration of a functional matrix. Histological analyses of 3%

250 kDa CMC hydrogels confirmed the presence of rounded cells in lacunae and the pericellular deposi-

tion of chondroitin sulfate proteoglycan, a phenotypic NP marker. Taken together, these studies support

the use of photocrosslinked CMC hydrogels as tunable biomaterials for NP cell encapsulation.

2009 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.

1. Introduction

The intervertebral disc (IVD) is a heterogeneous tissue that

functions to permit motion and flexibility, support and distribute

loads, and dissipate energy in the spine[1]. The IVD is composed

of the collagenous, lamellar annulus fibrosus, which maintains disc

shape and allows the spine to resist tensile loads[2], and the gelat-

inous nucleus pulposus (NP). The NP is a hydrated tissue, charac-

terized by high proteoglycan (i.e. aggrecan) and type II collagen

content [1]. This region functions to resist compressive loads

through the generation of a hydrostatic swelling pressure.

Degeneration of the IVD is strongly associated with back pain, a

significant healthcare problem, afflicting approximately 80% of

Americans during their lifetime [3] and costing over $80 billion

in annual related medical expenses[4]. Disc degeneration often re-

sults from traumatic injury or occurs naturally with aging. This

pathological condition is commonly attributed to increased degra-

dation of aggrecan molecules, giving rise to significant alterations

in disc biochemical composition and a loss of hydration [5]. The

NP is thus rendered more fibrous in structure and content, which

reduces nutrient diffusion and waste removal [6]. The resulting

increase in lactate concentration within the tissue lowers the local

pH [7]. The increased acidity compromises cell metabolism and

may precipitate cell death, as up to 50% of cells in adult discs have

been reported as necrotic [8]. Disc degeneration may be asymp-

tomatic, or the change in extracellular matrix (ECM) content may

contribute to increased disc stiffness and low back pain from the

altered distribution of loads [1]. Current clinical treatments focus

on alleviating pain rather than restoring the structure and function

of the disc. Tissue engineering strategies may provide a biologic

alternative capable of restoring both the structure and mechanical

function of the IVD.

Biomaterial scaffolds used in tissue engineering applications of-

ten attempt to mimic the native structure of the respective tissue.

The highly hydrated nature of the NP is similar to that of hydrogel

networks, making such materials prime candidates to serve as

scaffolds for NP regeneration. Hydrogels are hydrophilic, cross-

linked polymers which absorb large volumes of water and swell

without dissolution of the polymer [9]. Nucleus pulposus cells

are routinely cultured by encapsulation in hydrogels made from

alginate, a naturally derived polysaccharide originating from

brown algae [1014]. Alginate cell encapsulation promotes a

rounded, chondrocyte-like morphology, in contrast to the elon-

gated, fibroblast-like morphology seen in monolayer cultures.

The predominant method of alginate gelation is through ionic

1742-7061/$ - see front matter 2009 Acta Materialia Inc. Published by Elsevier Ltd. All rights reserved.doi:10.1016/j.actbio.2009.06.004

* Corresponding author. Tel.: +1 215 573 2626; fax: +1 215 573 2071.

E-mail address:[email protected](S.B. Nicoll).

Acta Biomaterialia 6 (2010) 179186

Contents lists available at ScienceDirect

Acta Biomaterialia

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http://dx.doi.org/10.1016/j.actbio.2009.06.004mailto:[email protected]://www.sciencedirect.com/science/journal/17427061http://www.elsevier.com/locate/actabiomathttp://www.elsevier.com/locate/actabiomathttp://www.sciencedirect.com/science/journal/17427061mailto:[email protected]://dx.doi.org/10.1016/j.actbio.2009.06.004


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crosslinking, achieved via diffusion of divalent cations to carbox-

ylic acid moieties on the polymer, resulting in a crosslinked net-

work. Although this initially produces stable gels, mechanical

integrity has been found to decrease over time, possibly due to a

loss of ions through diffusion[10]or depletion by the encapsulated

cells.

Photopolymerization has been widely used in situ to covalently

crosslink polymer networks in dental applications [9,15]. Thismethod employs biocompatible, light-sensitive photoinitiators

that absorb light, creating free radicals that can initiate polymeri-

zation to covalently crosslink functional groups along the polymer

backbone[9]. Elisseeff et al.[15]developed a photopolymerization

method to successfully encapsulate chondrocytes in poly(ethylene

oxide)-based (PEO) hydrogels for tissue engineering applications.

This technique was modified by Bryant et al. [16] to incorporate

degradable lactic acid units into poly(ethylene glycol)-based

(PEG) hydrogels to enhance the spatial distribution of ECM compo-

nents in these otherwise inert polymers. Photopolymerization has

also been employed to create polysaccharide-based hydrogels

using alginate and hyaluronic acid macromers modified with func-

tional methacrylate groups [17]. In an extension of this work,

methacrylated hyaluronic acid was used to engineer hydrogels

for cartilage cell encapsulation [18,19]. These constructs were

shown to allow for functional matrix production following

in vivo implantation in a subcutaneous murine pouch model

[20]. Although hyaluronic acid-based hydrogels have shown prom-

ising results, these biomaterials are often derived from an animal

source, which presents the risk of batch-to-batch variations and

the need for additional purification steps to reduce the possibility

of stimulating an immune response upon implantation. Neverthe-

less, the results observed for hyaluronic acid-based hydrogels indi-

cate the potential of similar polysaccharide-based systems for use

in additional orthopaedic tissue engineering applications.

One such candidate polysaccharide is carboxymethylcellulose

(CMC), a water-soluble derivative of cellulose, the primary struc-

tural component of plant cell walls. CMC is a biocompatible, low-

cost, FDA-approved material that is commercially available inhigh-purity forms, making this polymer a highly attractive option

for biomedical applications[21].

CMC-based materials have not been used previously for IVD re-

pair. Therefore, the objective of this study was to create photo-

crosslinked CMC hydrogels with tunable material properties for

NP cell encapsulation. We hypothesized that CMC macromers

could be synthesized with methacrylate groups that would allow

for photopolymerization. Moreover, an increase in CMC molecular

weight and weight percent would be expected to give rise to an in-

verse relationship between the equilibrium Youngs modulus and

the swelling ratio of the resulting hydrogels.

2. Materials and methods

2.1. Macromer synthesis

Methacrylated-carboxymethylcellulose (Me-CMC) was synthe-

sized through esterification of hydroxyl groups based on previ-

ously described protocols [17,18] (Fig. 1). Briefly, 1 g of 90 or

250 kDa CMC (Sigma, St. Louis, MO) was dissolved in 100 ml of

RNAse/DNAse-free water at 50 C and stirred for 30 min. The mix-

ture was then stirred at room temperature for 1 h before finally

being placed on an orbital shaker for 48 h at 4C to yield a

1 wt.% solution. Methacrylic anhydride (Sigma) at 20-fold excess

was reacted with 1% CMC over 24 h at 4C with 12 periodic adjust-

ments to pH 8.0 using 3 N NaOH to modify hydroxyl groups of the

polymer with functional methacrylate groups. The modified CMCsolution was purified via dialysis for 96 h against RNAse/DNAse-

free water (Spectra/Por1, MW 58 kDa, Rancho Dominguez, CA)

to remove excess, unreacted methacrylic anhydride. Purified Me-

CMC was recovered by lyophilization and stored at 20C. The de-

gree of substitution was confirmed using 1H NMR (360 MHz,

DMX360, Bruker, Madison, WI) following acid hydrolysis of puri-

fied Me-CMC. Briefly, a 20 mg sample of lyophilized Me-CMC was

dissolved in 20 ml of RNAse/DNAse-free water and hydrolyzed at

a pH of 2.0 at 80 C for 2.3 h. The pH of the hydrolyzed solution

was readjusted to 7.0, recovered via lyophilization and resus-

pended in deuterium oxide. Molar percent of methacrylation wasdetermined by the relative integrations of methacrylate proton

peaks (methylene peak, d= 6.2 and 5.8 ppm; methyl peak,d= 2.0 ppm) to carbohydrate protons.

2.2. Cell isolation

All cell culture supplies, including media, antibiotics and buffer-

ing agents, were purchased from Invitrogen (Carlsbad, CA) unless

otherwise noted. Discs C2C4 were isolated from bovine caudal

IVDs obtained from a local abattoir, and the NP was separated

through gross visual inspection based on previous protocols

[22,23]. Tissue was maintained in Dulbeccos modified Eagles

medium (DMEM) supplemented with 20% fetal bovine serum

(FBS) (Hyclone, Logan, UT), 0.075% sodium bicarbonate, 100 U ml1

penicillin, 100 lg ml1 streptomycin and 0.25 lg ml1 Fungizone

Fig. 1. Schematic of the synthesis of methacrylated carboxymethylcellulose.

180 A.T. Reza, S.B. Nicoll / Acta Biomaterialia 6 (2010) 179186


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reagent at 37 C in 5% CO2 for 2 days prior to digestion to ensure

that no contamination occurred during harvesting. A single serum

lot was used for all experiments to reduce potential variability in

the cellular response.

Tissue was diced and NP cells were released by collagenase

(Type IV, Sigma) digestion at an activity of 7000 U of collagenase

per g of tissue. Following incubation in collagenase, undigested tis-

sue was removed using a 40 lm mesh filter. Cells from multiplelevels (C2C4) were pooled and rinsed in sterile Dulbeccos phos-phate-buffered saline (DPBS). These primary cells were plated onto

tissue culture flasks and designated as passage 0. Cells were sub-

cultured twice to obtain the necessary number of cells, and passage

2 cells were used in all experiments[22].

2.3. Cell encapsulation in photocrosslinked hydrogels

Cell-encapsulated photocrosslinked constructs were prepared

at various weight percents. Prior to dissolution, lyophilized Me-

CMC was sterilized by a 30 min exposure to germicidal ultraviolet

(UV) light. The sterilized product was then dissolved in filter-ster-

ilized 0.05 wt.% photoinitiator, 2-methyl-1-[4-(hydroxyeth-

oxy)phenyl]-2-methyl-1-propanone (Irgacure 2959, I2959, CibaSpecialty Chemicals, Basel, Switzerland), in sterile DPBS at 4 C to

various weight percents (90 kDa Me-CMC: 3.2%, 4.2% and 5.2%;

250 kDa: 1.2%, 2.2% and 3.2%). Passage 2 NP cells were resuspended

in a small volume of 0.05% photoinitiator and then homogeneously

mixed with dissolved Me-CMC at 30 106 cells ml1. The seeding

density was selected based on previous studies using cell-seeded

constructs for engineering of cartilaginous tissues [2428]. Solu-

tions were cast at final concentrations of 3%, 4% and 5% (90 kDa

Me-CMC) and 1%, 2% and 3% (250 kDa Me-CMC) in a custom-made

glass casting device. The mixtures were exposed to long-wave UV

light (EIKO, Shawnee, KS, peak 368 nm, 1.2 W) for 10 min to pro-

duce covalently crosslinked hydrogel disks of 8 mm diame-

ter 2 mm thickness. Each hydrogel was incubated in 3 ml of

growth medium (DMEM with 10% FBS, 100 U ml

1

penicillin,100 lg ml1 streptomycin and 0.075% sodium bicarbonate) at37C under 5% CO2. At day 1, the medium was fully exchanged

with vitamin C (L-ascorbic acid) supplemented medium (growth

medium with 50 lg ml1 L-ascorbic acid), which was used for theremainder of the study and replaced every 23 days. Initial viabil-

ity studies (described below) were performed using gels cast at

5 mm diameter 2 mm thickness and were incubated in 1.5 ml

growth medium.

2.4. Cell viability and dynamic mechanical testing

Preliminary screening studies examined the effects of weight

percent and molecular weight on cell viability and the elastic

mechanical properties of 3%, 4% and 5% 90 kDa Me-CMC and 1%,2% and 3% 250 kDa Me-CMC. Cell viability was assessed at days 1

and 7 using the MTT (3-(4,5-dimethylthiazolyl-2)-2,5-diphenyltet-

razolium bromide) proliferation assay kit (ATCC, Manassas, VA).

Photocrosslinked Me-CMC hydrogels (n= 4) were incubated in

1 ml of growth medium supplemented with 100 ll of yellow tetra-zolium MTT for 4 h at 37 C, 5% CO2, shielded from light. Hydrogels

were then homogenized and formazan crystals were extracted

using the MTT detergent solution (MTT cell proliferation assay

kit, ATCC) for an additional 4 h at room temperature, shielded from

light. Total absorbance of the solubilized product was quantified at

570 nm using a Bio-Tek Synergy-HT microplate reader (Winooski,

VA). MTT absorbance values were compared between samples

and to cell-free control gels to quantify relative viability. Day 7

measurements were also evaluated against day 1 values to deter-mine loss of viability over time.

Cell viability of 2% 250 kDa CMC constructs was visually as-

sessed at day 0 (1 h after casting) and day 1 using the Live/Dead

kit (Invitrogen). Samples were rinsed in DPBS and then incubated

in Live/Dead solution (1 mM calcein AM, 1 mM ethidium homodi-

mer-2) for 45 min. Images were captured using a Zeiss Axiovert

200 microscope with fluorescent capabilities at excitation/emis-

sion wavelengths of 494/517 nm (calcein) and 528/617 nm (ethi-

dium homodimer-2/DNA complex). Live and dead cells werecounted using Image J software (National Institutes of Health).

At day 7, a Dynamic Mechanical Analyzer (DMA) 8000 (Perkin-

Elmer, Inc., Waltham, MA) testing apparatus was used to deter-

mine the elastic modulus of Me-CMC hydrogels at the weight per-

cents described above. Samples (n= 5) were rinsed in DPBS and

loaded into the DMA. Unconfined compression testing was per-

formed at 25 C at a strain rate of 10% min1. The modulus was

determined from the linear region of the stress vs. strain curves

at strains between 5% and 20%.

2.5. Swelling ratio

Following the initial studies examining cell viability and elastic

mechanical properties, 4% 90 kDa, 2% 250 kDa and 3% 250 kDa Me-

CMC hydrogels were chosen for further characterization. The equi-

librium weight swelling ratio,Qw, was determined for these formu-

lations at days 1, 7 and 14 for cell-laden and cell-free control

samples (n= 4). Constructs were weighed to determine the wet

weight (Ws), lyophilized and then weighed again to determine

dry weight (Wd). Qwwas calculated using the following equation:

Qw Ws=Wd

2.6. Characterization of equilibrium mechanical properties

Based on the early screening studies, unconfined compression

testing was conducted on 4% 90 kDa, 2% 250 kDa and 3% 250 kDa

Me-CMC cell-laden and cell-free control samples (n= 5) at days

1, 7 and 14 to measure the equilibrium Youngs modulus (Ey).The mechanical testing device is based on a similar setup described

by Soltz and Ateshian[29]. The device consists of a computer-con-

trolled stepper motor (Oriel Corp., Model 18515, Stratford, CT) that

prescribed a displacement on the specimen using a steel indenter

with glass platen attachment. A data card and LabVIEW software

(National Instruments, Austin, TX) were used for controlling the

stepper motor and data acquisition. Displacement was measured

using a linear variable differential transformer (Schaevitz, Model

PR812, Hampton, VA), and the load applied was measured using

a 250 g load cell (Sensotec, Model 31, Columbus, OH). Samples

were compressed between two impermeable glass platens in a

DPBS bath. The unconfined compression testing protocol was com-

prised of a creep test followed by a multi-ramp stressrelaxation

test. The creep test consisted of a 1 g tare load at 10 lm s1

rampvelocity for 1800 s until equilibrium was reached (equilibrium cri-

teria:


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ducted to visualize cellular distribution throughout the hydrogel.

Immunohistochemical analyses were performed to assess extracel-

lular matrix accumulation of chondroitin sulfate proteoglycan

(CSPG). Samples were treated with 0.5 N acetic acid for 2 h at

4 C. Non-specific binding was blocked using 10% goat serum

(Invitrogen) in DPBS. A monoclonal antibody to CSPG (1:100 dilu-

tion in blocking solution) (Sigma) was used, followed by incubation

in biotinylated goat/anti-mouse IgM secondary antibody (1:50dilution in blocking solution) (Vector Labs, Burlingame, CA). A per-

oxidase-based detection system (Vectastain Elite ABC, Vector Labs)

and 3,30-diaminobenzidine (Vector Labs) as the chromagen were

used according to the manufacturers protocols to detect ECM

localization. Non-immune controls were processed in blocking

solution without primary antibody. Samples were viewed with a

Zeiss Axioskop 40 optical microscope and images were captured

using AxioVision software.

2.8. Statistical analysis

A one-way analysis of variance (ANOVA) was performed on MTT

viability measurements to determine the effect of weight percent

and the effect of time. A two-way ANOVA was conducted on elasticmodulus data to determine the effects of weight percent and cells

(cell-laden vs. cell-free control constructs). A three-way ANOVA

was conducted on swelling and Ey data for 4% 90 kDa, 2%

250 kDa and 3% 250 kDa Me-CMC constructs to determine the ef-

fects of time, cells and starting material. A two-way ANOVA was

performed on equilibrium Youngs modulus measurements for 3%

250 kDa Me-CMC constructs to examine the effects of time and

cells. A Tukeys post hoc test was performed on all ANOVA calcula-

tions to detect significant differences between groups. All results

are presented as mean standard deviation with statistical signif-

icance defined as p


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4. Discussion

In this study, CMC was successfully modified with methacrylate

groups to produce photocrosslinked hydrogels with tunable prop-

erties. In addition, this is the first investigation to demonstrate suc-

cessful encapsulation of NP cells in photocrosslinked CMC

hydrogels, suggesting that these materials may serve as alternate

scaffolds for IVD replacement therapies.

Alginate is the most widely used biomaterial in NP tissue engi-

neering applications [1014]. However, the reversible nature of

conventional ionic crosslinking techniques has led to the investiga-

tion of photopolymerization methods. This approach has recently

been applied to alginate to produce mechanically stable hydrogel

constructs capable of supporting NP cell growth and viability

[30]. Nevertheless, raw alginate has been shown to stimulate an

immune response in vivo in mice and requires additional process-ing to remove impurities for biomedical applications[31]. Similar

purification procedures are required for animal-derived products,

such as hyaluronic acid, chitosan and chondroitin sulfate, whichhave also been used to engineer cartilaginous tissues [1720,32

35].

CMC is a well-established derivative of cellulose, which is

rendered water-soluble through the introduction of carboxy-

methyl groups along the polymer backbone. At physiological

pH, the carboxylic acid of the carboxymethyl group is deproto-

nated, resulting in a negatively charged polymer network,

which is similar to that provided by the glycosaminoglycans

in the ECM of cartilaginous tissues. CMC is commercially

available in high-purity forms, making it an appealing low-cost

alternative to other natural polysaccharides and inert polymers

currently used in tissue engineering applications. Additionally,

as a derivative of the plant-based polysaccharide cellulose,

CMC represents a renewable, environmentally friendlybiomaterial.

Fig. 2. Mitochondrial activity measurements (MTT) at days 1 and 7 for (A) photocrosslinked 90 and 250 kDa CMC hydrogels (n= 4) at various weight percents encapsulated

with bovine nucleus pulposus cells at 3 107 cells ml1. Representative day 7 MTT stereomicrograph images of 3% 90 kDa (B), 4% 90 kDa (C), 5% 90 kDa (D), 1% 250 kDa (E),

2% 250 kDa (F) and 3% 250 kDa (G) CMC cell-laden hydrogels (scale in mm). Live/Dead images of 2% 250 kDa CMC samples at days 0 (H) and 1 (I) with live cells stained green

and dead cells shown in red (bar = 100 lm). Significance set atp < 0.05. * , significant vs. 1% 250 kDa CMC within time point; #, significant vs. 1 and 3% 250 kDa CMC withintime point; +, significant effect of time within group.

A.T. Reza, S.B. Nicoll / Acta Biomaterialia 6 (2010) 179186 183


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Recent work has examined the efficacy of CMC-based hydrogels

for cell encapsulation [21,3638]. These studies have successfullyutilized various chemistries for CMC crosslinking, including phenol

modification of CMC carboxylic acid (COOH) groups [21,36], acry-

lation of CMC[37], amidation of CMC COOH groups[38]and elec-

trostatic interactions [39], highlighting the versatility of the CMC

polymer. Cytotoxicity has been tested using multiple cell lines with

positive results. However, some studies have examined CMC

hydrogels in sheet or membrane form, which may not be ideal

for orthopaedic applications. Additionally, modification of CMC

COOH groups limits the availability of charged moieties, thereby

reducing the swelling capability of the hydrogel and neutralizing

the negatively charged polymer network. Moreover, gels formed

through electrostatic interactions may not be as stable as cova-

lently crosslinked hydrogels an important feature of an orthope-

dic scaffold for load-bearing tissues.

Our early screening studies examined the effects of molecular

weight and macromer concentration on cell viability and elastic

modulus. Metabolic activity measurements using the MTT assay

showed no significant decrease over time for any group except

for the amorphous 1% 250 kDa CMC constructs. The lack of struc-tural integrity in these samples resulted in a significant loss of

material during transfer and may have contributed to the lower

than expected activity/viability measurements. Although the MTT

assay is routinely used to assess cell viability [15,30,4042], this

assay measures the activity of the mitochondrial enzyme succinate

dehydrogenase. Since the number of mitochondria can vary be-

tween cells, the MTT assay may not accurately reflect cell viability

[43]. Live/Dead fluorescent staining was used as a more direct eval-

uation of cell viability. Contrary to the MTT results, the staining

demonstrated a noticeable loss of viability over time. However, de-

creased cell viability in such hydrogel systems is not surprising, as

this trend has also been observed for bovine articular chondrocytes

encapsulated in PEO hydrogels and bovine NP cells encapsulated in

alginate, suggesting that additional environmental factors mayinfluence cell growth in photopolymerized hydrogels [30,40].

Although cell-interactive signals (i.e. growth factors, adhesive pep-

tides) have been shown to play an important role in modulating

cellular viability and function in engineered constructs [44,45],

the objective of this first study was to investigate cellpolymer

interactions in these novel photocrosslinkable CMC hydrogels,

excluding the influence of any exogenous factors.

Our initial screening study demonstrated the effects of CMC

molecular weight (90 and 250 kDa) and macromer concentration

on hydrogel properties. These studies underscore the influence of

crosslinking density on hydrogel material properties. Crosslinking

density is increased at higher macromer concentrations due to a

greater number of methacrylate functional groups available for

photoinitiated crosslinking. The theory of rubber elasticity predictsthat an increase in crosslinking density gives rise to an increase in

Fig. 3. Elastic modulus of day 7 photocrosslinked 90 kDa (A) and 250 kDa (B) CMC hydrogels encapsulated with bovine nucleus pulposus cells at 3 107 cells ml1 and

corresponding cell-free control gels (n= 5) at various weight percents. Significance set at p < 0.05. * , significant effect of weight percent.

Fig. 4. Equilibrium Youngs modulus for 4% 90 kDa and 2% 250 kDa CMC cell-free

control and cell-laden hydrogels (n= 5) over 14 days of in vitro culture. Significance

set at p < 0.05. * , significant vs. day 1 within group; +, significant vs. days 1 and 7

within group; #, significant vs. corresponding cell-free control.

Fig. 5. Equilibrium Youngs modulus for 3% 250 kDa CMC cell-free control and cell-

laden hydrogels (n= 5) over 14 days of in vitro culture. Significance set at p < 0.05.* ,

significant vs. day 1 within group; #, significant vs. corresponding cell-free control.



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hydrogel stiffness and a concomitant decrease in swelling ratio

[46]. The significant increases in elastic modulus associated with

increasing CMC molecular weight (as determined in our initial

studies), combined with the marked differences in swelling ratio

and equilibrium modulus of 2% vs. 3% 250 kDa CMC, are consistent

with the theory and with our original hypothesis.

Based on the results from our initial screening, the swelling ra-tio was characterized in cell culture medium for three formulations

of CMC: 4% 90 kDa CMC and 2% and 3% 250 kDa CMC. Qwremained

steady over time for all groups. Additional studies demonstrated

similar results in physiological saline and simulated body fluid

(data not shown). A stable swelling ratio is important for potential

IVD clinical applications as an intradiscal replacement material in

order to prevent bulging and extrusion into the annulus fibrosus.

AlthoughQw remained unchanged, the mechanical properties (Ey)

of 4% 90 kDa CMC and 2% 250 kDa CMC constructs experienced a

significant decrease over time for both cell-laden and cell-free con-

structs (Fig. 4). These two formulations were originally chosen for

more extensive characterization based on a study by Chou and Ni-

coll in which bovine NP cells were encapsulated in photocross-

linked methacrylated alginate hydrogels and implanted

subcutaneously in nude mice for 8 weeks [47]. The equilibrium

Youngs modulus of these alginate constructs was 1.25 kPa at

day 1 and increased to 4.31 kPa at 8 weeks, which indicated the

elaboration of a functional matrix that closely approximates values

of the native NP (5 kPa) reported by Cloyd et al. [48]. The results

of our initial study showed that the elastic modulus for 4% 90 kDa

CMC and 2% 250 kDa CMC constructs was 1 kPa at day 7 (Fig. 3).

As such, these formulations were selected for a more detailed anal-

ysis with the belief that the starting mechanical properties of the

scaffold would allow for matrix accumulation, resulting in a tem-

poral increase in modulus. However, Ey exhibited a continual de-

crease over time for both groups. Because CMC is a derivative of

cellulose, the polymer backbone is degraded by the plant-derived

enzyme, cellulase. As this enzyme was not introduced into the sys-

tem, the loss in mechanical properties was surprising. The decreasein modulus was observed for both cell-laden and cell-free con-

structs, indicating a non-cellular mediator of hydrogel weakening.

Although the schematic in Fig. 1 indicates methacrylation of the

hydroxyl group off of the C2 carbon, theoretically this could also

occur at a hydroxyl bonded to the C6 carbon. This arm would be

more susceptible to ester hydrolysis as the longer chain is less ste-

rically hindered, thereby resulting in the cleavage of periodic inter-

chain crosslinks without a significant loss in mass.

Due to the decrease in mechanical properties observed for 4%

90 kDa CMC and 2% 250 kDa CMC, a higher weight percent formu-

lation was chosen to provide a higher crosslinking density.

Although viability was robust in all concentrations of 90 kDa

CMC (Fig. 2A), a higher weight percent at this molecular weight

was not selected due to the large amount of starting material nec-

essary and the increased concentration of lingering free radicals.

Therefore, the 3% 250 kDa CMC formulation was selected. Similar

to 4% 90 kDa and 2% 250 kDa hydrogels, 3% 250 kDa cell-free con-

trol samples also experienced a temporal decrease in mechanical

properties (Fig. 5). However, the stiffer initial environment

(4 kPa) was on a par with native NP tissue (5 kPa) [48], and

cell-laden constructs elaborated a matrix that was able to over-come the decrease in mechanics and maintain the original modu-

lus. Unlike the softer 4% 90 kDa and 2% 250 kDa CMC hydrogels,

the partial hydrolysis of the stiffer 3% 250 kDa CMC constructs pro-

vided void space for the accumulation of secreted matrix macro-

molecules while maintaining sufficient structural integrity.

Histological analyses showed cells localized in lacunae throughout

the scaffold, as is typical of cartilaginous tissues (Fig. 6A), and the

pericellular deposition of CSPG was observed with pronounced

interterritorial staining at the periphery of the construct (Fig. 6B).

Although this study concentrated on characterizing the material

properties (degree of swelling and modulus) of cell-free and cell-

laden hydrogels, histological analyses confirmed the phenotypic

rounded morphology and elaboration of characteristic proteogly-

cans (i.e. CSPG) by encapsulated NP cells at 14 days in vitro. While

robust viability was verified at 7 days for all formulations, future

work will investigate the effects of time and environmental stim-

uli, such as growth factor supplementation (i.e. TGF-b3) [4952]

and mechanical loading (i.e. hydrostatic pressurization) [5358],

on cell viability and the functional assembly of phenotypic ECM

components.

Taken together, these findings indicate the utility of photocros-

slinkable CMC hydrogels for NP cell encapsulation, as these bioma-

terials support NP cell viability and may be easily tailored for

specific applications. Moreover, photocrosslinkable CMC may serve

as a cost-effective, biocompatible alternative to inert polymers,

including PEO and PEG, and expensive bacteria- and animal-de-

rived polysaccharides, such as hyaluronic acid and chondroitin sul-

fate, for use in the engineering of hydrated cartilaginous tissues.

Acknowledgements

This work was supported by NSF CAREER Award 0747968 and

an Early Career Translational Award from the Wallace H. Coulter

Foundation (S.B.N.) and an NSF Graduate Fellowship (A.T.R.). The

authors thank Dr. Jason Burdick and his laboratory for helpful dis-

cussions and access to equipment.

Appendix A

Figures with essential colour discrimination. Certain figures in

this article, particularly Figs. 1, 2 and 6, are difficult to interpret

in black and white. The full colour images can be found in theon-line version, at doi:10.1016/j.actbio.2009.06.004 .

Fig. 6. Hematoxylin and eosin staining (A) and chondroitin sulfate proteoglycan immunohistochemical staining (B) of cell-laden 3% 250 kDa CMC constructs at day 14.

Bar = 50 lm.

A.T. Reza, S.B. Nicoll / Acta Biomaterialia 6 (2010) 179186 185
http://dx.doi.org/10.1016/j.actbio.2009.06.004http://dx.doi.org/10.1016/j.actbio.2009.06.004


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References

[1] Buckwalter JA, Boden SD, Eyre DR, Mow VC, Weidenbaum M. Intervertebraldisk aging, degeneration, and herniation. In: Orthopaedic BasicScience. Rosemont, IL: American Academy of Orthopaedic Surgeons; 2000. p.55766.

[2] Hall S. Basic biomechanics. Boston, MA: McGraw-Hill; 2003. p. 276282.[3] Frymoyer JW, Cats-Baril WL. An overview of the incidences and costs of low

back pain. Orthop Clinics N Am 1991;22:26371.

[4] Martin BI, Deyo RA, Mirza SK, Turner JA, Comstock BA, Hollingworth W,Sullivan SD. Expenditures and health status among adults with back and neckproblems. JAMA 2008;299:65664.

[5] Larson JW, Levicoff EA, Gilbertson LG, Kang JD. Biologic modification of animalmodels of intervertebral disc degeneration. J Bone Joint Surg Am2006;88(Suppl. 2):837.

[6] Raj PP. Intervertebral disc: anatomyphysiologypathophysiology-treatment.Pain Pract 2008;8:1844.

[7] Buckwalter J. Do intervertebral discs deserve their bad reputation? IowaOrthop J 1998;18:111.

[8] Trout JJ, Buckwalter JA, Moore KC. Ultrastructure of the human intervertebraldisc. II. Cells of the nucleus pulposus. Anat Rec 1982;204:30714.

[9] Nguyen KT, West JL. Photopolymerizable hydrogels for tissue engineeringapplications. Biomaterials 2002;23:430714.

[10] Baer AE, Wang JY, Kraus VB, Setton LA. Collagen gene expression andmechanical properties of intervertebral disc cellalginate cultures. J OrthopRes 2001;19:210.

[11] Gruber HE, Fisher EC, Desai B, Stasky AA, Hoelscher G, Hanley EN. Humanintervertebral disc cells from the annulus: three-dimensional culture in

agarose or alginate and responsiveness to TGF-b1. Exp Cell Res1997;235:1321.

[12] Maldonado BA, Oegema TRJ. Initial characterization of the metabolism ofintervertebral disc cells encapsulated in microspheres. J Orthop Res1992;10:67790.

[13] Melrose J, Smith S, Ghosh P, Taylor TKF. Differential expression of proteoglycanepitopes and growth characteristics of intervertebral disc cells grown inalginate bead culture. Cells Tissues Organs 2001;168:13746.

[14] Wang JY, Baer AE, Kraus VB, Setton LA. Intervertebral disc cells exhibitdifferences in gene expression in alginate and monolayer culture. Spine2001;26:174751.

[15] Elisseeff JH, Anseth K, Sims D, McIntosh W, Randolph M, Langer R. Transdermalphotopolymerization for minimally invasive implantation. Proc Natl Acad SciUSA 1999;16:31047.

[16] Bryant SJ, Durand KL, Anseth KS. Manipulations in hydrogel chemistry controlphotoencapsulated chondrocyte behavior and their extracellular matrixproduction. J Biomed Mater Res A 2003;67:14306.

[17] Smeds KA, Pfister-Serres A, Miki D, Dastgheib K, Inoue M, Hatchell DL, GrinstaffMW. Photocrosslinkable polysaccharides for in situ hydrogel formation. JBiomed Mater Res 2001;54:11521.

[18] Burdick JA, Chung C, Jia X, Randolph MA, Langer R. Controlled degradation andmechanical behavior of photopolymerized hyaluronic acid networks.Biomacromolecules 2005;6:38691.

[19] Nettles DL, Vail TP, Morgan MT, Grinstaff MW, Setton LA. Photocrosslinkablehyaluronan as a scaffold for articular cartilage repair. Ann Biomed Eng2004;32:3917.

[20] Chung C, Mesa J, Miller GJ, Randolph MA, Gill TJ, Burdick JA. Effects of auricularchondrocyte expansion on neocartilage formation in photocrosslinkedhyaluronic acid networks. Tissue Eng 2006;12:266573.

[21] Ogushi Y, Sakai S, Kawakami K. Synthesis of enzymatically-gellablecarboxymethylcellulose for biomedical applications. J Biosci Bioeng2007;104:303.

[22] Chou AI, Bansal A, Miller GJ, Nicoll SB. The effect of serial monolayer passagingon the collagen expression profile of outer and inner anulus fibrosus cells.Spine 2006;31:187581.

[23] Chou AI, Reza AT, Nicoll SB. Distinct intervertebral disc cell populations adoptsimilar phenotypes in three-dimensional culture. Tissue Eng Part A

2008;14:207987.[24] Iwasa J, Ochi M, Uchio Y, Katsube K, Adachi N, Kawasaki K. Effects of celldensity on proliferation and matrix synthesis of chondrocytes embedded inatelocollagen gel. Artif Organs 2003;27:24955.

[25] Mauck RL, Seyhan SL, Ateshian GA, Hung CT. Influence of seeding density anddynamic deformational loading on the developing structure/functionrelationships of chondrocyte-seeded agarose hydrogels. Ann Biomed Eng2002;30:104656.

[26] Chang SCN et al. Injection molding of chondrocyte/alginate constructs in theshape of facial implants. J Biomed Mater Res 2001;55:50311.

[27] Puelacher WC, Kim SW, Vacanti JP, Schloo B, Mooney D, Vacanti CA. Tissue-engineered growth of cartilage: the effect of varying the concentration ofchondrocytes seeded onto synthetic polymer matrices. Int J Oral MaxillofacSurg 1994;23:4953.

[28] Vunjak-Novakovic G, Obradovic B, Martin I, Bursac PM, Langer R, Freed LE.Dynamic cell seeding of polymer scaffolds for cartilage tissue engineering.Biotechnol Prog 1998;14:193202.

[29] Soltz MA, Ateshian GA. Experimental verification and theoretical prediction ofcartilage interstitial fluid pressurization at an impermeable contact interface

in confined compression. J Biomech 1998;31:92734.

[30] Chou AI, Nicoll SB. Characterization of photocrosslinked alginate hydrogels fornucleus pulposus cell encapsulation. J Biomed Mater Res A 2008. doi:10.1002/

jbm.a.32191.[31] Orive G, Ponce S, Hernndez RM, Gascn AR, Igartua M, Pedraz JL.

Biocompatibility of microcapsules for cell immobilization elaborated withdifferent type of alginates. Biomaterials 2002;23:382531.

[32] Roughley P, Hoemann C, DesRosiers E, Mwale F, Antoniou J, Alini M. Thepotential of chitosan-based gels containing intervertebral disc cells for nucleuspulposus supplementation. Biomaterials 2006;27:38896.

[33] Di Martino A, Sittinger M, Risbud MV. Chitosan: a versatile biopolymer for

orthopaedic tissue-engineering. Biomaterials 2005;26:598390.[34] Liu Y, Shu XZ, Prestwich GD. Osteochondral defect repair with autologous bone

marrow-derived mesenchymal stem cells in an injectable, in situ, cross-linkedsynthetic extracellular matrix. Tissue Eng 2006;12:340516.

[35] Li Q, Williams CG, Sun DDN, Wang J, Leong K, Elisseeff JH. Photocrosslinkablepolysaccharides based on chondroitin sulfate. J Bio Mat Res A 2004;68:2833.

[36] Sa kai S , O gushi Y, Kawa ka mi K. Enz ymatica lly c rosslink edcarboxymethylcellulosetyramine conjugate hydrogel: cellular adhesivenessand feasibility for cell sheet technology. Acta Biomater 2009;5:5549.

[37] Pal K, Banthia AK, Majumdar DK. Development of carboxymethyl celluloseacrylate for various biomedical applications. Biomed Mater 2006;1:8591.

[38] Leone G, Fini M, Torricelli P, Giardino R, Barbucci R. An amidatedcarboxymethylcellulose hydrogel for cartilage regeneration. J Mater SciMater Med 2008;19:287380.

[39] Kim KS, et al. Electrostatic crosslinked in situ-forming in vivo scaffold for ratbone marrow mesenchymal stem cells. Tissue Eng Part A 2009. doi:10.1089/ten.tea.2008.0704.

[40] Elisseeff J , McIntosh W, Anseth K, Riley S, Ragan P, Langer R.Photoencapsulation of chondrocytes in poly(ethylene oxide)-based semi-interpenetrating networks. J Biomed Mater Res 2000;51:16471.

[41] Chung C, Erickson IE, Mauck RL, Burdick JA. Differential behavior of auricularand articular chondrocytes in hyaluronic acid hydrogels. Tissue Eng Part A2008;14:112131.

[42] Bryant SJ, Nuttelman CR, Anseth KS. Cytocompatibility of UV and visible lightphotoinitiating systems on cultured NIH/3T3 fibroblasts in vitro. J Biomater SciPolym Ed 2000;11:43957.

[43] Vistica DT, Skehan P, Scudiero D, Monks A, Pittman A, Boyd MR. Tetrazolium-based assays for cellular viability: a critical examination of selectedparameters affecting formazan production. Cancer Res 1991;51:251520.

[44] Nuttelman CR, Tripodi MC, Anseth KS. Synthetic hydrogel niches that promotehMSC viability. Matrix Biol 2005;24:20818.

[45] Mann BK, Schmedlen RH, West JL. Tethered-TGF-b increases extracellularmatrix production of vascular smooth muscle cells. Biomaterials2001;22:43944.

[46] Anseth KS, Bowman CN, Brannon-Peppas L. Mechanical properties of hydrogelsand their experimental determination. Biomaterials 1996;17:164757.

[47] Chou AI, Akintoye SO, Nicoll SB. Photo-crosslinked alginate hydrogels support

enhanced matrix accumulation by nucleus pulposus cells in vivo.Osteoarthritis Cartilage 2009.doi:10.1016/j.joca.2009.04.012.

[48] Cloyd JM, Malhotra NR, Weng L, Chen W, Mauck RL, Elliott DM. Materialproperties in unconfined compression of human nucleus pulposus, injectablehyaluronic acid-based hydrogels and tissue engineering scaffolds. Eur Spine J2007;16:18928.

[49] Miyanishi K et al. Effects of hydrostatic pressure and transforming growthfactor-beta 3 on adult human mesenchymal stem cell chondrogenesis in vitro.Tissue Eng 2006;12:141928.

[50] Mauck RL, Nicoll SB, Seyhan SL, Ateshian GA, Hung CT. Synergistic action ofgrowth factors and dynamic loading for articular cartilage tissue engineering.Tissue Eng 2003;9:597611.

[51] Byers BA, Mauck RL, Chiang IE, Tuan RS. Transient exposure to transforminggrowth factor beta 3 under serum-free conditions enhances the biomechanicaland biochemical maturation of tissue-engineered cartilage. Tissue Eng Part A2008;14:182134.

[52] Risbud MV et al. Toward an optimum system for intervertebral disc organculture: TGF-beta 3 enhances nucleus pulposus and anulus fibrosus survivaland function through modulation of TGF-beta-R expression and ERK signaling.

Spine 2006;31:88490.[53] Reza AT, Nicoll SB. Hydrostatic pressure differentially regulates outer and

inner annulus fibrosus cell matrix production in 3D scaffolds. Ann Biomed Eng2008;36:20413.

[54] Hutton WC, Elmer WA, Boden SD, Hyon S, Toribatake Y, Tomita K, Hair GA. Theeffect of hydrostatic pressure on intervertebral disc metabolism. Spine1999;24:150715.

[55] Hutton WC, Elmer WA, Bryce LM, Kozlowska EE, Boden SD, Kozlowski M. Dothe intervertebral disc cells respond to different levels of hydrostatic pressure?Clin Biomech 2001;16:72834.

[56] Kasra M, Goel V, Martin J, Wang S-T, Choi W, Buckwalter J. Effect of dynamichydrostatic pressure on rabbit intervertebral disc cells. J Orthop Res2003;21:597603.

[57] Neidlinger-Wilke C et al. A three-dimensional collagen matrix as a suitableculture system for the comparison of cyclic strain and hydrostatic pressureeffects on intervertebral disc cells. J Neurosurg Spine 2005;2:45765.

[58] Neidlinger-Wilke C et al. Regulation of gene expression in intervertebral disccells by low and high hydrostatic pressure. Eur Spine J 2006;15(Suppl.

3):S3728.

http://dx.doi.org/10.1002/jbm.a.32191http://dx.doi.org/10.1002/jbm.a.32191http://dx.doi.org/10.1089/ten.tea.2008.0704http://dx.doi.org/10.1089/ten.tea.2008.0704http://dx.doi.org/10.1016/j.joca.2009.04.012http://dx.doi.org/10.1016/j.joca.2009.04.012http://dx.doi.org/10.1089/ten.tea.2008.0704http://dx.doi.org/10.1089/ten.tea.2008.0704http://dx.doi.org/10.1002/jbm.a.32191http://dx.doi.org/10.1002/jbm.a.32191

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