Blocking of matrix metalloproteinases-13 responsive peptide in poly(urethane urea) for potential cartilage tissue engineering applications

Abstract

The matching of scaffold degradation rate with neotissue growth is required for tissue engineering applications. Timely provision of proper spaces especially for cartilage tissue engineering plays a pivotal role in chondrocyte cluster formation. In this study, poly(urethane urea) was synthesized using conventional two-stage method by extending the isocyanate group terminated prepolymers with different amounts of GPLGLWARK peptide, which responses the degrading induced by matrix metalloproteinase 13, the main proteinase for cartilage matrix degradation. The Fourier transform infrared spectrometer with the attenuated total reflection and 1H nuclear magnetic resonance spectra revealed that the peptides were introduced to poly(urethane urea) according to the characteristic absorption bands of the peptide and the newly formed urea bonds. The ultraviolet–visible spectroscopy spectra showed that the weight percentages of the peptide in the three poly(urethane urea) were 25%, 32%, and 35%. Atomic force microscopy images revealed that phase separation occurred in all poly(urethane urea) samples and became increasingly apparent with increasing amount of peptides introduced. Mechanical tests showed that the poly(urethane urea) strength increased with increasing amount of peptides in poly(urethane urea). Poly(urethane urea) proteolysis in matrix metalloproteinase 13 solution was more rapid than hydrolysis in aqueous buffer, and proteolysis rate was dependent on the amount of peptides in poly(urethane urea). Cell proliferation on the material surface in vitro displayed nontoxicity for all synthesized poly(urethane urea). In vivo subcutaneous implantation evaluation revealed the presence of local foreign body reactions triggered by poly(urethane urea) but was not due to peptide in poly(urethane urea). Moreover, the synthesized poly(urethane urea) with significant phase separation did not degrade under the matrix metalloproteinase 13 free subcutaneous environment, but poly(urethane urea) with minimal phase separation was degraded by attacking of the enzymes adsorbed on the hydrophobic surface through non-specific adsorption.

Keywords

Biodegradable polymers enzymatic degradation matrix metalloproteinase 13 tissue engineering

Introduction

Repairing cartilage defects remains challenging because connective tissues lack blood vessels, lymphatic tissues, and nervous tissues.¹ Cartilage tissue engineering is a promising approach for regenerating new cartilage on proper biodegradable scaffolds to restore the function of the defective cartilage and offers hope for patients waiting for organ repair.^2,3 The success of tissue regeneration relies on the scaffold properties to provide mechanical support, transfer biomechanical force, and deliver biological cues.⁴ The balance between degradation and strength is also a critical factor, given that their mismatch either slows down tissue ingrowth or accelerates scaffold collapse.⁵ Thus, for future applications, scaffolds should sense the growth of tissues and degrade according to the regeneration rate of the tissue.⁶ Although numerous studies on the improvement of scaffold bioactivity and strength are conducted, few studies have investigated on materials with responsive degradability.

The most commonly used degradable materials for tissue engineering scaffold are classified into natural (collagen, chitosan, hyaluronic acid, and elastin) and synthetic compounds (polyesters, polyanhydrides, polyphosphazenes, and their copolymers).^7–9 The degradation mechanism of these materials is a nonspecific chain scission process that is mainly controlled by hydrolysis, enzymolysis, and their combination. Degradation rate is generally predetermined by the chemistry of the macromolecules and cannot be changed once the material is fabricated.

Cell proliferation or migration requires spaces created by the degradation of collagens and proteoglycans in the extracellular matrix (ECM). This process is commonly conducted by proteases, such as matrix metalloproteinases (MMPs) and glycanase, which are secreted by cells to degrade specific sites on ECM molecules. Such protease reorganization molecules or structures could be introduced into synthetic materials to develop scaffolds with good mechanical strength and responsive degradation rate adaptive to cell growth for tissue engineering applications in the future.^10–12

Tissue engineering materials are fabricated by mixing ECMs with degradable polymers.^13,14 However, the blended materials exhibit the poor mechanical strength and unpredictable degradation rate of materials due to their heterogeneous structures caused by the weak intermolecular interactions between hydrophilic ECMs and hydrophobic polymers.¹⁵

Blocking protease sensitive molecular structures in polymer chains forms homogenous material structures with improved properties. For example, Benhardt et al. developed a poly(urethane urea) (PUU) by introducing the peptide segment GPQGIWGQGK into ether-based polyols as the primary MMP-2 responsive labile site for ligament tissue engineering application.¹⁶ Lee et al. incorporated GGGLGPAGGK into the backbone of polyethylene glycol (PEG)-based hydrogel to achieve 100% wet weight loss in collagenase at 37°C for 13 h and visualize collagenase activity.¹⁷ Guan et al. introduced the elastase-sensitive peptide AAK as chain extenders to increase the susceptibility of the material to enzyme-mediated degradation for potential soft-tissue engineering.¹⁸ Fonseca et al. grafted MMP-sensitive peptide PVGLIG to alginate hydrogel to promote extensive outward cell migration and invasion into the hydrogel.¹⁹ Paterson et al. prepared collagenase-degradable hydrogel by using a crosslinker containing the LGPA peptide sequence, showing the material’s potential as tissue engineering scaffold.²⁰ However, none of these sensitive molecular sequences display specificity for MMPs secreted by chondrocytes. In articular cartilage, MMP13 is the main protease continuously secreted by chondrocytes to degrade the major collagenous component of type II collagen in ECM.²¹ Previous work revealed that the peptide sequence GPLGLWAR could be selectively broken by MMP13 at the site between Gly and Leu.²² In the present study, a two-step method was used to block GPLGLWAR peptide segments into the elastic PUU chain to obtain a polymer with responsive degradation rate to MMP 13. PUU was selected because its mechanical properties may be easily tailored to be elastic for cartilage tissue engineering.²³ The synthesized PUU was investigated in terms of its chemical structures, tensile strength, degradation behavior, and toxicity both in vitro and in vivo.

Materials and methods

Materials

Hexamethylene diisocyanate (HDI) was supplied by TCI (Tokyo, Japan, 98%). A peptide composed by nine amino acids with the sequence of glycine–proline– leucine–glycine–leucine–tryptophan–alanine–arginine– lysine (GPLGLWARK) was purchased from Ningbo GL Biochemical (Ningbo, China, 98%). Polytetramethyleneglycol diol (PTMG, Mn = 1000 g/mol) was obtained from Sigma–Aldrich (Shanghai, China). Butanediol (BDO) and calcium hydride (CaH₂) were acquired from Aladdin (Shanghai, China). Stannous octoate was supplied by Alfa Aesar (Heysham, UK). MMP13 was provided by Worthington (Vassar Ave Lakewood, USA, 250 µ/mg). Dimethyl formamide (DMF), diethyl ether, and ethanol were obtained from Tianjin Baishi Chemical Reagent (Tianjing, China). DMF was dried with CaH₂, distillated under reduced pressure, and stored in a desiccator for not more than two weeks before use. PTMG was vacuum treated at 110°C overnight prior to use. The other reactants were used as received.

Mouse 129 teratocarcinoma AT805-derived chondrogenic ATDC5 cell line (the First Affiliated Hospital of the Sun Yat-Sen University), Dulbecco’s modified Eagle medium/Ham’s F12 (1:1, DMEM/F12, Hyclone, USA), fetal bovine serum (FBS, Hyclone, USA), and cell counting kit-8 assay (CCK8, Dojindo, Kumamoto, Japan) were used in cell experiments.

Synthesis of PUUs and sample preparation

PUU containing PTMG soft segment (SS) and HDI/GPLGLWARK peptide hard segment (HS) was synthesized by a modified two-step procedure.²⁴ PTMG and DMF were placed into a three-necked round bottom flask in an oil bath at 80°C under nitrogen protection. Stoichiometric HDI and stannous octoate were added into the flask and allowed to react for 2 h. The temperature was then decreased to room temperature and stoichiometric peptide/DMF (0.5 g/mL) solution was placed in the bottle for another 12-h block polymerization. The solution was added with ethanol and stirred for 30 min to quench the polymerization. Finally, the solution was precipitated into diethyl ether for further purification. The polymer was dried under vacuum for 24 h and stored in a desiccator. Three PUUs containing different amounts of peptide were synthesized (Table 1). Polyurethane (PU) with BDO as chain extender was also synthesized as control and termed as PUB.

Table 1.

Mole ratio and measured molecular weight of the synthesized polyurethane.

	PTMG	HDI	Peptide	BDO	HS (wt%)	Mn (Da)	Mw (Da)	polydispersity(PDI)
PU-P21	2	3	1	–	46	44,300	74,860	1.69
PU-P31	3	4	1	–	39	48,620	85,080	1.75
PU-P41	4	5	1	–	33	49,850	71,750	1.44
PUB	2	3	–	1	23	55,600	75,060	1.35

BDO: butanediol; HDI: hexamethylene diisocyanate; HS: hard segment; PTMG: polytetramethyleneglycol diol.

The film samples used for characterizations and toxicity experiments were prepared. The polymer/DMF solution (10%, w_t/v) was casted in a poly tetra fluoroethylene (PTFE) mold, heated it in an oven at 110°C for two days, and subjected to 24-h vacuum evaporation. Samples for cell and animal experiments were prepared by punching the film into 10-mm cyclic disks and small rectangular films (0.2 cm × 0.4 cm), respectively. The samples were sterilized by Gamma-ray irradiation under a dosage of 15 KGa before use.

Polymer degradation experiments

Hydrolysis and enzymolysis degradation were performed by immersing disk films into 5 mL plastic tubes containing 1 mL of tris(hydroxymethyl)methane-hydrochloric acid (TES) buffer (50 mM Tris-HCl, 10 mM CaCl₂, 150 mM NaCl, 0.02% NaN₃, pH 7.4) or buffer containing 0.1 mg/mL MMP13. The tubes were placed in a water bath rotary and shaken at 37°C up to 21 days. Samples were extracted at predetermined time intervals, and mass loss was calculated according to the equation: (W₀ − W₁)/W₀ × 100%, where W₀ is the initial weight and W₁ is the weight after degradation. For measuring the remaining weight, the samples were rinsed with deionized water for three times and dried at 50°C until a constant weight was obtained. Three parallel samples for each test point were used for measurement of mass loss.

Polymer characterizations

Gel permeation chromatography (GPC). Molecular weight was measured by a Viscotek GPC Max VE, 2001 series instrument equipped with a detector (Viscotek TDA 305 three detection combined system) and column (CLM, 3006 T6000M, Malvern) in tetrahydrofuran (THF) with a flow rate of 1 mL/min. Polystyrene standards were used for calibration. The samples were dissolved in THF (35 mg/mL), and the injection volume was 100 µL for the test.

Fourier transform infrared spectroscopy. Fourier transform infrared spectrometer with the attenuated total reflection (FTIR-ATR, Vector 33, Bruker, Germany) was utilized to determine the chemical structure of the polymer samples at 25°C. The wavenumber was scanned ranged from 600 cm⁻¹ to, 4000 cm⁻¹ with a step width of 4 cm⁻¹. Samples were randomly obtained from flat sheet membranes. ATR correction was used to enhance the absorption at high wavenumbers in the measurement.

Nuclear magnetic resonance (NMR) spectroscopy. ¹H NMR spectroscopy analysis of the polymer in deuterated dimethyl sulfoxide (DMSO-d6) was performed at 25°C with NMR spectrometer (400 MHz, AVANCE Digital, Bruker, Germany). The delay time for NMR tests was 2 s, and the scan number was 16 times with tetramethylsilane as internal reference.

Ultraviolet–visible spectroscopy (UV-Vis). The polymer/DMF solution (2 mg/mL) absorbance spectra within the wavelength range from 240 nm to 360 nm were collected by a UV-Vis spectrophotometer (TU-1901). Peptide amount in the final PUU was calculated according to the standard curve measured using the peptide/DMF solution with known concentration at a wavelength of 295 nm.

Polymer property tests

Water contact angle. The water contact angle was determined through sessile drop method with a water contact angle measurer (OCA15, DATAPHYSICS, Germany) at room temperature on PUU membrane. The droplet volume was 1 µL. Water contact angle was calculated according to the images captured by the equipment.

Atomic force microscopy (AFM) scanning. The phase morphology of the PUU membrane was obtained by AFM (MFP-3D-S, Asylum Research, America) in the dynamic force mode at room temperature. The phase shift angle under such mode suggested a SS darker contrast and an HS brighter contrast. A silicon cantilever with a nominal spring constant of 40 N/m, resonant frequency of 300 kHz, and tip radius of 10 nm was used for the tests.

Mechanical tests. Mechanical properties were tested by both tensile test and dynamic mechanic analysis (DMA). Dumb-bell samples (5 mm in width, 13 mm in length) cut from the PU films were tested using an INSTRON, 5967 at a crosshead speed of 5 mm/min. Mechanical properties were examined using the average value of measurements from five samples. Same-sized samples were evaluated for DMA tests by using a specific instrument (TA Q800 series DMA) over a temperature range of −120°C to100°C and a ramp rate of 3°C/min at a frequency of 1 Hz.

Scanning electron microscopy (SEM) observation. Film surface morphology images were captured using SEM instrument (Quanta 200, FEI, USA) at an accelerated voltage of 20 kV. Sample films were sputter coated with gold before SEM observation. Samples from cell experiments were treated by gradient dehydration before gold sputtering.

Cell proliferation on PUU surface

The ATDC5 cells were cultured in DMEM/F12 containing 5% FBS at 37°C under 5% CO₂. After 85% confluence, the cells were collected and subcultured until the required cell number was reached. The medium was changed every two days. Cell proliferation on films was tested by CCK8 method. Cyclic films were placed in a 24-well plate, and 10,000 cells were seeded per well. The cells were allowed to adhere within 5 h. The films were then extracted and transferred to another 24-well plate containing 400 µL fresh culture media per well. The test was performed at one, three, five, and seven days, and the absorbance (optical density) was recorded at 450 nm.

In vivo subcutaneous implantation

For animal experiments, six- to eight-week-old wild C57BL/6 mice were used. All experiments were performed in accordance with the guidelines of the Animal Experimentation Ethics Committee of the Southern Medical University. The animals were anesthetized by intraperitoneal injection of the mixture of ketamine and xylazine. The implant was placed under strict aseptic conditions. After cutting the dorsal skin, the PU films were implanted subcutaneously separately, and the skin was closed with a nylon suture. Mice were sacrificed at weeks 1, 2, 3, and 4 after implantation. The explanted subcutaneous tissue containing PU films were snap frozen, transversely cryosectioned (20 µm), and stained with hematoxylin and eosin. For immunofluorescence staining, the cryosections were fixed with cold acetone and incubated with rat anti-mouse CD11b (1:200, eBioscience). Rhodamine-conjugated goat anti-rat IgG (1:200, Santa Cruz) was used as secondary antibody. Nuclei were counterstained with 4',6-diamidino-2-phenylindole (DAPI). The slides were viewed under an Olympus BX51 fluorescence microscope (Olympus).

Statistics

Student’s t-test was conducted to analyze the differences among the groups (p < 0.05). All data are expressed as mean ± standard deviation.

Results

Chemical structure of synthesized PUU

A two-stage method was used to synthesize the PUU (Scheme 1) containing varied amounts of the GPLGLWARK peptide (Table 1). In the first step, HDI and PTMG were reacted to form prepolymer with isocyanate end groups. The amine end groups on the peptide reacted with the isocyanate terminal groups, forming urea bonds to extend the prepolymer chains and block the peptide segments within PUU. The amine group from the peptide exhibits higher nucleophilic affinity for the isocyanate group than the carboxyl and guadinol groups on the peptide; as such, the reaction preferentially occurs between amine and isocyanate groups. Table 1 shows the number–average molecular weight (Mn), weight–average molecular weight (Mw), polydispersity index, and calculated HS weight percentage of the synthesized PUU. The Mn for all PUU was higher than 40,000 Da, indicating that the polymer chains were extended by the peptide.

The results of FTIR-ATR and ¹H-NMR analyses display the characteristic molecular structures of PUU and PUB. In Figure 1, the PTMG SSs contained in both PUU and PUB exhibit C–O–C stretching vibration absorption band at the wavenumber of 1100 cm⁻¹ and their methylene stretching vibration bands at 2851 and, 2943 cm⁻¹. The characteristic urethane bonds in PUU and PUB show their N–H stretching vibration absorption band at approximately, 3336 cm⁻¹ and carbonyl stretching vibration band at 1720 cm⁻¹. The broad peak at approximately 3336 cm⁻¹ in PUU is due to N–H stretching vibration from the urea and amide bonds in PUU. The peptides in PUU also show their characteristic absorption bands of amide I carbonyl stretching vibration at, 1640 cm⁻¹, and the peak intensity increases with increasing amount of peptide incorporated in PUU. The band at 1675 cm⁻¹ is assigned to the carbonyl group due to urethane bonds involved in hydrogen bonding, which induces the band shifting to the low wavenumbers.

The ¹H-NMR spectra of the synthesized PUU are shown in Figure 2. The peaks at 3.3 and 1.5 ppm are assigned to the protons from PTMG methylene repeat units. The peaks at 1.1, 1.3, and 2.9 ppm are assigned to the protons of the hexamethylene adjacent to the urethane. The characteristic chemical shifts at 0.8, 4.2–4.5, 7.6–6.9, and 10.7 ppm are assigned to the protons of methyl, benzpyrole, guanidyl and amide, and carboxylic acid groups on the GPLGLWARK peptide, respectively. The peaks at 5.5–5.9 ppm show urea linkages are formed after the reaction between the peptide and the prepolymers.

Peptide amount in PUU and the phase structures

The absorption band at 295 nm in the UV-Vis spectra (Figure 3) is assigned to the tryptophan benzpyrole group of the GPLGLWARK peptide. The peptide is only present in PUU but not in PUB; as such, absorption is only observed among PUU samples. The peak intensity increases with increasing proportion of the peptide. The peptide weight percentages in PU-P41, PU-P31, and PU-P21 were calculated according to the standard curve (small figure in Figure 3) and were found to be 25.62%, 32.01%, and 35.13%, respectively. The peak shifts to short wavelength for the PUU containing low amounts of peptide because electron orbital transition is easily induced by the solvent in solutions of low concentrations.²⁵

The AFM phase images reveal the microphase separation in PUU and PUB. The high modulus HSs appear in light regions, whereas the low modulus SSs appear in the dark (Figure 3). The thermodynamic incompatibility between HS and SS is energetically favorable for them to avoid mixing. PUB with BDO as chain extender displays several regular structures in light area, with domain size of approximately 200 nm. For PUU, the domain size of light area increases with increasing amounts of peptide. The light area domain sizes for PU-P41, PU-P31, and PU-P21 are approximately 20, 100, and 300 nm, respectively.

Physical properties of PUU

Figure 4(a) shows the dynamic mechanical properties of the PU. The curves of storage modulus (solid line) for PUU and PUB decrease with temperature, whereas the loss modulus (dash line) shows two peaks on the curve. The small peak displayed on the curve of loss modulus at approximately −90°C is due to the energy dissipation caused by the motion of peptide pendent groups. Another peak at approximately −60°C is assigned to energy dissipation corresponding to glass transition (Tg) when the chain segment motion of SS fails to catch up with the changes in exterior force at the temperature. The moduli of PUU and PUB decreases remarkably when the temperature exceeds Tg and it deteriorates rapidly when the temperature exceeds softening point at 70°C.

The height of the peak displayed on the loss tangent curve (Figure 4(a)) provides a measure of the viscous SS portion to the elastic HS portion in the polymer. PU-P41 exhibits the highest SS to HS portion and the strongest peak intensity on the curve, whereas PU-P21 shows the weakest peak intensity. These findings are consistent with the test results from the UV-vis spectra. Figure 4(b) shows the PUU and PUB stress–strain curves in tensile strength tests. PU-P31 shows the longest stretch rate and PU-P21 display the highest tensile strength among all PUs.

According to the mechanical data displayed in Table 2, the tensile strength and the Young’s modulus increase with increasing amount of the peptide in PUU. PU tends to exhibit phase separation, where the SS units confer elastomeric behavior, and the HS units provide physical cross-linking. High HS (peptide segments) portion in PU-P21 enhances the phase separation and HS aggregation (Figure 3), resulting in improved mechanical properties (Figure 4(b)).

Table 2.

Properties of synthesized PUs.

Sample	Molecular weight	Water contact angle	Tg (°C)	Ts (°C)	Tensile strength (MPa)	Strain (%)	Young's modulus (MPa)
PU-P21	44,300	61 ± 3	−61.1	87.6	29.43 ± 3.99	823 ± 138	50.07 ± 5.04
PU-P31	48,620	65 ± 2	−60.2	75.3	10.49 ± 2.46	1120 ± 224	17.17 ± 4.20
PU-P41	37,850	79 ± 3	−59.6	–	4.16 ± 1.21	318 ± 68	8.42 ± 3.07
PU-B	55,600	83 ± 1	−68.5	80.4	7.43 ± 1.16	561 ± 164	49.12 ± 6.88
Cartilage	–	–	–	–	–	–	2.8–18.6^a25

^aThe statistical data are the “instantaneous” (loaded within 30 ms) compressive modulus of the human low limb joint specimens (ankle, knee, hip) of mortuary.

As shown in Table 2, the cartilage compressive modulus²⁵ has the same order of magnitude as the tensile modulus of the synthesized PU. Although the modulus for PU presented in the table is the tensile modulus, a correlation was observed between tensile and compressive moduli. The compressive modulus is higher than the tensile modulus, and their magnitudes are of the same order. Given that the PUU would be fabricated into porous scaffold, the mechanical properties should satisfy the applications to some extent.

Degradation property of PUU

Mass loss was observed among all PUU samples during degradation tests in enzymolysis solution. The mass loss increases with increasing amount of the peptide in PUU (Figure 5(a)). PU-P21, which has the highest amount of peptide, shows the fastest degradation rate because it offers the most abundant chain scission sites for MMP13 to target in the solution. The rate of mass loss in the first seven days is higher than that after seven days. This result may be caused by the high peptide content on the PUU surface during sample preparation. PUB shows low mass loss in the enzymolysis solution because of the migration of several dissolved small molecules into the solution.

Mass loss was also detected in the hydrolysis solution (Figure 5(b)). PUU samples display higher mass loss than PUB in the solution. This finding might be caused by both molecule dissolution and peptide hydrolysis. The peptide in the polymer increases the hydrophilicity of the materials (Table 2). The water contact angles are 61° ± 3° for PU-P21, 65° ± 2° for PU-P31, 79° ± 3° for PU-P41, and 83° ± 1° for PUB. The hydrophilic surface makes it easier for water to penetrate into the material interior resulting in the dissolution and degradation. The difference in the mass loss among PUU samples is not as significant as that in the enzymolysis solution. Thus, the phase separation is not uniform for all PUU samples.

The surface morphology of all the PUU and PUB films changed after 21-day degradation in both hydrolysis and enzymolysis solutions (Figure 5). Mass loss was compared among the films at the same time in the solution. The results are consistent with the extent of change in the surface morphology. Surface erosion was observed in PU-P21 immersed in the hydrolysis solution but not as intensive as that in the enzymolysis solution. PUB, the polymer without the peptide, displays the least changes in surface morphology in both solution. PU-P21 is broken into small pieces and particles at later degradation stage in the solution (data not shown).

Cell experiment and subcutaneous implantation

Figure 6 displays the cell proliferation on different substrates for one week. The cell numbers on all PUU and PUB samples increase with culture time (Figure 6(a)). Hence, the substrates are biocompatible for the adhering cells. Cell proliferation is not significantly difference among all PUU substrates. Cells show a more rounded shape and less stretched pseudopodium on the PUB surface than those on the PUU surface (Figure 6). This result could be due to the peptide offering possible cell binding sites offered by the peptide for adhesion.

Analysis of histological slices shows that after implantation, no hemorrhage or necrosis (septic or aseptic) occurred at the site. The PUU and PUB samples may undergo foreign body reactions and initiate mononuclear cell infiltration over four weeks (Figure 7(a)). Fluorescence detection further confirmed that PU polymers mainly initiated CD11b⁺ cell (monocytes/macrophages) invasion (Figure 7(b)). Despite the variation in the weight percentages of peptide segments in PUU, analysis of CD11b⁺ fluorescence intensity integrated optical density/area of interesting (IOD/AOI) indicates that those PUUs have triggered different levels of local responses compared with PUB (Figure 7(c)). Hence, the peptide is not a potential antigen and has no direct influence on the biocompatibility of the polymers in vivo.

We also investigated the in vivo biodegradation of the polymers implanted subcutaneously. We did not monitor the apparent degradation that occurred in the implants over four weeks, except for PU-P41. On the later stage of post-implant (day 28), degradation debris was observed on the PU-P41 film surface (Figure 7(a)). Through autofluorescence observation (wavelength of 530–550 nm), we found several tiny erosive poles scattered in PU-P41 (Figure 7(d)). This result is unexpected because the peptide content in PU-P41 is lower than that in PU-P21 and PU-P31. The MMP13 gene is mainly involved in cartilage and bone development, as such, its expression of MMP13 is insufficient in local subdermal microenvironment to decompose peptides in polymers. Thus, the degradation of PU-P41 in vivo was not exerted by MMP13 but could be due to the nonspecific degradation by several enzymes. In addition, the poor hydrophilicity (Table 2) and the less microphase separation of PU-P41 (Figure 3) could be more favorable for its in vivo nonspecific degradation. Before reaching the experimental stage at cartilage defect, the compositions must be precisely adjusted to minimize those adverse effects.

Discussion

Scientists incorporated the ECM degradable structures into synthesized PEG hydrogels to promote the chondrocyte function in vitro. For example, Yongdoo incorporated the GCRDGPQGIWGQDRCG peptides into PEG to form metalloproteinase-sensitive hydrogel; and demonstrated large cell clusters and more cells’ diffusion in the hydrogel.²⁶ Stacey blocked the CRDTEGE–ARGSVIDRC peptide, an aggrecanase-cleavable site in aggrecan, into PEG hydrogel; the result showed less hypertrophic cartilage in such hydrogels than that in nondegradable PEG hydrogel.²⁷ Hydrogel with cell proliferation sensitivity offers abundant available spaces and low restraint over cell clone cluster formation, there by promoting chondrocyte differentiation; however, but the enzyme specificity must be further improved for future applications.

MMP13 is the main and important cartilage ECM degradation enzyme secreted by chondrocytes.²⁸ In damaged cartilage, highly expressed MMP13 exhibits a substrate preference for collagen II over collagens I and III by recognizing a well-known specific site at Gly⁷⁷⁵–Leu⁷⁷⁶ in collagen II.²² Studies have investigated on MMP13 deficiency protecting C57BL/6 mice from antibody-induced arthritis; the results reveal the major presentation of such enzyme involved in matrix degradation.²⁹

In our research, we fabricated the PUU blocked with the GPLGLWARK peptide containing the amino acid combination of Gly–Leu, which may respond to MMP-13 enzymolysis in high kcat/Km value.^30,31 Traditional ester-bond-containing polymers are mainly degraded by water but exert few responses to local substances secreted by the cells at the implantation site; as such, PUU may show adjustable degradation rate in response to MMP13 secreted by chondrocytes after implantation and offers suitable spaces for tissue growth.

PUU can be easily fabricated into porous scaffold and offer improved and extended mechanical support for new tissue regeneration. Compared with chondrocytes on hydrogel, those on porous scaffold express similar secretion of glycosaminoglycan (GAG) and collagen type II but higher levels of collagen type I and collagen type X. However, both hydrogels and scaffolds show comparable chondrogenesis.³²

Further advantage for the enzyme-induced degradable material is the possible control over the material degradation time after implantation. Given that the enzyme activity could be neutralized by antagonists or promoted by activators through local injection or oral administration of those medicines, doctors may adjust the degradation rate according to the situation of patients. This convenient strategy is extremely attractive for controlling degradation of the implanted scaffold in vivo.

The adaptive degradation through the sensitive peptide segments blocked in PUU by MMP13 attack is expected in this research. However, both hydrolysis in vitro and nonspecific enzymolysis in vivo were detected. Further studies must be conducted to understand the detailed mechanism and factors (enzyme concentration, phase structure, molecular weight, and kinds and concentration of ions) affecting the expected ideal adaptive degradation.

After implantation, the peptide segments in PUU may act as antigens and induce attack on the immune system. In research on subcutaneous implantation, PUUs with varied amounts of peptide and PUB led to foreign body reactions at the implantation site. Hence, the peptide is not a major antigen with regard to the penetration of monocytes after implantation. This process is possibly triggered by other factors. Given that inflammation influences MMP13 expression level, relevant factors of PUU affecting the inflammation must still be investigated in the future.

Conclusion

We have successfully incorporated MMP13-sensitive peptide into PUU molecule. The FTIR and NMR results confirmed that the predicted characteristic structures were obtained. The peptide weight percentages in the PUU were calculated based on the UV-visible spectra and were found to be 25.6%, 32.0%, and 35.1%. The peptide incorporated into the molecule would not interfere in the microphase separation. The mechanical properties of the synthesized PUU were adjustable by modifying peptide content. The peptide was degraded in both hydrolysis and enzymolysis solutions but was determined to be more sensitive to the latter. Cell experiments show that the material supported the cell proliferation. Moreover, subcutaneous implantation confirmed that the peptide does not cause foreign body reactions in vivo. These experimental results indicate that PUU is one of the promising alternative degradable polymers for cartilage tissue engineering applications.

Footnotes

Authors’ Note

Gang Wu, Huan Wang, and Jiangwei Xiao contributed equally to this work. Correspondence to this article can also be referred to Hua Liao, Southern Medical University, No. 1023 Shatai south road, Baiyun District, Guangzhou, Guangdong 510515, China. Email: hua-liao@163.com.

Declaration of Conflicting Interests

The author(s) declared no potential conflicts of interest with respect to the research, authorship, and/or publication of this article.

Funding

The author(s) disclosed receipt of the following financial support for the research, authorship, and/or publication of this article: This research was financially supported by the NSFC Project (31470934, 81572102, 51572110), the key project of Guangdong Science and Technology Program (2014B010105007).

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