Device for repair surgery of cylindrical organs, particularly ruptured tendons, comprising a therapeutic agent for stimulating regrowth, and method of producing such device

Abstract

A device for repair surgery of cylindrical organs, particularly of ruptured tendons, is configured as a tubular sheath (T) made of a mesh of elastic fibers formed by electrospinning a biocompatible and biodegradable polymer. The tubular sheath has a Young elasticity modulus of about 0.1 to about 4 MPa and a strain at break of about 50 to about 1,000%, and it has a first wall surface and a second wall surface substantially parallel thereto, with said first wall surface being comparatively smooth (W.sub.S) and said second wall surface being comparatively rough (W.sub.R). According to the invention, the elastic fibers comprise first fibers consist of polymer in neat form and second fibers consist of polymer with an admixture of a therapeutic agent for stimulating regrowth processes of a predetermined cylindrical organ. The tubular sheath comprises a first tubular region adjacent to the first wall and a second tubular region adjacent to the second wall, said first tubular region being formed of said first fibers and said second tubular region being formed of said second fibers.

Claims

1. A device for repair surgery of cylindrical organs comprising a tubular sheath (T) made of a mesh of elastic fibers formed by electrospinning at least one biocompatible and biodegradable polymer, said tubular sheath having a Young elasticity modulus of about 0.1 to about 4 MPa and a strain at break of about 50 to about 1,000%, said tubular sheath having a first wall surface and a second wall surface substantially parallel thereto, said first wall surface being smooth (W.sub.S) and said second wall surface being rough (W.sub.R), wherein said elastic fibers comprise first fibers consisting of a first one of said polymers in neat form, and second fibers consisting of a second one of said polymers with an admixture of a therapeutic agent adapted to stimulate regrowth processes of a predetermined cylindrical organ, said tubular sheath comprising a first tubular region adjacent to said first wall and a second tubular region adjacent to said second wall, said first tubular region being formed of said first fibers and said second tubular region being formed of said second fibers.

2. The device according to claim 1, wherein said first polymer and said second polymer are the same polymer.

3. The device according to claim 1, wherein at least said second polymer is a biodegradable polyester urethane block copolymer with poly-hydroxy-butyrate as a hard segment and -caprolactone as a soft segment.

4. The device according to claim 3, wherein said soft segment has an average molecular weight of about 900 g/mol to about 1,250 g/mol and wherein the relative content of said soft segment is about 60 to about 75 parts by weight whereas the relative content of said hard segment is about 40 to about 25 parts by weight.

5. The device according to claim 1, wherein said tubular sheath has a Young elasticity modulus of about 0.4 to about 2.5 MPa and a strain at break of about 200 to about 1000%.

6. The device according to claim 1, wherein said tubular sheath is of substantially frustoconical shape.

7. The device according to claim 1, wherein said second fibers are heterogeneous filaments having included cavities filled with said therapeutic agent.

8. The device according to claim 1, wherein said second fibers are hollow filaments having a central core filled with said therapeutic agent.

9. The device according to claim 1, wherein said first tubular region and said second tubular region each form about one half of the sheath's wall thickness.

10. The device according to claim 1, wherein said therapeutic agent is selected from the group consisting of growth hormones, pharmaceutical agents and growth promoting cells.

11. The device according to claim 10, wherein said therapeutic agent is platelet-derived growth factorBB.

12. A method of producing a device according to claim 1, wherein said elastic fibers are formed by solution electrospinning.

13. The device of claim 1, wherein the device is adapted for the repair of a ruptured tendon.

14. The device of claim 10, wherein the growth promoting cells are stem cells.

15. Method for repairing surgery of cylindrical organs comprising providing the device of claim 1, pulling the device over a wound site, wherein the therapeutic agent stimulates the regrowth processes of the predetermined cylindrical organ.

16. The method of claim 14, wherein the cylindrical organ is a tendon.

Description

BRIEF DESCRIPTION OF THE DRAWINGS

(1) The above mentioned and other features and objects of this invention and the manner of achieving them will become more apparent and this invention itself will be better understood by reference to the following description of various embodiments of this invention taken in conjunction with the accompanying drawings, wherein are shown:

(2) FIG. 1 Single and emulsion electrospinning of DegraPol. A) Structure of the co-block-polymer DegraPol consisting of a hard and a soft segment joined by isocynate junction unit. B) During the process of single electrospinning, only simple polymer fibers are obtained. C) Emulsion electrospinning allows for incorporation of biomolecules in an aqueous phase, into the fibers, resulting with either randomly distributed molecules within the fiber (bottom) or core-shell architecture of the fiber (top).

(3) FIG. 2 Schematic overview of the design of the bioactive DegraPol tube and its anticipated clinical application for tendon repair. A) Reversibly expendable DP tube, made from a new highly elastic DP. B) Design of single electrospun (pure) DP tube (one-layer) and bioactive electrospun DP tube (two-layered), incorporating PDGF-BB in the emulsion electrospun layer. C) Application of the bioactive DP tube over a conventionally sutured ruptured site of a tendon, delivering PDGF-BB at the wound site. D) Enlarged schematic cross-sectional view of bioactive DegraPol tube with flat first wall surface (bottom) and rough second wall surface (top) with groove-type features.

(4) FIG. 3 Characterization of polymer emulsions and scaffolds produced with different wt % of DegraPol. A) Fluorescence microscopy images of 8, 10 and 12 wt % DP emulsions with FITC-BSA dissolved in Tris-buffer (pH 7.4) as aqueous phase, prepared before electrospinning. B) Average diameter [m] of droplets and C) average circularity of droplets in each DP-emulsion prepared. D) SEM micrographs of resulting single electrospun DP fibers from only polymer solution (pure DP). E) SEM micrographs of resulting emulsion electrospun DP fibers, at different wt % of DP polymer in the polymer solution used for electrospinning. (***p<0.001.) Scale bars: A), D) and E) 20 m.

(5) FIG. 4 Effect of emulsion electrospinning and electrospinning parameters on fiber diameter of DegraPol fibres. A) Effect of different electrospinning parameters (flow rate [ml h.sup.1], voltage [kV]) on fiber diameter of single and emulsion electrospun DP fibers, using different wt % of DP polymer solution. B) Fiber diameter as a function of different wt % of DP during single and emulsion electrospinning. (*p<0.05, ***p<0.001.)

(6) FIG. 5 Anisotropy in mechanical properties of single and emulsion electrospun DegraPol tubes. A) Appearance of single (pure) and emulsion electrospun DP tubes. B) SEM micrographs of the top surface of single and emulsion electrospun DP tubes, showing tertiary structure formation by fiber fusion. C) Ultimate tensile stress (UTS) [MPa] and D) strain at break [%] of DP tubes measured in the axial direction. E) Ultimate tensile stress (UTS) [MPa] and F) strain at break [%] of DP tubes measured in the transverse direction. G) Young's modulus [MPa] of 8, 10 and 12 wt % DP scaffolds, single and emulsion electrospun. (*p<0.05, ** p<0.01, ***p<0.001.) Scale bars: B) 150 m.

(7) FIG. 6 Characterization of DegraPol emulsion electrospun scaffolds as a delivery system using two model molecules. A) CLSM images of emulsion electrospun DegraPol fibers with incorporated FITC-BSA at different flow rates (ml h.sup.1) and wt % of DegraPol in the polymer solution. B) In vitro cumulative release [%] of fluorescein from DegraPol scaffolds as a function of flow rate used (1 ml h.sup.1 and 3 ml h.sup.1, 10 wt % DP). C) In vitro cumulative release [%] of fluorescein from DegraPol scaffolds as a function wt % of DP polymer solution used, at a constant flow rate (1 ml h.sup.1). D) In vitro cumulative release [%] of FITC-BSA from DegraPol scaffolds as a function of flow rate used (1 ml h.sup.1 and 3 ml h.sup.1, 10 wt % DP). E) In vitro cumulative release [%] of fluorescein from DegraPol scaffolds as a function wt % of DP polymer solution used, at a constant flow rate (1 ml h.sup.1). Scale bars: A) 10 m.

(8) FIG. 7 PDGF incorporation within DegraPol scaffolds and its release profile from emulsion electrospun DegraPol scaffolds. A) CLSM images of physically adsorbed PDGF-BB on pure DP scaffolds detected by immunostaining methods. B) Emulsion electrospun DP scaffolds with incorporated PDGF-BB. C) In vitro cumulative release of PDGF-BB [pg ml.sup.1]. D) In vitro cumulative release of PDGF-BB [%] calculated from theoretical loading. E) pg of PDGF-BB released per mg of DP scaffold in in vitro conditions. Electrospinning parameters: C), D) and E) 1 ml h.sup.1, 11-12.5 kV, 12 wt % DP. Scale bars: A) and B) 20 m.

(9) FIG. 8 Effect of PDGF-BB on rabbit tenocytes and bioactivity of incorporated and released PDGF-BB. A) CLSM images of screening range of PDGF-BB concentrations (ng ml.sup.1) for proliferative effect on Achilles tendon rabbit tenocytes in serum free conditions using EdU proliferation assay. B) CLSM images of screening range of PDGF-BB concentrations (ng ml.sup.1) for proliferative effect on Achilles tendon rabbit tenocytes in serum.sup.+ conditions using EdU proliferation assay. C) EdU-positive cells [%] for each PDGF-BB concentration in serum free condition. D) EdU-positive cells [%] for each PDGF-BB concentration in serum.sup.+ condition. E) Bioactivity of released PDGF-BB from DP scaffolds in serum free condition, expressed in EdU-positive cells [%]. F) Bioactivity of incorporated PDGF-BB on DP scaffolds in serum.sup.+ condition, expressed in EdU-positive cells [%]. (** p<0.01, ***p<0.001.) Scale bar: A), B), E) and F) 100 m.

(10) FIG. 9 Biomechanical properties of extracted rabbit Achilles tendons three weeks post-surgery. In both groups, an electrospun DegraPol tube was implanted around the fully transsected and sutured tendon, however, in one group (denoted by DegraPol Tube+PDGF-BB), the growth factor PDGF-BB was incorporated. As can be concluded, the growth factor affected ultimate failure load as well as ultimate stress of the corresponding specimen: They were higher compared to the tendon treated without PDGF-BB.

(11) FIG. 10 Biomechanical properties of extracted rabbit Achilles tendons three weeks post-surgery. In a first group (left, denoted by DegraPol Tube), a DegraPol tube was implanted around the fully transsected and sutured tendon; in a second group (middle, denoted by DegraPol Tube+PDGF-BB), the growth factor PDGF-BB was incorporated in the tube; in a third group (right, denoted by Control, no tube was applied. Black bars (left side of each group) show data from experiments conducted to test emulsion electrospun tubes whereas grey bars (right side of each group) show data from experiments conducted to test coaxially electrospun tubes. The measured properties were: (a) tendon length in mm, (b) cross sectional area (CSA) in cm.sup.2, (c) load until failure in N, (d) failure stress im MPa, (e) stiffness in N/mm, and (f) elastic modulus in MPa.

DETAILED DESCRIPTION OF THE INVENTION

(12) The device for repair surgery of cylindrical organs, particularly of ruptured tendons, is shown in FIGS. 1 and 2. It is configured as a tubular sheath (T) made of a mesh of elastic fibers formed by electrospinning a biocompatible and biodegradable polymer. The tubular sheath T is approximately cylindrically shaped along a longitudinal axis A and has a first wall surface, which in the example shown in FIG. 2 is the outer tube surface, and a second wall surface substantially parallel thereto, which in the example shown is the inner tube surface. The configuration shown in FIGS. 2A to 2C is the one to be used for the intended surgical application. This configuration is obtained from a freshly produced electrospun tubular sheath, as shown in FIG. 2D, by everting the same, i.e. by switching the inner and outer side thereof. The first, outer wall surface is comparatively smooth (W.sub.S) whereas the second, inner wall surface is comparatively rough (W.sub.R), thus allowing good adhesion to the ruptured tendon surface. A first tubular region adjacent to the first, outer wall surface is formed of first fibers consisting of the polymer in neat form. A second tubular region adjacent to the second, inner wall surface is formed of second fibers consisting of the polymer with an admixture of a therapeutic agent for stimulating tendon regrowth processes. The tubular sheath as a whole has a Young elasticity modulus of about 0.1 to about 4 MPa and a strain at break of about 50 to about 1,000%.

(13) As shown in FIG. 2B at the right hand side, the boundary region Z denoted with an arrow is not necessarily a sharp transition between the two types of fibers. Depending on how the electrospinning process is carried out, there may be a gradual transition between the two fiber types if the polymer solution in the reservoir of the electrospinning device is gradually changed from a solution of the neat polymer to a solution also containing the therapeutic agent.

Example 1

(14) Several electrospinning methods for production of bioactive scaffolds include physical adsorption of biomolecules onto scaffolds, electrospinning of emulsions that contain the biomolecules of interest, i.e. emulsion electrospinning or coaxial electrospinning, thus creating an external polymer shell filled with aqueous core carrying biomolecules..sup.[10] Major difficulties using these different methods for biomolecule incorporation are achieving controlled release profiles of the molecules and maintaining their bioactivity. Among these, emulsion electrospinning at least offers easy incorporation of molecules in the scaffold, with a more sustained release profile, compared to a burst release exhibited with physical adsorption methods..sup.[11] Thus, for production of bioactive DP scaffolds emulsion electrospinning was used. Successful incorporation of lysozyme, Rhodamine B, bovine serum albumin (BSA) and some growth factors (NGF and PDGF-BB) into PLGA, PCL and PLLA polymer fibers by emulsion electrospinning has been reported..sup.[12] However, no previous research has been done exploiting DP as a drug delivery system, thus careful morphological and mechanical characterization of the produced scaffolds is necessary to determine whether changes in scaffold production and design affect the scaffold's mechanical and delivery device properties.

(15) The goals of this study were (i) detailed investigation of the impact of electrospinning parameters (voltage, flow rate, weight % of polymer solution) on the morphology (fiber diameter) of the new DP scaffolds; (ii) comparison of pure and emulsion electrospun DP tubes in terms of morphological and mechanical properties; (iii) characterization of DP scaffolds as delivery devices in terms of release kinetics of two model compounds (fluorescein (FL) and fluorescein isothiocyanate-conjugated bovine serum albumin (FITC-BSA)), as well as PDGF-BB, the molecule of interest and (iv) assessing PDGF-BB bioactivity and effect upon release on rabbit tenocytes in in vitro conditions.

EXPERIMENTAL SECTION

(16) Incorporation of Fluorescein and FITC-BSA in DP Polymer Solutions:

(17) DegraPol (Ab medica, Italy) polymer solutions with 8, 10, 12, 14 or 16 wt % were prepared by dissolving the polymer overnight at room temperature in a mixture of chloroform/HFP (80:20 w/w) (Sigma). For emulsion preparation, fluorescein (M.sub.w=376.27 g mol.sup.1, Polysciences, Inc., USA) was dissolved in MilliQ water at 5 mg mL.sup.1 and FITC-BSA (Sigma) was dissolved in 10 mM Tris-HCl buffer (pH=7.4) at a concentration of 5 mg mL.sup.1, each representing the aqueous phase of the emulsion. Water-in-oil emulsions were prepared by drop-wise addition of 200 L of aqueous phase to the polymer solution while stirring at 500 rpm for 2 minutes. Afterwards the mixtures were sonicated using a probe ultrasonicator (Sonopuls HD 2070, Bandelin, Germany) for 2 minutes at 50% amplitude. All the procedure was done immediately prior to electrospinning. For emulsion characterization prior to electrospinning, few drops of emulsions of different wt % DP polymer solutions and FITC-BSA were mounted on a glass slide and observed under fluorescence microscope (Zeiss Axiovert 200M; Carl Zeiss, Germany). Images were analyzed for average diameter and circularity of emulsion droplets and presented as meanstandard deviations (n=3). Circularity (4[area]/[perimeter].sup.2) ranges from 0 (infinitely elongated polygon) to 1 (perfect circle).

(18) Incorporation of PDGF-BB in DP Polymer Solution:

(19) Recombinant human PDGF-BB (PeproTech) was diluted in 0.1% BSA in DI water at a concentration of 40 g mL.sup.1. 200 l of this, containing a total of 8 g PDGF-BB, were added drop-wise to 5 g of DP polymer solution, under stirring for 2 minutes at 500 rpm. Afterwards the mixture was sonicated with a probe ultrasonicator for 2 minutes at 50% amplitude. Immediately afterwards the emulsion was used for electrospinning.

(20) Scaffold Production by Electrospinning:

(21) In-house assembled electrospinning device was used, consisting of a spinning head with a blunt end made of stainless steel tube (1 mm inner diameter and 0.3 mm wall thickness, Angst & Pfister AG, Zrich, Switzerland), a DC high voltage supply (Glassman High Voltage Inc., High Bridge, N.J., USA), hollow cylindrical aluminum mandrel as a collector and a syringe pump (SP210cZ, WPI, Germany). The DP polymer solutions or emulsions created were loaded in 2 mL syringe (B. Braun Melsungen AG, Germany) and pumped into the spinning head. The electrospinning conditions were varied by changing the concentration of polymer in the solution (8, 10, 12, 14 and 16 wt %), the flow rate (0.5-3 ml h.sup.1) used and the voltage (10-17.5 kV) applied between the spinning head and the collector. The distance between the spinning needle and the collector (working distance) was kept constant at 15 cm. The process was done at room temperature (22-23 C.) and 35% humidity. For DP tube production, same electrospinning device was used, where a cylindrical rotating aluminum mandrel (length: 200 mm, diameter: 4 mm) was used as a collector. Pure or emulsion DP polymer solutions were used for DP tube production, using electrospinning parameters of 11-12.5 kV, 1 ml h.sup.1 flow rate and 20 cm distance from the collector, under 35% humidity and room temperature. In addition, the spinning head through which the polymer solution was ejected was moving left and right in a range of 5 cm to facilitate equal deposition of fibers on the rotating collector. Spun tubes were removed from the collector with 50% ethanol and washed with water and dried under vacuum at room temperature.

(22) PDGF-BB Adsorption on DP Scaffolds:

(23) PDGF-BB was physically adsorbed on pure DP scaffolds. Electrospun scaffolds were cut into 1.5 cm1.5 cm pieces and placed in low-binding micro tubes (Sarstedt). 1 mL of 300 ng mL.sup.1 PDGF-BB (in 0.1% BSA in 1PBS) was added to each tube and incubated with the scaffolds for 24 hours at 4 C. Afterwards, the scaffolds were rinsed in distilled water and used for further experiments.

(24) Scanning Electron Microscopy (SEM):

(25) Electrospun scaffolds were dried in vacuum overnight and then samples were mounted on metal stubs with conductive double-sided tape. Samples were sputter coated (SCD500, Bal-tec) with platinum in order to obtain 10 nm coating and then examined by SEM (Zeiss SUPRA 50 VP, Zeiss, Cambridge, UK) at an accelerating voltage of 5 kV. Fiber diameters of each sample were measured using SEM images and the image analysis software platform Fiji. First a diagonal line was drawn on the image, and the fiber diameter of the fibers was measured perpendicular to the fiber length at the points where the diagonal line crossed the fibers. The measurement was done manually with the measurement tool in Fiji, after calibration with the scale bar of the microscope image. Three images of each scaffold were analyzed, with an average of 30 counts per scaffold (n=10/image). Fiber diameters are given as averagesstandard deviations.

(26) Confocal Laser Scanning Microscopy (CLSM):

(27) Emulsion electrospun DegraPol scaffolds with incorporated fluorescein or FITC-BSA were analyzed with confocal microscopy in order to visualize the presence and distribution of the molecules within the electrospun fibers. A thin layer of different emulsion electrospun DP fibers was collected on glass coverslips and then observed by CLSM (SP5, Leica Microsystems, Wetzlar, Germany). The excitation wavelength for both fluorescein and FITC-BSA was 488 nm, and images were taken with 63x/1.4 NA objective.

(28) Mechanical Testing of Scaffolds:

(29) The mechanical properties of pure and emulsion electrospun DP tubes (8, 10 and 12 wt % DP) were obtained from stress/strain curves measured using a uniaxial load test machine (Instron tensile tester, High Wycombe, Buck, UK: model 5864) equipped with 10 N load cell. The mechanical properties of the tubes were measured in two directions, the axial and transverse direction of the tube. Standard dog-bone shaped samples with a testing region of 122 mm.sup.2 and thickness range of 500-800 m were punched out from the tubes in the longitudinal direction and an elongation rate of 12 mm/min was applied until failure. Rectangular samples with a testing region 122 mm.sup.2 were cut in the transverse direction and elongation rate of 12 mm/min was applied until failure. The Young's modulus [MPa], strain at break (%) and tensile strength [MPa] were determined for every direction in every condition (n=3).

(30) In Vitro Release of Fluorescein, FITC-BSA and PDGF-BB from DP Scaffolds:

(31) Emulsion electrospun scaffolds incorporating fluorescein or FITC-BSA were cut in pieces (20-30 mg) and shortly wetted in 50% ethanol and rinsed in MilliQ water. The scaffolds (n=9) were placed in micro test tubes and 1 mL of release medium was added. The release of fluorescein was performed in MilliQ water, while that of FITC-BSA was performed in 10 mM Tris-HCl buffer (pH 7.4). Samples were incubated at 37 C. and 5% CO.sub.2 under mild shaking. At each respective time point (t=12, 24, 48, 72, 120, 168, 336 hours) samples were taken out and placed in new micro test tubes and 1 mL of fresh release medium was added. The concentration of fluorescein or FITC-BSA in the release medium was determined by measuring the fluorescence of each sample by a fluorescence plate reader (Infinite200, Tecan, Switzerland) at 485 nm excitation and 540 nm emission. The concentration of the samples was calculated based on standard curves of known concentration of fluorescein and FITC-BSA in each respective release medium used. At the end of the release time period, extraction of not released molecules from the DP scaffolds was performed. The extraction of the scaffolds was done by dissolving the scaffold in 1 mL of chloroform by shaking at 800 rpm at room temperature, followed by addition of 1 mL of aqueous phase (MilliQ water for fluorescein and 10 mM Tris-HCl buffer for FITC-BSA). The two phases were mixed together overnight and afterwards the samples were centrifuged at 10 000 rcf for 30 minutes. The concentration of fluorescein or FITC-BSA was measured in the supernatant with fluorescence plate reader. The total loaded amount of molecule was the determined extracted amount added to the cumulative release at the end time point. The results are presented as cumulative release [%] as function of time and calculated by Equation (1):

(32) $\begin{array}{cc}\mathrm{Cumulative}\mathrm{release}\left[\%\right]=\left(\frac{M_r}{M_t}\right)100& \left(1\right)\end{array}$
where M.sub.r is the amount of released molecule at time t and M.sub.t is the total released amount of molecule from the scaffold at the end time point plus the extracted leftover amount of molecule from the scaffold upon release.

(33) Emulsion electrospun DP scaffolds incorporating PDGF-BB were cut in pieces (20-30 mg), placed in low-binding micro tubes (Sarstedt) and 500 L of 0.1% BSA in 1PBS were added as release medium. Samples (n=9) were incubated at 37 C. and 5% CO.sub.2 under mild shaking. At each respective time point (t=1, 2, 3, 5, 7, 10, 14, 21 and 30 days) samples were taken out and placed in new micro test tubes and 500 L of fresh medium were added. Release samples were stored at 20 C. until further quantification and bioactivity assays. Human recombinant PDGF-BB was quantified with PDGF-BB ELISA kit (PeproTech) according to manufacturer's protocol and samples were measured with an absorbance plate reader (Infinite200, Tecan, Switzerland) at 405 nm and correction set at 650 nm. Release of PDGF-BB is represented as cumulative release over time in pg ml.sup.1, cumulative release [%] based on theoretical loading or normalized to the weight of DP scaffolds as pg mg.sup.1 of DP scaffold.

(34) PDGF-BB Detection on Electrospun DP Scaffolds:

(35) Physically adsorbed PDGF-BB on pure DP scaffolds or incorporated on emulsion electrospun DP scaffolds was detected by immunostaining. Samples were washed with 1PBS, blocked with 3% BSA (Sigma) in 1PBS for 1 hour at room temperature, washed 3 with 1PBS and incubated with 1 g ml.sup.1 rabbit anti-PDGF-BB polyclonal antibody (PeproTech) overnight at 4 C. The samples were then washed 3 with 1PBS and incubated with 10 g ml.sup.1 AlexaFluor488 goat anti-rabbit (Life Technologies) for 1 hour at room temperature. Afterwards, samples were washed 3 with 1PBS and mounted with mounting medium (DAKO) and imaged with confocal microscope (SP5, Leica Microsystems, Wetzlar, Germany). 408 excitation/455 emission was used for autofluorescence of the DP fibers and 498 excitation/519 emission for secondary antibody visualization. For negative control, scaffolds were incubated with only secondary antibody in the presence of PDGF-BB or with primary and secondary antibodies in the absence of PDGF-BB to detect any non-specific binding of the antibodies to the scaffolds.

(36) Cell Culture:

(37) Rabbit tenocytes were isolated from Achilles tendons of New Zealand White rabbits. Briefly, Achilles tendons were cut out and washed with DPBS (Biowest) with 200 g mL.sup.1 gentamycin (Biowest) and 2.5 g mL.sup.1 amphotericin B (Biowest). Small pieces from the central part of the tendons were cut with a scalpel, put in digestion solution (Ham's F12 (Biowest), 200 gentamycin, 2.5 g ml.sup.1 amphotericin B and 3 mg ml.sup.1 collagenase II (147 U mg.sup.1, PAN Biotech) and left to react between 12 and 18 hours, at 37 C. and 5% CO.sub.2, without shaking. The samples were centrifuged at 400 g for 5 minutes. The pellet was resuspended and added directly to 75 cm.sup.2 tissue culture flask with 12 ml of culture medium (Ham's F12, 10% FBS (Biowest) and 50 g mL.sup.1 gentamycin). After 24 hours, the adherent cells were 3 washed with DPBS and fresh 25 ml culture medium were added. In 4 to 5 days, tenocytes migrated out and formed monolayer and were cryopreserved before use. Cryopreserved rabbit tenocytes were thawed and resuspended in culture medium (Ham's F12 with 10% FBS and 1% Penicillin/Streptomycin (P/S) (Life Technologies)). Tenocytes between passages 1 and 4 were used for all experiments.

(38) In Vitro Bioactivity Assays:

(39) The effect of PDGF-BB on rabbit tenocytes was tested by increase in cell proliferation assessed by Click-iT EdU proliferation kit (Life Technologies). The proliferative effect was studied in serum.sup.+ (Ham's F12 with 10% FBS and 1% P/S) and serum free (Ham's F12, 1RPMI vitamins solution (Sigma), 1 non-essential amino acids solution (Life Technology) and 1% P/S) culture medium and PDGF-BB concentrations of 1-50 ng mL.sup.1 were tested. Cells were seeded in 8-well slides, ibiTreat (Ibidi) at 410.sup.4 cells mL.sup.1 (300 L per well) in serum free medium. Cells were cultured in these conditions for serum starvation synchronization with daily change of serum free medium. After 3 days, different PDGF-BB concentrations in serum free or serum.sup.+ medium, respectively (1, 3, 5, 10, 25, 50 ng mL-1 for serum free condition and 10, 25, 50 ng mL.sup.1 for serum.sup.+ condition), together with 10 M 5-ethynyl-2-deoxyuridine (EdU) were added and incubated for 24 hours at 37 C. and 5% CO.sub.2. Cells were then fixed, permeabilized and EdU stained according to the kit's protocol. Cell nuclei were stained with 5 g mL.sup.1 46-diamidino-2-phenylindole dilactate (DAPI) for 10 minutes. Samples were imaged with confocal microscope and 30 random images were taken per sample (n=6). Results are expressed as EdU-positive cells [%]standard deviations.

(40) For determining of the bioactivity of the released PDGF-BB from the emulsion electrospun DP scaffolds, the aliquots from each release experiment respectively were pooled together and concentrated using 10K Amicon Ultra-0.5 Centrifugal Filter units (EMD Millipore) according to manufacturer's protocol. The concentrated samples were diluted in serum free medium containing 10 M EdU, added to serum starved synchronized tenocytes and incubated for 24 hours. Afterwards, the cells were fixed, permeabilized and EdU stained and the percentage of EdU-positive cells was determined.

(41) For determining PDGF-BB bioactivity directly on DP scaffolds, tenocyte proliferation on pure DP, emulsion electrospun DP with PDGF-BB, empty emulsion electrospun DP without PDGF-BB (Emulsion DP+BSA) and pure DP with physically adsorbed PDGF-BB was studied. Scaffolds (1.51.5 cm) were placed in 24-well tissue culture plates and UV-sterilized for 30 minutes. Scaffolds were 1 washed with culture medium and tenocytes (110.sup.5 cells/scaffold) were seeded in serum.sup.+ medium. After overnight culturing of the cells on the scaffolds, the medium was replaced with a fresh one containing 10 M EdU and samples were incubated for additional 24 hours. The samples were fixed, permeabilized and EdU stained according to the kit's protocol. Scaffolds were mounted on microscope slides with mounting medium (DAKO) and imaged with confocal microscope. 20 random images were taken per scaffold (n=8) and the percentage of EdU-positive cells was determined for each sample.

(42) Statistics:

(43) Data was analyzed with Origin (OriginLab, Northampton, Mass.). Values are expressed as meansstandard deviation. One-way analysis of variance (one-way ANOVA) was performed to test the differences between groups in all the experiments, using comparison post-hoc test for significance. p values of less than 0.05 were considered statistically significant results and are indicated with an asterisk within graphs (*p<0.05, **p<0.01, ***p<0.001).

(44) Results and Discussion

(45) Bioactive DP Scaffold Design

(46) To produce bioactive DP scaffolds that can release PDGF-BB and promote the healing process of tendon rupture repair, we used a newly synthesized, highly elastic type of DP (FIG. 1A). Being slightly different from the standard single electrospinning (FIG. 1B), emulsion electrospinning (FIG. 1C) was chosen as a method for incorporation of biomolecules inside DP electrospun fibers. This new type of DP allows for production of extendable, elastic electrospun tubes that can be pulled over an injured tendon (FIG. 2A). These elastic tubes are surgeon-friendly and easily applicable in clinical settings. Similar to previously tested pure DP scaffold.sup.[4] (FIG. 2B), the bioactive DP scaffold was designed to be in the form of a two-layered tube, having a rough electrospun inner layer delivering PDGF-BB at the site of injury and a smooth outer layer, consisting of only pure electrospun DP (FIG. 2B). This DP tube is intended to be applied over the wound of conventionally sutured ruptured tendon (FIG. 2C)..sup.[4] Having two-layered tube will allow local delivery of PDGF-BB at the site of injury, however, having a protective layer of pure DP on the outside should limit diffusion of PDGF-BB to the surrounding tissue causing cell migratory effect.sup.[13] and not desired adhesion formation..sup.[6]

(47) Morphology and Physical Properties of Emulsion Electrospun DP Scaffolds

(48) As previously not used in emulsion electrospinning and not explored as a drug delivery device, several characterizations of emulsion electrospun DP scaffolds were performed in terms of wt % of DP solution to be used, initial emulsion formation and morphology of obtained scaffolds.

(49) Water-in-oil emulsions were prepared, whereby the organic phase consisted of DP polymer solutions (8, 10, 12 wt %), while the aqueous phase contained FITC-BSA in Tris-buffer. Fluorescence microscopy revealed that these emulsions were homogenous in appearance when using 8-12 wt % DP polymer solutions (FIG. 3A). 8 wt % and 10 wt % emulsions had droplets with a diameter of 9.411.45 m and 9.370.68 m, while 12 wt % emulsions experienced a significant decrease in the diameter of the droplets down to 5.050.35 m, thus resulting in more fine emulsions (FIG. 3B). Average circularities of the droplets in each emulsion did not show significant differences between conditions and were 0.650.03 on average (FIG. 3C). Classification after Pal et al., defines emulsions with droplet side of 4-12 m as fine and 25-30 m as coarse..sup.[14] Thus, the DP emulsions used can be classified as fine emulsions, with droplet size of 5-10 m and without big phase separations. This suggests that a 0.4% of aqueous phase content is well suited for the preparation of DP emulsions.

(50) The effect of wt % of the DP polymer solution on the fiber morphology of pure and emulsion electrospun DP scaffolds was compared using SEM imaging (FIGS. 3D and E). Proper fiber formation with pure DP solution was possible in the range of 8-12 wt % DP (FIG. 3D). Higher wt % of DP increased the viscosity of the polymer solution, resulting in blob on the fibers, inhomogeneity and increased fiber fusion. For emulsion electrospinning with DP, higher than 14 wt % polymer solutions did not allow for proper emulsion formation and jet formation during the electrospinning process. Emulsion electrospun DP fibers showed the anticipated smooth fiber surfaces, without blobs visible on the fibers, with 14 wt % DP emulsion electrospun scaffolds having high inhomogeneity (FIG. 3E). More viscous polymer solutions require more time and voltage during the electrospinning process for jet stretching and beads and blobs on the fibers can be present due to difficulties in proper jet formation..sup.[15]

(51) The porosity of DP scaffolds produced with single or emulsion electrospinning was gravimetrically determined and compared between 8, 10 and 12 wt % DP. The scaffold porosity for different wt % of DP single electrospun scaffold ranged from 68.353.07%, 66.335.58% to 63.431.96% for 8, 10 and 12 wt % DP scaffolds respectively, without significant differences between them (Figure S1A). For emulsion electrospun DP scaffolds, the porosity ranged from 76.112.26%, 74. 423.45 to 71.333.22 for 8, 10 and 12 wt % respectively (Figure S1A). Emulsion electrospun DP scaffolds exhibited significant increase in scaffold porosity when compared to each of the respective wt % of only DP electrospun scaffolds.

(52) In addition the hydrophilicity of the scaffolds was determined by measuring the water contact angle for each sample. Both pure and emulsion electrospun scaffolds (with FITC-BSA incorporated) are hydrophobic after the electrospinning, with static water contact angles from 104.611.49 to 124.862.87 (Figure S1B).

(53) Effect of Electrospinning Parameters on DP Fiber Diameter

(54) Different electrospinning parameters (voltage, flow rate, wt % of polymer solution, working distance) can affect the electrospinning process, the jet stability and morphology of polymer fibers obtained..sup.[16] Previous research on electrospinning parameters affecting DP electrospun scaffolds were performed using the original classic DP.sup.[17], which has significantly smaller strain at break (61.3311.37%) when compared to the new DP synthesized and explored here DP..sup.[4] Moreover, the solvent used in previous research was pure chloroform, while here, a mixture of chloroform and hexafluoro-2-propanol in a ratio of 80:20 was used in order to allow easy dissolution of DP without involving subsequent heating step. Taking this into consideration, the fiber diameter of pure and emulsion electrospun DP fibers was compared using different wt % of DP (8, 10, 12 and 14) and different electrospinning parameters (voltage (10-17.5 kV) and flow rate (0.5-3 mL h.sup.1)). Similar trends were observed in single and emulsion electrospun DP fibers (FIG. 4A). For single electrospun DP fibers, decrease in average fiber diameter [m] was observed with the increase in wt % of DP used, but without significant differences. A clear effect on fiber diameter upon differences in applied voltage [kV] was not visible. However, significant differences were obtained with differences in flow rates (mL h.sup.1) applied. Increasing the flow rate from 0.5 mL h.sup.1 to 3 mL h.sup.1 resulted in an average increase of fiber diameter from 3.832.12 m to 6.312.35 m. With increase in flow rate applied, a greater volume of solution is pumped out and while stretched with the same voltage, it leads to formation of thinner or thicker jets. However, too high flow rate may not allow for proper solvent evaporation before fiber deposition and increase in fiber defects or fiber fusion can be present.

(55) For emulsion electrospun DP scaffolds, a significant increase in fiber diameter was observed with the increase of wt % of DP, from 8 to 14 wt % DP and increase in flow rate, from 0.5 mL h.sup.1 to 3 mL h.sup.1. Higher fiber diameters obtained with an increased wt % of DP can be correlated with increased in viscosity of the polymer solutions. As a result, low voltages (10 kV) with higher wt % (14 wt %) did not allow for proper fiber formation. The applied voltage only showed an effect in the range of 12.5, 15 and 17.5 kV (FIG. 4A). These findings are in agreement with earlier studies..sup.[18]

(56) Emulsion versus single electrospinning of DP had a major effect on the range of fiber diameters obtained. Emulsion electrospun DP scaffolds resulted in an average diameters from 1.790.17 m up to 4.540.19 m, while pure DP scaffolds ranged from 4.50.79 m to 3.980.12 m. Emulsion DP fibers produced with 8-12 wt % DP polymer solutions had significantly smaller fiber diameters when compared to pure electrospun DP scaffolds in the same range (FIG. 4B). This effect can be accounted for by the change in properties of the polymer solution (viscosity and surface tension) once the aqueous phase is introduced..sup.[19] During emulsion electrospinning, the polymer jet undergoes increased and earlier splitting as well as increased instability, thus leading to decreased fiber diameters.

(57) Mechanical properties of emulsion electrospun DP scaffolds Variations of scaffold morphology and composition induce changes in the mechanical properties of the scaffolds and are highly dependent on the polymer..sup.[20] The new DP explored here, with very high elasticity, allows for easy and surgeon-friendly handling, when it is pulled over the wound site. With the goal of adjusting for this desired property, mechanical testing of emulsion electrospun DP tubes (with incorporated FITC-BSA) was performed. To determine differences arising due to changes in electrospinning method, DP tubes were produced from different wt % of DP (8, 10 and 12) with single and emulsion electrospinning. Formation of tertiary structures by fiber fusion was observed in thick electrospun DP scaffolds (500-800 m), resulting in fused fiber in the axial direction of the electrospun tubes (FIGS. 5A and B). This led to formation of parallel ridges along the rough surface of DP tubes in both methods. This effect was present only when DP was dissolved in chloroform and hexafluoro-2-propanol mixture and not in pure chloroform. Since this formation of ridges can influence the mechanical properties of the DP tubes, both directions of the tubes (axial and transverse) were tested. The major stretching of the DP tube during surgery handling happens in the transverse direction of the tube.

(58) When compared to pure DP scaffolds, the ultimate tensile stress (UTS) [MPa] of emulsion electrospun DP scaffolds decreased in both tube directions (FIGS. 5C and E) and the UTS was lower in the transverse direction, compared to the axial direction. Emulsion electrospun scaffolds experienced significantly lower UTS in the transverse direction (0.7590.055, 0.2280.034, 0.3380.013) when compared to pure DP scaffolds (0.9620.055, 0.670.123, 1.2950.132) for all wt % of DP. This difference can be due to scaffold's morphological differences, but also due to protein presence (in aqueous droplets) within the fibers, thus resulting in internal structural defects. Failure stresses of healthy rabbit Achilles tendons which are similar in strength as human flexor tendons have UTS in the range of 30 MPa..sup.[6] The UTS of the electrospun DP scaffolds (1-2 MPa) are too low to act as reinforcement of the structural tendons which are not intended when applied in tendon repair in a clinical setting. Significant decrease in strain at break [%] from pure DP to emulsion electrospun DP was observed for both directions of the tube (FIGS. 5D and F). Moreover, significant decrease in strain at break [%] in the transverse direction when compared to the axial direction for all samples was observed. In the transverse direction, emulsion electrospun scaffolds had strain at break of 311.3616.8, 224.1618.9 and 175.486.4%, for 8, 10 and 12 wt % DP, respectively. Pure DP scaffolds had higher values, i.e. 843.922.75, 827.565.96 and 79591.61, for 8, 10 and 12 wt % DP respectively (FIG. 5F). However, the more fragile emulsion electrospun DP scaffolds could still be stretched in the transverse direction on average 2 times from their initial size before breaking. This decrease in the strain at break is present due to the tertiary structure obtained by fiber fusion (FIG. 5B) when scaffolds started becoming thicker. Thicker scaffolds had ridges formed by fibers fused that could act as scaffold defects and decrease the scaffold's mechanical properties. These fused fiber bundles created along the axial direction of the tube, run along in the same direction as tendon collagen fibers, but perpendicular to the direction of tube stretching, thereby resulting in a decreased strain at break in the transverse direction. However, the properties of the emulsion electrospun scaffolds still allow successful handling and application during surgery.

(59) Young's modulus [MPa] of emulsion electrospun scaffolds (1.80.24, 1.500.15 and 1.750.21) when compared to single spun scaffolds (2.260.15, 1.600.17 and 2.060.13) for each wt % of DP decreased slightly, but not significantly (FIG. 5G). As Young's moduli of healthy rabbit tendons are in the range of 100 MPa.sup.[21], the comparatively low moduli of both emulsion and pure DP scaffolds are far from being similar to the tendons, but this is not relevant when scaffolds are used as drug delivery devices.

(60) Release Kinetics of Fluorescein and FITC-BSA from DP Scaffolds

(61) To study the release kinetics of entrapped molecules from emulsion electrospun DP scaffolds, fluorescein and FITC-BSA were incorporated as model biomolecules with low (376.27 g mol.sup.1) and high (66 kDa) molecular weights, respectively. The scaffolds were thereby spun with different wt % of DP, as well as, two different flow rates (1 and 3 mL h.sup.1) to tune the fiber diameters while maintaining the same applied voltage (12.5-13 kV). As revealed by confocal laser scanning microscopy, the molecules were randomly dispersed within the fibers (FIG. 6A) similar to other polymers.sup.[21] and no core-shell fibers.sup.[22] were formed.

(62) The cumulative release of incorporated fluorescein reached already 95% after 48 hours without further changes up to 5 days (FIGS. 6B and C), suggesting a burst-mediated release independent of the mentioned adjustable parameters. On the other hand, the release profile of FITC-BSA from DP scaffolds was a more sustained one. In 24 hours, less than 20% of FITC-BSA were released from the fibers, with a gradual increase every next time point. The maximum release within 14 days ranged from 58.426.18% for 12 wt % DP up to 64.904.26% for 8 wt % DP (FIG. 6E). No significant differences are visible in the FITC-BSA release profile from fibers with different wt % of DP (FIG. 6E) or if produced with different flow rates (ml h.sup.1) (FIG. 6D).

(63) As a polymer powder, DP is biodegradable within 3 months, mainly by hydrolytic degradation. Degradation of the scaffolds in aqueous conditions over the time period of 14-30 days was not visible (data not shown). Taking this into account, the release of molecules from the electrospun fibers in the studied aqueous environments (PBS buffer) is expected to happen primarily due to diffusion mechanism and not due to the degradation of the fibers within the time frame studied here (up to 14 days). This significantly can differ in in vivo setting, where at a wound site, the presence of different macrophages or MMPs can accelerate the polymer degradation..sup.[24]

(64) The differences in the release of fluorescein and FITC-BSA are due to their differences in molecular weight and size too..sup.[25] Fluorescein is readily dissolvable in water and upon placing the scaffolds in aqueous environment it can readily dissolve in it, without being adsorbed on the surface of the scaffolds. In contrast, BSA as a protein easily adsorbs on surfaces. It has stronger interactions with hydrophilic surfaces and can cover up to 95% of the surface. On hydrophobic surfaces it usually covers around 50% of the surface..sup.[26] DP being hydrophobic, released BSA can adsorb up to some extent on the surface and be in an equilibrium of protein release and protein re-adsorption on the scaffold surface.

(65) In addition, the DP fibers are experiencing quite smooth surfaces without visible nanopores on their fiber surface. Fluorescein having a Stokes radius of 0.55 nm can easily diffuse out of the DP fibers. On the other hand, larger proteins like BSA (Stokes radius=3.48 nm) might experience limited diffusion through the fibers, with major part of released BSA being close to or on the fiber surface, and the rest that is inside the fibers to be released only after scaffold degradation.

(66) Release Kinetics of PDGF-BB from DP Scaffolds

(67) To assess the possible interaction of PDGF-BB (24.3 kDa disulfide-linked homodimer of two P chains) with DP surfaces that might play a role in its release from the scaffolds, PDGF-BB was physically adsorbed on pure DP scaffolds and immunostaining was performed in order to detect it on the surface of the fibers. From the image (FIG. 7A) it is visible that its distribution on the surface of DP fibers was uniform. On the other hand, emulsion electrospun DP scaffolds were also immunostained for PDGF-BB and sparse, randomly distributed PDGF-BB was visualized, that got on the surface of the scaffolds during the production procedure, with the rest incorporated within the polymer fibers and not accessible for detection (FIG. 7B).

(68) From all electrospinning parameters screened and release kinetics of model molecules studied, 12 wt % DP, 11-12 kV applied voltage, 1 ml mL.sup.1 flow rate and 20 cm working distance were chosen as best parameters for the production of the bioactive emulsion electrospun DP scaffolds. The in vitro PDGF-BB release from emulsion electrospun DP fibers showed similar trend to FITC-BSA, i.e. a more sustained release, rather than burst release as fluorescein exhibited (FIG. 7). Within 30 days, the cumulative release of PDGF-BB in pg ml.sup.1 reached around 3500 g ml.sup.1, i.e. 3.5 ng ml.sup.1 (FIG. 7C). Due to the sensitivity of growth factors, the DP scaffolds were not dissolved in chloroform and PDGF-BB was not extracted at the end of the release period. To obtain quantitative data on the release, the cumulative percentage of PDGF-BB released was calculated using a theoretical loading (13.33 ng/mg of DP scaffold). This resulted in a very low cumulative release of PDGF [%] of average 1.3% within 30 days (FIG. 7D). Emulsion electrospinning does not offer 100% loading efficiency, thus we hypothesize that this cumulative release % is larger. In addition, the amount of PDGF-BB released was normalized to the weight of DP scaffold, resulting with an initial release of 31.811.71 g mg.sup.1 of DP and within 30 days having a cumulative release of 170.635.71 g mg.sup.1 of DP (FIG. 7E). PDGF-BB physiological concentrations in whole blood have been determined to be in the range of 3.30.9 ng mL.sup.1 and increase to 178 ng mL.sup.1 in platelet-rich plasma (PRP).sup.[27]. PRP has been suggested as a strategy for promoting wound-healing cascade..sup.[28] In in vitro conditions, within 30 days period the released PDGF-BB amounts from the scaffolds were within the physiological range of whole blood. Due to different conditions and enzymatic DP degradation possible in in vivo settings, one should expect different release of PDGF-BB, with possibly larger amount of PDGF-BB released, within 2-4 weeks.

(69) PDGF-BB Effect on Rabbit Tenocyte Proliferation and Bioactivity Assays

(70) To determine the concentration at which PDGF-BB has a significant effect on rabbit tenocyte proliferation, different concentrations of PDGF-BB were screened in serum free and serum.sup.+ medium. Serum starved synchronized cells were exposed to 1-50 ng/ml PDGF-BB for 24 hours. As quantitative measure of proliferating cells, the percentage of EdU-positive cells (marking DNA synthesis in proliferating cells) was determined for each condition. In serum free medium, the lowest PDGF-BB concentrations (1 and 3 ng mL.sup.1) already induced an increase in cell proliferation when compared to a control, but the increase in cell proliferation being significant only at 5 ng mL.sup.1 and higher PDGF-BB concentrations (FIGS. 8A and C). These results are in agreement with the effect of PDGF-BB on proliferation of avian tendon fibroblasts..sup.[29] In the presence of serum, the tenocytes showed high proliferation rate already at day 1 and a slight, but not significant increase in proliferation with increase in PDGF-BB concentrations (10-50 ng mL.sup.1) (FIGS. 8B and D). Taking this into account, bioactivity of concentrated samples of released PDGF-BB was assessed in serum free conditions, where its amount was in the lower concentration ranges tested (FIG. 8E). Released PDGF-BB overtime from the electrospun scaffolds lead to increase in tenocyte proliferation when compared to no PDGF-BB presence, thus confirming that PDGF-BB retains bioactivity after electrospinning and release (FIG. 8E). In addition, WST-1 proliferation assay was performed with rabbit tenocytes cultured in serum free and serum.sup.+ conditions to observe their metabolic activity (proliferation) and viability within 14 days, without or with PDGF-BB (5, 10, 25 and 50 ng mL.sup.1). Tenocytes cultured in serum free conditions did not show increased mitochondrial metabolic activity and were viable, but their metabolic activity and proliferation increased significantly with the addition of PDGF-BB (Figure S2A). Tenocytes cultured in serum.sup.+ medium, showed higher proliferation rates than tenocytes in serum free conditions, however the proliferative effect of PDGF-BB after 14 days, was mostly visible and significant at 25 and 50 ng mL.sup.1 PDGF-BB concentrations (Figure S2B).

(71) To obtain a more physiologically relevant bioactivity assay in serum.sup.+ conditions, tenocyte proliferation was tested directly on pure and bioactive DP scaffolds, expecting that local PDGF-BB concentration is higher due to the equilibrium of growth factor bound to the scaffold surface and released into the medium. Tenocytes seeded on electrospun scaffolds (in serum.sup.+ conditions) generally had decreased proliferation, 18% (FIG. 8F) when compared to seeded tenocytes on glass coverslips (Figure S3B). These differences can be correlated with recently published study where has been shown that cell mediated fiber recruitment drives extracellular matrix mechanosensing and regulates cellular functions in fibrillary microenvironments. Lower fiber/network stiffness enabled cells to recruit nearby fibers leading to increased local adhesive ligand density, cell spreading and proliferating signaling..sup.[30] We hypothesize that through this mechanism, tenocytes seeded on DP electrospun scaffolds, sensed quite dense fiber networks, resulting in a decrease in proliferation when compared to cell seeded on coverslips.

(72) Bioactive emulsion electrospun DP scaffolds showed a significant increase in proliferation (31.575.09%), when compared to pure DP scaffolds without PDGF-BB (18.966.82%) or empty emulsion electrospun DP with only BSA incorporated as a control (18.703.34%). Pure DP scaffolds with physically adsorbed PDGF-BB on the surface were used as a positive control. Tenocytes seeded onto these scaffolds experienced significant increase in proliferation (28.375.21) when compared to the pure DP scaffolds or empty emulsion electrospun DP scaffolds (FIG. 8F). Tenocytes seeded on scaffolds incorporating PDGF-BB in any of the two ways, had increased proliferation without significant differences between physically adsorbed or emulsion incorporated PDGF-BB. This data suggests indeed higher local concentration of PDGF-BB on the scaffold surface sensed by the cells seeded onto the bioactive emulsion electrospun scaffolds, with some amount of it eventually being released into the medium. It also suggests that the incorporated PDGF-BB into the scaffolds is still biologically active and leads to increase in cellular proliferation.

(73) Exposure of NIH3T3 cells to low concentrations of PDGF-BB (<2 ng mL.sup.1) was shown to have greater effect on cell migration rather than cell proliferation, due to differences in endocytotic routes of the PDGF receptor after PDGF binding. Higher PDGF-BB concentrations (30 ng mL.sup.1) resulted in increased cell proliferation..sup.[13] Similar behavior can be seen in the rabbit tenocyte proliferation, where 1 and 3 ng mL.sup.1 did not lead to significant increase in cell proliferation, but >5 ng mL.sup.1 PDGF-BB did. Since bioactive DP tubes are to be directly and locally applied at the wound site, with cells binding to the scaffold directly, we expect to have a condition closely resembling the serum.sup.+ condition of bioactivity assay, where some tenocytes would be exposed to higher local PDGF-BB concentration and some experience the migratory effect of PDGF-BB administration.

CONCLUSION

(74) Out data suggest that emulsion electrospun DP tubes are promising for Achilles tendon rupture repair and we identified conditions under which they can be produced to be sufficiently elastic and bioactive. We have screened and defined a range of electrospinning parameters best suited for emulsion electrospinning of DP, incorporating different biomolecules within the polymer fibers. Emulsion electrospinning of DP lead to fiber diameter decrease and decrease in the strain of break [%] in both tube's directions, compared to pure DP, but still offering enough scaffold elasticity for successful clinical use. High molecular weight FITC-BSA and PDGF-BB experienced a sustained release kinetics from the emulsion electrospun DP scaffolds, compared to a burst release of low molecular weight fluorescein form the same. However, large fraction of bioactive molecules was left inside the DP fibers and was not released under the tested in vitro conditions. The released PDGF-BB was shown to be bioactive, leading to increased proliferation of rabbit tenocytes in in vitro under serum free conditions. In addition, tenocytes seeded directly onto bioactive DP scaffolds showed increased proliferation in serum.sup.+ conditions, suggesting higher local PDGF-BB concentration on the scaffolds surface. As a next step, the bioactive DP tube, delivering PDGF-BB is aimed to be implanted and tested in in vivo rabbit model.

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Example 2

(76) Application of PDGF-BB Loaded DegraPol Tubes in Rabbit Achilles Tendons

(77) Three weeks post-surgery, double-layered DegraPol tubes having PDGF-BB incorporated by emulsion electrospinning and implanted around a fully transsected rabbit Achilles tendon being repaired with a conventional 4-strand Becker suture led to a higher ultimate failure load of extracted tendons compared to DegraPol tubes fabricated in the same way, however, without incorporation of PDGF-BB (FIG. 9a) (t test to compare failure load of the two groups (PDGF-BB): p=0.083). The same was observed for ultimate stress (FIG. 9b) (t test to compare the ultimate stress of the two groups (PDGF-BB): p=0.026). The addition of this growth factor showed therefore to beneficially enhance tendon healing. As such, most occurring problems after tendon rupture and repair can be addressed; adhesion formation is reduced by the tube acting as a physical barrier and tendon strength is enhanced by incorporation of this growth factor-lowering the risk of re-rupture[1]. [1] Elliot, D. & Giesen, T. Primary Flexor Tendon Surgery: The Search for a Perfect Result. Hand Clin. 29, 191-206, doi:10.1016/j.hcl.2013.03.001 (2013).

Example 3

(78) Application of Different DegraPol Tubes in Rabbit Achilles Tendons

(79) FIG. 10 shows a comparison of various biomechanical parameters determined three weeks post-surgery of fully transsected rabbit Achilles tendons being repaired with a conventional 4-strand Becker suture. One series of experiments represented with black bars in the figure was made using emulsion electrospun DegraPol tubes, whereas another series of experiments represented with grey bars was made with coaxially electrospun tubes. In each of these two series, a comparison was made between three groups, namely, a group receiving the application of a tube made with DegraPol only (denoted by DegraPol Tube), a group receiving the application of a tube made with DegraPol incorporating the growth factor PDGF-BB (denoted by DegraPol Tube+PDGF-BB) and a group receiving no tube (denoted by Control). The measured properties were: (a) tendon length in mm, (b) cross sectional area (CSA) in cm.sup.2, (c) load until failure in N, (d) failure stress in MPa, (e) stiffness in N/mm, and (f) elastic modulus in MPa.

(80) According to a one-way analysis of variance (one-way ANOVA) performed to test the differences between groups in all the experiments, the load until failure achieved after tendon repair using coaxially electrospun tubes containing PDGF-BB was significantly larger than in the case of unfilled coaxially electrospun tubes (see FIG. 10c). If using a simple t-test, a significant difference also holds for a corresponding comparison regarding stiffness (see FIG. 10e).