Low-Temperature Additive Manufacturing of Biomimic Three

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Low temperature additive manufacturing biomimic threedimensional hydroxyapatite/collagen scaffolds for bone regeneration Kai-Feng Lin, Shu He, Yue Song, Chun-Mei Wang, Yi Gao, JunQin Li, Peng Tang, Zheng Wang, Long Bi, and Guo-Xian Pei ACS Appl. Mater. Interfaces, Just Accepted Manuscript • DOI: 10.1021/acsami.6b00815 • Publication Date (Web): 01 Mar 2016 Downloaded from http://pubs.acs.org on March 4, 2016

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ACS Applied Materials & Interfaces

Low temperature additive manufacturing biomimic three-dimensional hydroxyapatite/collagen scaffolds for bone regeneration Kai-Feng Lin

†,§

, Shu He

†,§

, Yue Song

†,§

, Chun-Mei Wang †, Yi Gao †, Jun-Qin Li †,

Peng Tang †, Zheng Wang †, Long Bi †,*, Guo-Xian Pei †,* †

Department of Orthopaedics, Xijing Hospital, The Fourth Military Medical

University, Xi’an, 710032, PR China.

1

ACS Paragon Plus Environment


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KEYWORDS: Three dimensional printing, collagen, hydroxyapatite, Bone substitutes, Tissue engineering

ABSTRACT Low temperature additive manufacturing (AM) holds promise for fabrication of three dimensional (3D) scaffolds containing bioactive molecules and/or drugs. Due to the strict technical limitations of current approaches, few materials are suitable for printing at low temperature. Here, a low temperature robocasting method was employed to print biomimic 3D scaffolds for bone regeneration using a routine collagen-hydroxyapatite (CHA) composite material, which is too viscous to be printed via normal 3D printing methods at low temperature. The CHA scaffolds had excellent 3D structure and maintained most raw material properties after printing. Compared to non-printed scaffolds, printed scaffolds promoted bone marrow stromal cell proliferation and improved osteogenic outcome in vitro. In a rabbit femoral condyle defect model, the interconnecting pores within the printed scaffolds facilitated cell penetration and mineralization before the scaffolds degraded and enhanced repair, compared to non-printed CHA scaffolds. Additionally, the optimal printing parameters for 3D CHA scaffolds were investigated; 600 µm-diameter rods were optimal in terms of moderate mechanical strength and better repair outcome in vivo. This low temperature robocasting method could enable a variety of bioactive molecules to be incorporated into printed CHA materials and provides a method of bioprinting biomaterials without compromising their natural properties.

2


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1. INTRODUCTION Autograft and allograft tissues have been used to repair bone defects for several decades. However, traditional bone graft supply has many practical and surgical complications and cannot meet the escalating demand for bone defect repair 1. Synthetic biomaterials with specific structures represent promising alternatives

1, 2

,

and numerous bone grafting materials have been fabricated via conventional approaches such as solvent casting, particle/salt leaching

3, 4

,chemical/gas foaming 5,

freeze drying 6, phase separation 7, and foam-gel techniques 8. Although mean pore size of scaffolds can be controlled using such approaches, it is challenged to prepare materials with precise shape, pore size and pore interconnectivity of materials via these approaches

9, 10

. Additive manufacturing (AM) techniques have been widely

developed and possess the advantages of enabling a tailored shape, precise pore size control and a high pore interconnection ratio, in combination with rapid production 11. Commercially-available AM techniques include powder-based printing (3DP), stereolithography (SLA), fused deposition modeling (FDM), selective laser sintering (SLS) and robocasting

12

. Powder-based printing possesses significant potential for

the manufacture of bone grafting scaffolds as almost any powder can be used, provided the powder is combined with an adequate binder. For example, ceramic/glass (hydroxyapatite, β-TCP, glass etc.) and polymeric (poly-L-lactide [PLLA], poly(lacticco-glycolic) acid [PLGA] etc.) powders have been used to fabricate various bioabsorbable scaffolds

13

. Unfortunately, pure ceramic/glass and

synthetic polymeric-based scaffolds have the disadvantages of excessive mechanical 3



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strength, poor osteoinductivity and biodegradability, and toxic degradation by-products. Therefore, the scaffolds using these conventional process methods have demonstrated limited clinical success 14. One common strategy to improve the biocompatible properties of such scaffolds is to incorporate bioactive molecules

15, 16

. However, the due to limitations of AM

techniques, such as the limited percentage of high viscosity polymer that can be incorporated

17

, the use of extremely harsh chemical solvents

temperatures during processing or post-processing

18

and the high

12

, it is difficult to incorporate

bioactive molecules using AM techniques. Coating the scaffolds with bioactive molecules after printing is regarded as a second-choice, and the effects are limited as the coating process is only a surface modification 19, 20. In order to disperse bioactive molecules or drugs evenly within scaffolds by including these components within the printing inks, a medium that can be printed at a relatively low temperature is required. The inks used in powder-based printing should be low viscosity. However, collagen is a well-characterized high viscosity material. In order to print collagen via powder-based printing at low temperature, additional chemical solvents are necessary to reduce the surface tension of the binder. Since the ability of a chemical solvent to reduce viscosity is limited, the potential quantity of the collagen in inks is restricted 21. Moreover, any change in the raw materials has to also meet the requirements of the printing process, which makes it difficult to freely modify the raw materials of the inks 17. Furthermore, if a large, complicated scaffold is printed, residual power present inside the pores after printing may cause series of adverse consequences. Additionally, 4


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the high cost of the powder not used during printing would be an economic waste 22. To achieve the ultimate goal of clinical application, an ideal approach should possess the advantages of few technical limitations on the ink, be free of unnecessary chemical solvents, economical and be able to produce a stable structure via 3D printing at a low temperature. Hydroxyapatite (HA) and collagen are two major constituents of bone and are 23-26

frequently used for bone reconstruction osteoconductivity and bioactivity osteoinductive

. HA has excellent biocompatibility,

27

. Collagen is biocompatible, biodegradable and

23, 24

, and provides natural binding sequences for cell attachment that

are absent from bioceramics and synthetic polymers

28

an

molecules

excellent

delivery

system

for

bioactive

. Moreover, collagen provides or

drugs

29,

30

.

Collagen-hydroxyapatite (CHA) scaffolds have been demonstrated to combine the advantages of the mechanical strength of ceramics with the biological advantages of collagen

31

. Therefore, it seems a logical step to fabricate CHA scaffolds via 3D

printing to support and promote bone regeneration. However, to our knowledge, the feasibility of 3D printed CHA composites containing 33% percent collagen has not yet been demonstrated. Robocasting, also termed direct-write assembly, is a unique AM technique as it enables the creation of ceramic scaffolds using water-based inks with minimal organic content and does not require any sacrificial support material or mold recently-developed filament-free printing (FFP) technique

33, 34

32

. The

further improved the

robocasting approach, employing a progressive cavity pump to freely mix materials to 5



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enable the printing of a variety of composite materials. Consequently, it is possible to print high viscosity inks at low temperature, while the structure of the scaffold is freely controllable and can have a high interconnected pore ratio. In the present study, a CHA scaffold with a collagen/HA ratio of 1/2 (w/w) was printed using this modified robocasting approach at 4°C. The potential of these scaffolds for bone defect repair were systemically evaluated in vitro using rabbit bone marrow stromal cells (BMSCs) and in vivo in a rabbit femoral condyle defect model. We hypothesize that the excellent biomaterial properties of CHA are maintained after printing and the resulting 3D structure will further improve the adhesion, proliferation and osteogenic differentiation of bone marrow stromal cells, and thus promote graft-to-bone healing.

2. MATERIALS AND METHODS

2.1 Scaffold fabrication Lyophilized type I collagen isolated from bovine achilles tendon was purchased from Sannie Bioenginnering Technology Company (Tianjing, China). Hydroxyapatite powder (particles smaller than 100 nm in diameter) was purchased from Emperor Nano-material Company (Nanjing, China). Scaffolds of CHA with a grid-like microstructure were prepared using a define FFP printer (PC Printer, Particle Cloud Biotechnology, Xi’an, China) as illustrated in Figure 1A. Briefly, collagen and hydroxyapatite were individually homogenized in 0.5 M acetic acid solution (AAS). 6


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Then the two slurries were homogenized to create inks (2 g, collagen, 4 g HA, 10 ml AAS) using an overhead blender (Ultra Turrax T18; IKA Works Inc., Wilmington, USA) at 4°C. The cartridge was filled with the ink, tapped vigorously under slight vacuum to remove bubbles and placed on the robotic deposition device, which is controlled by a FFP software program integrated within the printer. The inks were quantitatively delivered through a conical nozzle (diameter = 300 µm) and deposited on a stainless steel panel at a constant linear printing speed of 10 mm/s. The thickness of each layer was set to 200 µm and each layer was repeatedly printed twice in the z-axis before changing printing direction to form side holes, resulting in an actual layer thickness of 400 µm. An x-y axis repeated writing method was used to produce microstructures with same pore widths of 400 µm and one of three rod widths: Group I, 300 µm; Group II, 600 µm and Group III, 900 µm. To prevent non-uniform drying, the environment was maintained at 90% relative humidity (RH) and 4°C during assembly. Cubic (10 x 10 x 3 mm) and cylindrical (10 x Φ 5 mm) scaffolds were printed for the in vitro and in vivo tests, respectively. To produce non-printed scaffolds (Group IV), the ink was placed directly into a mold with a size and shape corresponding to the printed scaffolds. The scaffolds were cooled to -80°C at a constant cooling rate of 0.9°C/min and lyophilized using a freeze-dryer (Alpha 1-2 LD plus, Christ, Osterode, Germany). All scaffolds were further crosslinked by immersion in 1% (w/v) genipin solution (Wako Pure Chemical Industries, Ltd., Kanagawa, Japan) to improve their mechanical character and were sterilized using ethylene oxide

35, 36

. Finally, the scaffolds were aseptically packaged until used in 7



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experiments.

2.2 Scaffolds characterization 2.2.1 Scanning electron microscopy The microstructure and morphology of the scaffolds were characterized by scanning electron microscopy ([SEM] S-4800; Hitachi, Tokyo, Japan). The surface element content of the scaffolds was examined using an energy dispersive spectrometer (EDS) integrated within the SEM. Integrated mapping software was used to analyze the element distribution.

2.2.2 Wettability properties The wettability of the scaffolds was evaluated via contact-angle measurements using an EasyDrop Standard (KRUSS, Hamburg ,Germany). Briefly, a drop of water was dropped from 2 cm above the scaffold. When the drop contracted with surface of the scaffold, photographs were captured and the contact angle was calculated (n = 5).

2.2.3 Porosity The open porosity of the scaffolds (n = 5) was measured using the liquid displacement method in ethanol

37

. Briefly, a dry porous scaffold was placed into a

pycnometer and weighed (Wps), then the pycnometer was completely ﬁlled with ethanol and the net weight was registered (Wpse). Subsequently, the scaffold was removed and the net weight of the residual ethanol and pycnometer was measured 8


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(Wpr). Wp is the weight of the dried pycnometer; Vp is the volume of the pycnometer (Vp = 25 mL); ρe is the density of ethanol at the measurement temperature. The porosity of the scaffolds was calculated as follows: (%) =

( − ) − (Wpr − Wp) /ρe ∗ 100% Vp − (Wpr − Wp) /ρe

2.2.4 Mechanical properties Unconfined compression testing was carried out using a Model 858 Material Testing System (MTS Systems, Minneapolis, USA). The scaffolds (n = 5) were submerged in phosphate-buffered saline (PBS) for 4 h prior to testing then placed between two platens. Compressive tests were conducted up to a maximum compressive strain of 10%, at a strain rate of 10% per minute. The compressive modulus was defined as the slope of a linear fit to the stress-strain curve over 2-5% strain.14

2.3 In vitro tests 2.3.1 Isolation and culture of rabbit bone marrow stromal cells (BMSCs) All animal procedures were approved by the institutional animal care committee of The Fourth Military Medical University and carried out in strict accordance with the Guidelines on the Care and Use of Laboratory Animals issued by the Chinese Council on Animal Research and Animal Care. BMSCs were prepared as previously described

38

. Briefly, bone marrow was harvested under aseptic conditions from the 9



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tibia and femur condyle of anaesthetized male New Zealand rabbits (1 month-old; 0.8 kg) and resuspended in α-minimum essential medium (α-MEM; Hyclone, Logan, USA) containing 10% fetal bovine serum (FBS, Gibco, Grand Island, USA), 100 U/ml penicillin and 100 mg/L streptomycin. Cells were cultured at 37°C in 5% CO2/95% humidity. After 48 h, non-adherent cells were carefully removed by washing and the culture medium was changed. Thereafter, the culture medium was replaced every 2-3 days. The third-passage cell subcultures were used in the experiments. Flow cytometry was used to confirm surface antigen marker CD34 and CD44 expression. BMSCs were incubated with anti-CD34-APC (Bioss, Beijing, China), anti-CD44-PE (Bioss, Beijing, China) for 30 min at 4°C, washed, collected and analyzed using the Facscan Flow Cytometry system (BD, Franklin Lakes, USA). Antibody was replaced with PBS for negative control staining. 2.3.2 Seeding of BMSCs onto scaffolds Scaffolds were washed in α-MEM three times before seeding. BMSCs were seeded onto the scaffolds at 3 × 105 cells/cm³, then cultured at 37°C in 5% CO2/95% humidity; the medium was changed every 2 days.

2.3.3 Evaluation of scaffold cytotoxicity towards BMSCs The cytotoxicity of the scaffolds towards BMSCs was evaluated using the CCK8 assay (Dojindo, Kumamoto, Japan) and Annexin V-FITC/propidium iodide (PI) staining (Liankebio Company, Hangzhou, China)

39

. For the CKK8 assay, BMSCs

10


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were seeded into 96-well plates (1 × 104 cells/well) and 8 h later, the medium was replaced with control media or conditioned media in which the CHA scaffolds had been soaked in for 48 h at 37°C. After 24 h, CCK8 solution was added and the samples were incubated at 37°C for 1 h. The optical density (OD) values were measured using a microplate reader (Dragon Wellscan MK3, Labsystems , Bucharest, Romania) at 450 nm (n = 5). For the Annexin V-FITC/PI assays, the cells on the scaffolds or culture dish were digested and collected 24 h after seeding, stained with Annexin V-FITC for 30 min at 4°C while protected from light, followed by PI for 10 min at RT while protected from light, then analyzed using the Facscan Flow Cytometry system (n = 3).

2.3.4 Analysis of BMSC adhesion and proliferation on scaffolds The seeding efficiency of the scaffolds was measured via counting the number of cells released from the scaffold by trypsin digestion at 8 h after seeding using an Automated Cell Counter (TC-10; Bio-Rad, Hercules, USA). The seeding efficiency was calculated by dividing the numbers of cells on the scaffolds by the numbers of cells seeded (n = 5). Additionally, the CCK8 assay was used to measure the proliferation of cells on the scaffolds at 1, 4, 7 and 11 days after seeding (n = 3). Vinculin, F-actin and nuclei were identified using anti-vinculin antibody (AbD Serotec, kidlington, England), fluorescent Rhodamine-phalloidin (Cytoskeleton, Denver, USA) and DAPI (Sigma Aldrich, St. Louis, USA), respectively, At 24 h after seeding, the cells adhered to the scaffolds or culture dish were fixed with 4.0% w/v 11



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paraformaldehyde for 20 min, permeabilized with 0.1% Triton X-100 for 5 min and blocked using 1% w/v bovine serum albumin (BSA) for 30 min at room temperature. Then, the samples were incubated with anti-vinculin antibody overnight at 4°C, fluorescein isothiocyanate (FITC)-labeled goat anti-mouse IgG antibody for 4 h at 4°C (Cwbiotech, Bejing, China), fluorescent Rhodamine-phalloidin for 30 min at RT, followed by DAPI for 5 min at RT, with three PBS washes after every incubation. The scaffolds were mounted on coverslips and examined using a confocal laser scanning microscope (TCS SP5; Leica, Wetzlar, Germany). The integral optical density (IOD) values of the vinculin immunostaining images (green) were calculated using Image pro-Plus 6.0 software. Cell adhesion morphologies were assessed via SEM at 1, 4, 7 and 11 days after seeding. Briefly, the adhered cells and scaffolds were fixed in 2% v/v glutaraldehyde at 4°C overnight, dehydrated through an ethanol series, critical-point dried, sputtered with gold and subjected to SEM.

2.3.5 Analysis of BMSC differentiation on scaffolds Alkaline phosphatase (ALP) activity (n = 3) was assayed using an ALP assay kit (Beyotime, Jiangshu, China). The media was replaced with osteoinductive medium at 24 h after seeding. The cells were digested, collected and lysed after 7 and 14 days culture in osteoinductive media. Cells cultured with osteoinductive medium but not seeded on a scaffold were prepared as a control (Group N). Subsequently, the lysates were reacted with p-nitro-phenyl phosphate (p-NPP), and the absorbance of 12


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p-nitrophenol was determined at 405 nm using a microplate reader. The total intracellular protein content was determined using the BCA protein kit (Beyotime). ALP activity was normalized to total intracellular protein synthesis and expressed as nanomoles of p-nitrophenol produced per min per mg of protein (nmol/min/mg protein). Cells (n = 3) were also collected after incubation in osteoinductive media for 3, 7 and 14 days to evaluate the gene expression levels of collagen type I alpha 1 (COL1A1), ALP, runt-related transcription factor-2 (Runx2), osteopontin (OPN) and glyceraldehyde-3-phosphate dehydrogenase (GAPDH). Briefly, total RNA was extracted according to the instructions of the RNA extraction kit (Omega, Norcross, USA).The extracted RNA samples were reverse transcribed to generate cDNA, which was used to perform real-time PCR using a Bio-Rad CFX Manager system. The primers used are listed in Table 1. Table 1: Primer sequences Gene

Forward primer sequence (5`-3`)

Reverse primer sequence (5`-3`)

13



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COL-1A1

AAGAACGGAGATGACGGAGAAG

GCACCATCCAAACCACTGAA

ALP

GGCACAAGCACTCCCACTTT

TGGTCAATCCTGCCTCCTTC

RUNX-2

GCAGGAGGAAAATGAGCAAAG

GTGAGTGAGCAGAGCCGAGA

OPN

CACCGCAGAATGCTATGTCC

GTGGTCATCGTCCTCATCCTC

GAPDH

CCACTTTGTGAAGCTCATTTCCT

TCGTCCTCCTCTGGTGCTCT

2.4 In vivo tests 2.4.1 Implantation surgery procedure New Zealand rabbits (male, 12-weeks-old, 3.25 ± 0.25 kg) were fasted overnight with free access to water before the experiment. The rabbits were anaesthetized via intramuscular injection of xylazine hydrochloride (4 mg/kg) and 2% w/v pentobarbital (30 mg/kg). After disinfection and incision, a tunnel (Φ = 5 mm, L > 10 mm) was drilled through the condyle of the femur using a stainless steel drill. Printed scaffolds (Φ = 5 mm, L = 10 mm) with various rod widths (Group I, II, III; n = 12) or non-printed scaffolds (Group IV; n =12) were placed into the tunnel defects. Bone wax was used to seal the tunnels. Tunnels without any scaffolds were prepared as negative controls (Group N; n =12). The wounds were sutured in layers. During the three days post surgery, the rabbits were administered gentamicin (5 mg/kg) and penicillin (50 kU/kg) each day. Three rabbits of each group were humanely euthanized at 2, 4, 8 or 12 weeks after surgery for subsequent examinations. Rabbits

14


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(n = 3 per group) were intramuscularly injected with tetracycline (80 mg/kg; Sigma) at 14 days and calcein (8 mg/kg; Sigma) at 3 days before being euthanized at the 8-week time-point.

2.4.2 Micro-CT analysis The femur were carefully removed, fixed in 80% ethanol for 4 days, placed in the sample holder and scanned using a Micro-CT system (Y. Cheetah; Yxlon, Hamburg, Germany). The X-ray source voltage was set to 80 kV, beam current to 200 mA; the scanning angular rotation was 360° and the angular increment was 0.40°. The projections were reconstructed and a 5 × 10 mm2 cylindrical region was selected in the middle of the tunnel defects along the same major axis as the region of interest (ROI) using VGStudio MAX software (Volume Graphics, Heidelberg, Germany). The scaffold degradation rate was evaluated using RSV/SV, where RSV is the volume of the residual scaffold and SV is the total volume of the scaffold before implantation. BV/TV was calculated as the volume of the new bone formation divided by the total volume of the ROI.

2.4.3 Histology analysis After Micro-CT analysis, the specimens were carefully pruned to the femoral condyle and dehydrated in a graded ethanol series (70−100%), soaked in methyl methacrylate (MMA) solution for 3 weeks, solidified. Then, sections were prepared for pathological analysis using the modified interlocked diamond saw (Leica 15



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Microtome, Wetzlar, Germany). Sections prepared from animals euthanized at 8 weeks were observed using a fluorescence microscope to calculate the mineral apposition rate (MAR) of new bone formation, which is the length between the two different labels (µm/d) (n = 10). Then, all sections were stained with 1.2% trinitrophenol and 1% acid fuchsin (Van Gieson staining) and observed using a light microscope (DM6000B, Leica Microsystems).

2.5 Statistical analysis All quantitative data were analyzed using SPSS 14.0 or Graphpad Prism 5 statistical software. The results are presented as the mean ± SD values for each group. Statistical comparisons were carried out using the Student’s t-test or one-way ANOVA followed by Bonferroni's multiple comparison tests. Significance was defined as P < 0.05.

3. RESULTS 3.1 Scaffolds fabrication and characterization As shown in Figure 1B, four groups of CHA scaffolds were produced by robocasting printing (Group I, II, III) or mold (Group IV). High-resolution SEM images showed both printed (Figure 2A, B) and non-printed (Figure 2C, D) scaffolds had a relatively smooth surface and a plate-like crystal internal microstructure, similar to cancellous bone. The element distribution of calcium (Ca) and carbon (C) on the surface of scaffolds were mapped and merged to reflect the degree of dispersion of 16


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collagen and HA (Figure 2E, F, G). Both C and Ca were uniformly dispersed within the scaffolds except that some HA particles aggregating the spherical masses reflected by the Ca distribution. The water contact angle (wettability) of the scaffolds result indicated that there was no significant difference between the four groups of scaffolds and the average water contact angle of all scaffolds was approximately 46° (Figure 2H). The printed scaffolds (Group I, II, III) had a significantly higher porosity than the non-printed scaffolds (Group IV; P < 0.05). As the rod width of the printed scaffolds increased from 300 to 900 mm, porosity decreased from 82.9% to 71.8%; the porosity of Group I and Group III scaffolds was significantly different (P < 0.05; Figure 2I). Additionally, the compressive modulus of the Group I and II scaffolds were significantly lower than those of the Group III and IV scaffolds (P < 0.05). However, the compressive modulus of Group III was slightly higher than that of the Group IV scaffolds (P > 0.05; Figure 2J).

Figure 1: (A) Schematic of the robocasting fabrication process. (B) Gross view of the surface morphology of the experimental collagen/hydroxyapatite (CHA) 17



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composite scaffolds. The printed scaffolds had a grid-like microstructure with pore widths of 400 µm and rod widths of 300 µm (Group I), 600 µm (Group II) or 900 µm (Group III). Group IV was non-printed scaffolds. Gross views of the structure of example printed complicated scaffolds: bone with marrow cavity (C), femoral head (D) and tibial plateau (E). Scale bar: 5 mm

Figure 2. SEM images of the surface morphology (A) and internal microstructure (B) of the printed CHA scaffolds, and the surface morphology (C) and internal microstructure (D) of the non-printed CHA scaffolds. Element distribution of calcium (E) and carbon (F) within the scaffolds and the corresponding merged images (G). Water contact angle analysis (H) for the four groups of scaffolds. Porosity (I) and compressive modulus (J) for the four groups of scaffolds (αP < 0.05 vs. Group I; βP < 0.05 vs. Group II; γP < 0.05 vs. Group III; δP < 0.05 vs. Group IV). CCK8 assay (K) of the cytotoxicity of conditioned media prepared from CHA scaffolds (1 day incubation) or 18


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control media towards BMSCs. Assessment of BMSC survival and apoptosis at 1 day after seeding on the scaffolds or culture dishes (control) using the Annexin V-FITC/PI double-staining assay (L).

3.2 Isolation and culture of rabbit BMSCs After the third passage, the isolated cells had a uniform spindle morphology and formed swirl-shaped cell colonies (Supporting Information, Figure S1A). Flow cytometry revealed that 95.6% of the cells were CD34-CD44+, which confirmed that the isolated cells were BMSCs (Supporting Information, Figure S1B).

3.3 Evaluation of scaffold cytotoxicity towards BMSCs As shown in Figure 2K, the CCK8 assay indicated that conditioned media prepared from CHA scaffolds did not exert apparent cytotoxicity towards BMSCs compared to control media (P > 0.05). Flow cytometry analysis also revealed there was no significant difference in the numbers of live or apoptotic cells on the printed scaffolds compared to control cells seeded on culture dishes (P > 0.05; Figure 2L).

3.4 BMSC adhesion on scaffolds Counting the numbers of cells digested from the scaffolds indicated that the seeding efficiency decreased from 92.6% to 78.8% as the porosity increased; the seeding efficiencies of the Group I and II scaffolds were significantly lower than that of the Group IV scaffolds (P < 0.05) at 8 h after seeding (Figure 3A). SEM images (Figure 3C) revealed that BMSCs were firmly adhered to the scaffolds at after 24 h seeding. Some cells on the scaffolds altered to a plate-like morphology in contrast to 19



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the spindle morphology observed in culture dishes, and cells also formed numerous cellular pseudopods attached to the scaffold. To further investigate the biocompatibility of the scaffolds, immunostaining for vinculin (green), F-actin (red) and DNA (blue) was performed 24 h after cells were seeded on the scaffolds or culture dishes (Figure 3D). As shown in Figure 3B, determination of the IOD values for vinculin immunostaining revealed significantly higher levels of focal adhesion for the cells on the scaffolds compared to cells in culture dishes (P < 0.05). However, there was no significant difference in focal adhesion between the cells on the four groups of scaffolds (P > 0.05).

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Figure 3. (A) In vitro seeding efficiency for BMSCs on the scaffolds at 8 h after seeding. (B) Integral optical density values for vinculin expressed by BMSCs at 8 h after seeding on the scaffolds. (αP < 0.05 vs. Group I; βP < 0.05 vs. Group II; γP < 0.05 vs. Group III; δP < 0.05 vs. Group IV). (C) SEM images of the interactions between the scaffold and BMSCs at 24 h after seeding. (D) Confocal laser scanning microscopy images of BMSCs at 1 day after seeding on the scaffolds. Scale bar: 50 µm.

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3.5 BMSC proliferation on scaffolds The CCK8 assay was used to evaluate cell growth on the scaffolds at 1, 4, 7 and 11 days after seeding (Figure 4B). The OD values of all groups increased over time, with significantly higher OD values observed for the scaffolds in printed groups (I, II, III) than non-printed group (IV) at 11 days (P < 0.05). In printed groups, the scaffold with higher porosity had higher OD values at 11 days (P < 0.05). SEM was also used to evaluate cell growth on scaffolds (Figure 4-A). At 1 day after seeding, the cells were sporadically distributed on all scaffolds, by 4 and 7 days the scaffolds were partially or fully covered by cells, and at 11 days, the cells had begun to grow in stacks. Moreover, the scaffold was cut into two parts through midpoint to reveal cross section. We observed that large numbers of cells were growing inside the printed scaffolds at 11 days (Groups I, II, III), while no cells were growing inside the non-printed scaffolds (Groups IV).

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Figure 4. (A) SEM images of BMSCs on the surface of the scaffolds at 1, 4, 7 and 11 days after seeding in vitro. Note cells growing inside the printed scaffolds (Groups I, II, III) at 11 days after seeding. (B) Measurement of BMSC proliferation on scaffolds using the CCK8 assay at 1, 4, 7 and 11 days after seeding. (C) ALP activity of BMSCs on scaffolds after 7 and 14 days after seeding. (αP < 0.05 vs. Group I; βP < 0.05 vs. Group II; γP < 0.05 vs. Group III; δP < 0.05 vs. Group IV). 23



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3.6 BMSC differentiation on scaffolds Intracellular ALP activity was measured at 7 and 14 days after seeding (Figure 4C). ALP expression was not significantly different between the four scaffold groups (P > 0.05). However, ALP expression was significantly higher for the cells in all four scaffold groups than Group N cells in culture dishes (P < 0.05). The transcription levels of the osteoblast-specific genes COL1A1, ALP, Runx2 and OPN were evaluated via qRT-PCR at 3, 7 and 14 days (Figure 5). No significant differences were observed at 3 days between any groups. At 7 days, the transcription of COL1A1, ALP and Runx2 were significantly higher for the cells in all four scaffold groups than Group N cells in culture dishes (P < 0.05). At 14 days, COL1A1, ALP, Runx2 and OPN were expressed at higher levels by the cells in all four scaffold groups than Group N (P < 0.05); however, the levels of these four genes were not significantly different between the four scaffold groups at any time point (P > 0.05).

Figure 5. Quantitative RT-PCR analysis of osteoblast-specific gene expression by BMSCs on the surface of 24


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scaffolds at 3, 7 and 14 days after seeding in vitro. (A) Collagen type I alpha 1, (B) Alkaline phosphatase, (C) Runt-related transcription factor-2, (D) Osteopontin. (αP < 0.05 vs. Group I; βP < 0.05 vs. Group II; γP < 0.05 vs. Group III; δP < 0.05 vs. Group IV).

3.7 Micro-CT analysis Three dimensional stereoscopic pictures of the implants (blue) and bone tissue (yellow) were reconstructed from micro-CT projections (Figure 6A). The microstructural parameters are summarized in Figure 6B and C. At the 2-week time point, there was no significant difference in BV/TV or RSV/SV between the four scaffold groups (P > 0.05). Analysis of the degradation rate (RSV/SV) revealed that the Group I and II scaffolds had degraded more than the Group III and IV scaffolds at 4 weeks (P < 0.05). All scaffold groups (I, II, III, IV) had significantly higher BV/TV values than the blank controls (Group N) at 4 weeks and subsequent time points (P < 0.05). By 8 weeks, most scaffolds had completely degraded; the RSV/SV values were not significantly different between the four scaffold groups at 8 weeks and subsequent time points (P > 0.05), Additionally, the BV/TV values of the printed scaffolds (Groups I, II, III) were significantly higher than that of the non-printed scaffolds (Group IV) at 8 weeks and subsequent time points (P < 0.05). At 12 weeks, Group I and II had significantly higher BV/TV values than Group III (P < 0.05); however, the BV/TV values of Group I and II were not significantly different at any time points (P > 0.05).

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Figure 6. (A) Micro-CT 3D reconstruction images of the scaffolds and new bone formation at 2, 4, 8 and 12 weeks after in vivo implantation in the rabbit femoral condyle defect model. Scale bar: 1 mm. (B, 26


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C) Quantitative analysis based on the micro-CT evaluation. (αP < 0.05 vs. Group I; βP < 0.05 vs. Group II; γP < 0.05 vs. Group III; δP < 0.05 vs. Group IV; θP < 0.05 vs. Group N).

3.8 Histological analysis At 8 weeks after implantation, MAR was evaluated by measuring the interval distance between the centers of the yellow (tetracycline) and green (calcein) bands (Figure 7). The distance between bands in scaffolds Groups (I, II, III, IV) were significantly wider than those in Group N (P < 0.05). Although the distance between bands was slightly larger in Groups I, II and III than Group IV, the differences were not significant (P > 0.05). Histological analysis of the bone/implant interface was performed using Von-Gieson staining to assess the osteointegration of the implants at 2, 4, 8 and 12 weeks after implantation (Figure 8). By 2 weeks after implantation, numerous cells and fibrous tissue (blue) surrounded the scaffolds (black) and, moreover, were present within the pores of the printed scaffolds (Group I, II, III), although no obvious newly-formed bone was observed. At 4 weeks, the scaffolds had partially degraded and the original outline of the scaffolds became unclear. Tiny newly-formed bone trabeculae (red) were present inside the scaffolds. At 8 weeks, most scaffolds had degraded and bone trabeculae could be easily observed. Additionally, a larger number of bone trabeculae were observed in the printed scaffolds (Group I, II, III) than non-printed scaffold (Group IV). At 12 weeks, the scaffolds were almost completely degraded. Mature bone trabeculae repairing the defect were observed in the Group I, II and III scaffolds. 27



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However, the spaces created by degradation of the scaffolds Group IV were not fully replaced with bone tissue. In contrast, only a small amount of bone tissue was observed within the defects in Group N at any time point.

Figure 7. (A) Fluorochrome double-labeling for tetracycline (yellow) and calcein (green) at 8 weeks after in vivo implementation in the rabbit femoral condyle defect model. (B) mineral apposition rate of new bone formation of each group at 8 weeks. Scale bar: 100 µm. (αP < 0.05 vs. Group I; βP < 0.05 vs. Group II; γP < 0.05 vs. Group III; δP < 0.05 vs. Group IV).

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Figure 8. Histological analysis of new bone formation around and within the scaffolds in the rabbit femoral condyle defect model. After Van Gieson staining, newly formed bone stains red; the tissue stained dark blue is fibrous tissue. Scale bar: 100 µm.

4. DISCUSSION CHA offers several biological advantages as a bone analogue and has been widely applied in the clinic as a commercial product

23

. Although a variety of

materials have been processed into porous 3D scaffolds using AM techniques, CHA materials have not previously been simply printed into 3D scaffolds without other comprise measures

21, 40

. In this study, a modified robocasting printer was adopted to

create CHA composite of bone graft substitutes at low temperature for tissue engineering applications. Our in vitro and in vivo analyses demonstrated that these printed CHA scaffolds combine the biological advantages of CHA materials with the 29



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superior 3D structure constructed by AM techniques. Low temperature 3D printing allows bioactive molecules or drugs that improve bone healing or combat infection to be incorporated into printed scaffolds

41

. To

achieve this aim, the physical properties of the materials, such as powder flowability, roughness, wettability and binder viscosity, must meet the requirements of powder-based printing technology, which limits the types of materials that can be printed at low temperature

13

. Additional processing unavoidably alters the primary

properties of materials and limits the clinical use of the resulting scaffolds. Given that a large variety of bone repair materials are currently available, devising a processing method with minimal printing material limitations to maximize preservation of bioactive properties may represent a convenient and efficient processing approach. Here, to assess the potential of an existing modified robocasting process, a previously-reported CHA material

14, 31

was printed into 3D scaffolds. Using this

modified robocasting approach, experimental 3D structures, as well as large, complicated scaffolds suitable for clinical application could be printed (Figure 1C, D, E). Scaffolds with varied rod widths (Groups I, II, III) were printed to identify the appropriate printing parameters for CHA materials. Based on previous research

35

,

genipin was applied for crosslinking, to enhance the mechanical properties and maintain the structure of the scaffolds. In powder-based printing, the materials are mixed during printing; however, in this study, the collagen and HA were mixed before printing. The three dimensional scaffolds printed using robocasting can be regarded as a structure consisting of 30


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numerous small micron-grade cylindrical non-printed scaffolds. These “small non-printed scaffolds” are similar to the non-printed scaffolds created as a control (Group IV); both have a relatively smooth surface created by the compression of the smooth nozzle and mold, respectively. Internally, the scaffolds presented a plate-like crystal microstructure with the typical characteristics of CHA materials in which the pore size is less than 100 µm and created by the sublimation of the ice crystals during lyophilization

23

. The uniform chemical element distribution of C (indicative of

collagen) and Ca (indicative of HA) demonstrated all parts of the scaffolds have consistent bioactive properties. Contact-angle measurements demonstrated the materials had good wettability, which could promote cell adhesion, proliferation and differentiation 42, 43. The compressive modulus of this material is about 0.1MPa which still can not match that of cancellous bone (about 2-20MPa)

23

. The compressive

modulus of Group I and II scaffolds were lower than those of Group III and IV scaffolds, which reflects the higher porosity of the Group I and II scaffolds. Interestingly, the compressive modulus of Group III was slightly higher than that of the Group IV scaffolds, though this difference was not significant. It is possible that the ordered printed structure partially offsets the decrease in mechanical strength associated with the increased porosity of the printed scaffolds. Assessment of the safety of scaffolds is of primary importance. Both the CCK8 assay and Annexin V-FITC/PI staining demonstrated the scaffolds were non-toxic to BMSCs in vitro. Measurement of seeding efficiency demonstrated that more BMSCs adhered to the Group IV scaffolds than the Group I, II (P < 0.5) or III (P > 0.05) 31



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scaffolds. Seeding efficiency decreased as the porosity of the printed scaffolds increased, consistent with other research

44

. Further evaluations via SEM and

immunostaining revealed that BMSC morphology and adhesion ability were similar on all four types of scaffolds; we concluded the decrease in seeding efficiency as porosity increased was associated with our observation that small amount of the rich-cells culture medium flowed out from printed scaffolds through the printed pore on seeding. BMSCs are reported to have higher adhesion ability towards CHA materials compared to untreated culture dishes, as indicated by assessing the mean IOD values for vinculin

45

. Both the CCK8 assay and SEM indicated that BMSCs

proliferated on the scaffolds and, after approximately 11 days, the cells completely enveloped the scaffolds. Although the printed scaffolds had relatively lower seeding efficiencies than the non-printed scaffolds, the higher porosity of the printed scaffolds promoted BMSC proliferation more significantly over time. BMSCs had limited ability to penetrate the non-printed scaffolds, whereas analysis of cross sections via SEM indicated BMSCs were widely distributed inside the printed scaffolds with 400 µm, which is in accord with previous report

46

. As we expected, all four groups of

CHA scaffolds enhanced the osteogenic differentiation of BMSCs compared to the Group N cells on culture dish, as indicated by ALP activity assays and qRT-PCR. A previous study indicated that a lower scaffold porosity enhanced osteogenesis in vitro 47

; however, porosity had no significant effect on the osteogenic ability of the cells

seeded on the four groups of CHA scaffolds. The previous studies mainly focused on the micropore diameter, which was less than 300 µm. In present study, the increased 32


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porosity of the printed scaffolds was mostly due to an increase in the number of 400 µm-wide macropores. As mentioned above, the “small non-printed scaffolds” in Group I, II and III had a similar microstructure as the non-printed scaffolds of Group IV; therefore, it is reasonable that BMSCs on all four scaffold groups had similar ALP activity and transcription of osteoblast-specific genes. As osteogenic outcome is a product of the number of BMSCs, we suppose that scaffolds containing a higher numbers of BMSCs will have a stronger osteogenic outcome. Nevertheless, both printed and non-printed CHA proved biocompatible with BMSCs and the printed scaffolds demonstrated better properties of promoting BMSCs proliferation and osteogenic differentiation than non-printed scaffolds in vitro. Since the 3D CHA scaffolds had good biocompatibility in vitro, a preclinical rabbit femoral condyle defect model was conducted to assess their bone-defect repair ability in vivo. Micro CT revealed that, compared to the blank control Group N, CHA facilitated new bone growth within the scaffolds, as the scaffolds were resorbed or incorporated into newly-forming bone. A good scaffold should be degraded at an appropriate rate consistent with the process of tissue repair

48

. Van Gieson staining

demonstrated that the non-printed scaffolds (Group IV) promoted a pattern of repair from the edges towards the center of defect, since cells could only penetrate the scaffold a limited distance. If newly tissue cannot timely form before scaffold is degraded due to the complicated in vivo environment, a defect is likely to be left in the core of the tunnel drilled in the bone. As 400 µm-wide pores interconnect within the printed scaffolds, BMSCs could migrate into all regions of the scaffold and 33



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mineralize at the same time. Hence, the osteogenic property of the CHA material was further improved by creating a 3D scaffold via the AM approach. In summary, the current approach provides a method of printing CHA materials without additional chemical or physical processing of the raw material. In both in vitro and in vivo tests, the printed CHA scaffolds had the advantages of promoting BMSCs proliferation and differentiation and promoting defect repair compared to the non-printed CHA scaffolds. Considering the mechanical properties and osteogenic outcome of the three types of printed scaffolds created using the existing CHA material tested, the printing diameter of the Group II scaffolds (600 µm) resulted in the moderate mechanical strength, appropriate degradation rate and excellent repair outcome. One limitation of the present study is the weak mechanical strength of CHA materials. However, these materials are likely to be suitable for the repair of low load-bearing bone defects or cancellous bone defects

23, 49

. In this study, we achieved

our objectives of printing a scaffold using existing materials without additional chemical or physical processing of the raw material and confirmed the resulting printed scaffolds promoted a better osteogenic outcome. Future studies will focus on incorporating growth factors or drugs into the printed scaffolds. Additionally, the use of other materials such as autopolymerizing materials could be investigated.

5. Conclusion In this study, a low temperature modified robocasting method was applied to print biomimic 3D scaffolds for bone regeneration using chemical solvent-free CHA 34


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composite materials. This method combines the biological advantages of CHA materials with the superior 3D structure provided by AM techniques. Future studies are required to confirm whether incorporating growth factors or drugs into such printed scaffolds could improve bone healing efficacy or combat infection. Besides, the potential of other excellent bioactive materials remains to be assessed.

ASSOCIATED CONTENT Supporting Information Morphology of third-passage rabbit BMSCs. Flow cytometric analysis of CD34/CD44 expression ratio for third-passage rabbit bone marrow stromal cells.

AUTHOR INFORMATION Corresponding Authors *Phone: 86-029-84775289. Fax: 86-029-84775289. E-mail: [email protected] *Phone: 86-029-84771860. Fax: 86-029-84771860. E-mail: [email protected] Author Contributions

§These authors contributed equally to this study. Funding Sources National Natural Science Foundation of China, grant number: 81371982. Notes The authors declare no competing financial interest.

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ACKNOWLEDGEMENTS We thank Dr. Qingfeng Zeng and Mr. Xin Zhang from Xi’an Particle Cloud Biotechnology Company for their assistance with FFP printing and fruitful discussion on CHA composite materials.

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Figure 1: (A) Schematic of the robocasting fabrication process. (B) Gross view of the surface morphology of the experimental collagen/hydroxyapatite (CHA) composite scaffolds. The printed scaffolds had a grid-like microstructure with pore widths of 400 µm and rod widths of 300 µm (Group I), 600 µm (Group II) or 900 µm (Group III). Group IV was non-printed scaffolds. Gross views of the structure of example printed complicated scaffolds: bone with marrow cavity (C), femoral head (D) and tibial plateau (E). Scale bar: 5 mm 91x46mm (300 x 300 DPI)


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Figure 2. SEM images of the surface morphology (A) and internal microstructure (B) of the printed CHA scaffolds, and the surface morphology (C) and internal microstructure (D) of the non-printed CHA scaffolds. Element distribution of calcium (E) and carbon (F) within the scaffolds and the corresponding merged images (G). Water contact angle analysis (H) for the four groups of scaffolds. Porosity (I) and compressive modulus (J) for the four groups of scaffolds (αP < 0.05 vs. Group I; βP < 0.05 vs. Group II; γP < 0.05 vs. Group III; δP < 0.05 vs. Group IV). CCK8 assay (K) of the cytotoxicity of conditioned media prepared from CHA scaffolds (1 day incubation) or control media towards BMSCs. Assessment of BMSC survival and apoptosis at 1 day after seeding on the scaffolds or culture dishes (control) using the Annexin V-FITC/PI double-staining assay (L). 138x143mm (300 x 300 DPI)



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Figure 3. (A) In vitro seeding efficiency for BMSCs on the scaffolds at 8 h after seeding. (B) Integral optical density values for vinculin expressed by BMSCs at 8 h after seeding on the scaffolds. (αP < 0.05 vs. Group I; βP < 0.05 vs. Group II; γP < 0.05 vs. Group III; δP < 0.05 vs. Group IV). (C) SEM images of the interactions between the scaffold and BMSCs at 24 h after seeding. (D) Confocal laser scanning microscopy images of BMSCs at 1 day after seeding on the scaffolds. Scale bar: 50 µm. 205x237mm (300 x 300 DPI)


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Figure 4. (A) SEM images of BMSCs on the surface of the scaffolds at 1, 4, 7 and 11 days after seeding in vitro. Note cells growing inside the printed scaffolds (Groups I, II, III) at 11 days after seeding. (B) Measurement of BMSC proliferation on scaffolds using the CCK8 assay at 1, 4, 7 and 11 days after seeding. (C) ALP activity of BMSCs on scaffolds after 7 and 14 days after seeding. (αP < 0.05 vs. Group I; βP < 0.05 vs. Group II; γP < 0.05 vs. Group III; δP < 0.05 vs. Group IV). 219x270mm (300 x 300 DPI)



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Figure 5. Quantitative RT-PCR analysis of osteoblast-specific gene expression by BMSCs on the surface of scaffolds at 3, 7 and 14 days after seeding in vitro. (A) Collagen type I alpha 1, (B) Alkaline phosphatase, (C) Runtrelated transcription factor-2, (D) Osteopontin. (αP < 0.05 vs. Group I; βP < 0.05 vs. Group II; γP < 0.05 vs. Group III; δP < 0.05 vs. Group IV). 68x55mm (600 x 600 DPI)


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Figure 6. (A) Micro-CT 3D reconstruction images of the scaffolds and new bone formation at 2, 4, 8 and 12 weeks after in vivo implantation in the rabbit femoral condyle defect model. Scale bar: 1 mm. (B, C) Quantitative analysis based on the micro-CT evaluation. (αP < 0.05 vs. Group I; βP < 0.05 vs. Group II; γP < 0.05 vs. Group III; δP < 0.05 vs. Group IV; θP < 0.05 vs. Group N). 209x313mm (300 x 300 DPI)



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Figure 7. (A) Fluorochrome double-labeling for tetracycline (yellow) and calcein (green) at 8 weeks after in vivo implementation in the rabbit femoral condyle defect model. (B) mineral apposition rate of new bone formation of each group at 8 weeks. Scale bar: 100 µm. (αP < 0.05 vs. Group I; βP < 0.05 vs. Group II; γP < 0.05 vs. Group III; δP < 0.05 vs. Group IV). 94x104mm (300 x 300 DPI)


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Figure 8. Histological analysis of new bone formation around and within the scaffolds in the rabbit femoral condyle defect model. After Van Gieson staining, newly formed bone stains red; the tissue stained dark blue is fibrous tissue. Scale bar: 100 µm. 112x71mm (300 x 300 DPI)



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Low-Temperature Additive Manufacturing of Biomimic Three

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