In vivo and in vitro study of a novel nanohydroxyapatite sonocoated scaffolds for enhanced bone regeneration

In vivo and in vitro study of a novel nanohydroxyapatite sonocoated scaffolds for enhanced bone regeneration

Accepted Manuscript In vivo and in vitro study of a novel nanohydroxyapatite sonocoated scaffolds for enhanced bone regeneration Julia Rogowska-Tylma...

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Accepted Manuscript In vivo and in vitro study of a novel nanohydroxyapatite sonocoated scaffolds for enhanced bone regeneration

Julia Rogowska-Tylman, Janis Locs, Ilze Salma, Bartosz Woźniak, Mara Pilmane, Vita Zalite, Jacek Wojnarowicz, Aleksandra Kędzierska-Sar, Tadeusz Chudoba, Karol Szlązak, Adrian Chlanda, Wojciech Święszkowski, Aharon Gedanken, Witold Łojkowski PII: DOI: Reference:

S0928-4931(18)31424-3 https://doi.org/10.1016/j.msec.2019.01.084 MSC 9342

To appear in:

Materials Science & Engineering C

Received date: Revised date: Accepted date:

28 May 2018 13 January 2019 14 January 2019

Please cite this article as: J. Rogowska-Tylman, J. Locs, I. Salma, et al., In vivo and in vitro study of a novel nanohydroxyapatite sonocoated scaffolds for enhanced bone regeneration, Materials Science & Engineering C, https://doi.org/10.1016/j.msec.2019.01.084

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In vivo and in vitro study of a novel nanohydroxyapatite sonocoated scaffolds for enhanced bone regeneration

Julia Rogowska-Tylmana,b, Janis Locsc, Ilze Salmad, Bartosz Woźniaka, Mara Pilmanee , Vita Zalitec, Jacek Wojnarowicza, Aleksandra Kędzierska-Sara, Tadeusz Chudobaa, Karol Szlązakb, Adrian

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Chlandab, Wojciech Święszkowskib, Aharon Gedankenf, Witold Łojkowskia

Institute of High Pressure Physics, Polish Academy of Sciences, Warsaw, Poland

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Faculty of Materials Science and Engineering, Warsaw University of Technology, Warsaw, Poland

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Rudolfs Cimdins Riga Biomaterials Innovations and Development Centre, Institute of General Chemical

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Engineering, Faculty of Materials Science and Applied Chemistry, Riga Technical University, Riga, Latvia Department of Oral and Maxillofacial Surgery, Riga Stradins University, Riga, Latvia

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Faculty of Anatomy and Anthropology, Riga Stradins University, Riga, Latvia

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Department of Chemistry, Bar-Ilan University, Ramat Gan, Israel

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ABSTRACT

There still remains a need for new methods of healing large bone defects, i.e., gaps in

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bone tissue that are too big to naturally heal. Bone regrowth scaffolds can fill the bone gap and enhance the bone regeneration by providing cells with a support to for new tissue formation.

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Coating of the scaffolds surface with nanocrystalline hydroxyapatite may enhance the

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osteoinductivity or osteoconductivity of such scaffolds. Here we present the sonocoating method to coat scaffolds with bioactive hydroxyapatite nanoparticles. We show a method, where the material to be coated is immersed in a colloidal suspension of nanoparticles with mean sizes of 10 nm and 43 nm in water, and high-power ultrasound waves are applied to the suspension for 15 min at 30°C. High power ultrasounds lead to growth of cavitation bubbles in liquid, which implode at a critical size. The implosion energy propels the nanoparticles towards the material surface, causing their attachment to the scaffold. Using this technique, we produced a uniform layer of nanohydroxyapatite particles of thickness in the range 200 to 300 nm on two types of

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scaffolds: a porous -TCP ceramic scaffold and a 3D-printed scaffold made of PCL fibers. In vivo tests in rabbits confirmed that the novel coating strongly stimulated new bone tissue formation, with new bone tissue occupying 33% for the nHAP-coated PCL scaffold and 68% for the nHAPcoated -TCP after a 3-month test. The sonocoating method leads to formation of a bioactive

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layer on the scaffolds at temperature close to room temperature, very short time and in water. It

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is a green technological process, promising for bone tissue regeneration applications.

1. Introduction

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Many studies have investigated the problem of healing bone defects, which are too large to naturally heal [1,2]. Such defects may occur following tumor extraction operations, or trauma

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due to various accidents. A common treatment approach involves the use of allografts, i.e., donor tissues or autografts from the same individual. However, this method requires two operations, can

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infection and immune rejection [3,4].

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cause a bone defect in another part of the patient’s body, and carries risks of post-operative

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Various materials have been developed as alternatives to autografts or allografts, including bone repair scaffolds and cements [5–10]. Cements are mainly used for small gaps;

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they prevent soft tissues from filling the bone cavity, stimulate bone regrowth, and are gradually resorbed. Scaffolds provide mechanical support for the cells, enabling to form new tissue in defect

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site [9–19]. To support bone regeneration, a scaffold should present mechanical stability, good biocompatibility, controllable biodegradability, osteoconductivity (i.e., ability to stimulate active bone cell growth), and no cytotoxic effects towards the surrounding tissues [20–23]. However, many challenges have arisen in designing suitable scaffold materials, including frequent bone infections, insufficient vascularization, bioactivity, and osseointegration [24–26]. Numerous materials have been tested for this purpose, many based on biodegradable polymers and calcium phosphate ceramics [26–30]. Modification of such materials with bioactive layers provides a

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natural means of enhancing a material’s biocompatibility and bioactivity by changing surface properties [31], especially when hydroxyapatite (HAP) or nanocrystalline hydroxyapatite (nHAP) are used as coating materials [32–36]. Scaffolds coated with nHAP can reportedly shorten the defect healing process [37–43], and such coating can also exhibit osteoinductive properties [44– 46]. Compared to their microcrystalline counterparts, nanocrystalline HAP particles display better

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bioactivity due to their high specific area and enhanced solubility, which increases the calcium ion

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concentration in the surrounding environment [47,48]. There were also many strategies reported involving mesenchymal stem cells, growth factors, etc. for the fabrication of the bone grafting

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scaffolds [49-50]. Nevertheless, none of the research works were able to identify the fast and reproducible way for bone tissue regeneration scaffold fabrication.

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Multiple methods have been developed to deposit nHAP on implants, including simulated

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body fluid, electrophoretic coating, sputtering, dip coating, spin coating, and plasma spraying. Nevertheless, coatings obtained by such methods show limitations such as layer homogeneity in

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nanoscale, layer adherence and biological response in vitro and in vivo [10,17,20,29,32,33]. Electrodeposition provide coatings of the micron-size thickness with flower and flake-like

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morphologies. Pulsed laser deposition method enables to obtain island-like agglomerated apatite

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coatings of 2-10 µm thickness. Simulated body fluid immersion in supersaturated medium is one of the most commonly used techniques of apatite deposition on implants. Layers produced by this

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method are precipitated in various topographies, such as: flower-like, plate-like, needle-like structures of high brittleness [10,17,20,29]. The SBF deposited structures show limitations in biological response due to micron-size roughness. Literature shows that for PCL 3D-printed scaffolds, SBF did not result in increased osteoconductivity and cell proliferation as previously reported by many authors [51,52]. The aim of this study was to show that bioactive coatings significantly enhancing bone tissue growth can be produced using sonocoating method. Sonocoating is a technique, where the object

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to be coated is immersed in a colloidal suspension of nanoparticles, and high-power ultrasounds are applied to the said suspension [53,54]. The technique was applied to sonocoat two types of scaffolds with a nanometric thickness layer composed of nHAP. The layer should exhibit optimum surface roughness, nanoscale homogeneity, and sufficient adhesion to substrates. The additional aim of the study was to perform the coating process at temperature of 30°C, suitable for

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temperature-sensitive materials, such as biodegradable polymers. Low temperature of the

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process opens the way to possible use of antibiotics or bone growth factors during the nHAP deposition process [53,54]. Low-temperature deposition is also important for preserving the

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thermodynamically unstable nanostructure of nHAP particles, which grain size may increase in higher temperatures and the particles may lose their unique biological properties.

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In our present study, we examined the potential of sonocoating - i.e., using a sonochemical

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method to coat porous scaffolds with nHAP particles at temperatures below 40°C. With the sonocoating technique, materials are coated via immersion in a colloidal suspension of

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nanoparticles and application of high-energy ultrasound waves that cause acoustic cavitation

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nearby the coated object [53-55]. In recent years, much focus was visible towards biodegradable bone regeneration biomaterials. Here we tested the application of sonocoating on two types of

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bone scaffolds: a 3D-printed poly-ɛ-caprolactone (PCL) structure and a porous β-tricalcium phosphate (β-TCP) ceramics. PCL is a polymer that has recently attracted much interest in the

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field of tissue engineering, since it can be used in 3D printing, exhibits relatively high mechanical strength, good biodegradability, and a low melting point [56-58]. PCL is a polymer with the degradation rate matching the pace of new bone tissue formation (the total degradation time up to 2–3 years) but its surface hydrophobicity is not favorable for cells attachment [59-61]. However, it was previously reported by Heo et al. and Domingos et al. that 3D printed PCL, when combined with nHA ceramics can exhibit better bone-like and human mesenchymal stem cells (hMSC)

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adhesion, proliferation and osteogenic differentiation in comparison to PCL with micron size hydroxyapatite [61,62]. Currently available treatments for large bone defects often fail in recreating the anatomical features and function of the lost tissue due to the complexity in structure and properties of bone in the population [2,21]. 3D printing used for scaffold preparation shows great perspectives for

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the fabrication of patient-specific bone grafts, since it allows customizing the architecture of

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implants, e.g. producing different lay-down patterns of the fibers with varying the layer deposition angles [57,62]. Among the 3D printed PCL scaffolds with different architectures (e.g. 0°/15°/30°,

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0°/30°/60°, 0°/45°/90°, 0°/60°/120°, 0°/75°/150°,0°/90°/180°) we have chosen the structure composed of 0o/60o/120o deposited layers as a promising construct due to combined features:

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high elastic modulus, dynamic stiffness, optimal size of pores, good cells long-term cells response and bone tissue regeneration in vivo in New Zealand rabbit animal model [49,57-58]. Additionally,

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such structure provides sufficient surface area and porosity for the structures penetration and

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modification with nanoparticles. As a reference material we selected β-TCP porous ceramics due to its well-known high bioactivity and favorable mechanical properties [12,14]. Since, the

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bioactivity and biocompatibility of biomaterials are related to their cell attachment and proliferation

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behavior, we selected materials which show great promise for tissue engineering from the good cellular activity point of view [37,38,56–62]. For the in vitro study we have selected human

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osteosarcoma cell line MG-63 due to wide use and well defined protocols for bone scaffolds testing [60,62,63]. The aim of our investigation was to determine whether sonocoating could effectively create a uniform nanometric nHAP coating, comprising bioactive nanoparticles on the porous scaffolds, and promote bone-like cells (MG-63) attachment and proliferation behavior as well as new bone tissue growth in vivo.

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2. Materials and methods 2.1.

Hydroxyapatite nanoparticles synthesis Hydroxyapatite nanocrystals were obtained via a hydrothermal reaction using microwave

energy as previously described by Kuśnieruk et al. [65] and Smoleń et al. [66]. The synthesis

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precursors were calcium hydroxide (Ca[OH]2, pure; CHEMPUR, Poland) and phosphoric acid

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(H3PO4, 85% solution, pure for analysis; CHEMPUR, Poland), which were mixed at a Ca/P molar

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ratio of 1.67. The reaction solution was prepared by drop-wise addition of phosphoric acid to calcium hydroxide at a rate of 0.01 ml every 3 seconds, with intensive stirring of the solution at

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ambient temperature. After 30 minutes of additional stirring, the reaction solution was transferred into a Teflon vessel, which was heated under microwave radiation with a frequency of 2.45 GHz

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in a microwave reactor Ertec Magnum II (Ertec Poland) [65,67,68].

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Microwave technology enabled the delivery of an extremely high-power density of 5 W/ml to the reactant solution, heating the precursors at a rate of 1 K/s. The fast uniform heating and

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the ability to immediately stop the reaction by turning off the microwave power allowed precise control of the synthesis time. This permitted to obtain a narrow particle size distribution. The

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obtained colloidal suspension solutions were transferred to a plastic bottle, and stored at 4°C

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[65,68].

The Laboratory of Nanostructures of IHPP produces six types of nHAP (registered in

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Polish Patent office as GoHAP™) [67], which have mean grain sizes ranging from 8 to 43 nm depending on the synthesis time. Two nHAP types were selected for the present study. We used GoHAP type II nanoparticles, further called nHAP2, which are synthesized in 90 seconds, and are characterized by a mean grain size of 10 ± 1 nm, specific surface area of 200 ± 20 m2/g, Ca/P ratio of 1.61 ±0.04, and calcium ion solubility of 20 mg/dm3 during a 28-day test according to ISO 10993-14 norm. We also used GoHAP type VI nanoparticles, further called nHAP6 which are synthesized in 600 s, and are characterized by a mean grain size of 43 ± 4 nm, specific surface

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area of 49 ± 5m2/g, Ca/P ratio of 1.61 ± 0.04, and calcium ion solubility of 4.6 mg/dm3 during a 28-day test according to ISO 10993-14 norm [69]. We have used of two types of nHAP nanoparticles in an attempt to find whether there are any differences visible in biological response of the coating. The X-ray diffraction patterns and other characteristics of nHAP2 and nHAP6 nanoparticles correspond very well to natural bone. A detailed description of the nHAP

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nanostructures and synthesis details has been previously published [65,68].

Scaffold fabrication and characterization

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2.2.1. Ceramic scaffold fabrication

Porous ceramic β-TCP scaffolds were prepared from as-synthesized calcium-deficient

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hydroxyapatite (CDHAP) powder with a Ca/P ratio of 1.5, which was produced at the Rudolfs

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Cimdins Riga Biomaterials Innovations and Development Centre of Riga Technical University following the method described by Loca et al. [70]. A wet precipitation synthesis reaction was

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performed between calcium hydroxide (0.45 M Ca(OH)2 suspension) and orthophosphoric acid (2.00 M H3PO4 solution). After synthesis, the precipitates were vacuum filtered and dried in an

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oven [71,72]. To form the scaffold structure, CDHAP powder was mixed with a glycerol solution

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(purity > 99.8%; BIO-VENTA Ltd., Latvia), distilled water, and ammonium bicarbonate (NH4HCO3; Ltg. Enola, ES/BASF). The foaming process was conducted in a drying oven at temperatures

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ranging from 80°C to 120°C. To ensure the transformation of CDHAP to a β-TCP porous ceramic scaffold, the foam was sintered at 1200°C for 2 hours.

2.2.2. Polymer scaffold fabrication PCL scaffolds were fabricated by using a 3D printer (Bioscaffolder; SYSENG; Germany). Poly-ɛ-caprolactone polymer (PCL; MW 70,000–90,000) was purchased from Sigma-Aldrich, USA. Scaffolds were printed as cylindrical forms with a height of 4 mm and diameter of 5 mm.

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Every structure comprised 18 layers, of 7 fibers each, with a layer-to-layer angle alignment of 0°/60°/120°. The average fiber thickness was 200m.

2.2.3. Sonocoating procedure

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Sonocoating is a method of depositing nanoparticles on a material’s surface using high

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intensity power ultrasounds [73-75]. The material to be coated is submerged in a colloidal

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suspension of nanoparticles, and ultrasonic waves are applied to cause cyclic changes of the acoustic pressure in the suspension. At high ultrasound power, vapor voids, called cavitation

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bubbles, are formed [53,54], which grow throughout the pressure oscillations until violent

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implosion occurs at a critical bubble size. Implosions close to the substrate surface generate liquid micro-jets that propel the nanoparticles towards the material surface. Fig. 1 shows a schematic

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of the sonocoating process.

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Prior to sonocoating, the β-TCP and PCL scaffolds were cleaned by rinsing with a mixture of distilled water and ethanol (75:25), followed by additional cleaning in an ultrasonic bath for 5

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min, and drying under air in a laminar flow cabinet at room temperature. Next, the scaffolds were

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mounted in a Teflon handle arm, and submerged in a vessel with a nanoparticle concentration of 0.1 (w/v%). An ultrasonic Ti horn (Sonics & Materials VCX750, USA) was immersed 2 cm above

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the scaffold (Fig.1) and placed in sound absorbing chamber. To prevent heating of the suspension due to ultrasonic energy absorption, the process temperature was stabilized at 30°C. After a coating time of 15 min, the coated samples were rinsed with distilled water, and dried under air in a laminar flow cabinet at room temperature. Subsequently, the samples were vacuum packed and stored under the hood. The ultrasonic coating method has been previously described in detail [75]. To investigate the structure of the layer, PCL in a form of flat pellets coated with nHAP under the same conditions as the above scaffolds were prepared.

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Fig. 1. Schematic presentation of the sonocoating process.

2.2.4. Examination of scaffold structure

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To investigate scaffolds morphology, we used a Field Emission Scanning Electron

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Microscope (FE-SEM; Ultra Plus GEMINI; Carl Zeiss, Germany) with operating voltage 2 kV and 3 kV. During SEM investigations of the PCL and β-TCP scaffolds, we additionally examined

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fracture surfaces. To investigate the pore structure of the scaffolds, we performed microcomputed tomography (µ-CT) using an X-Ray micro-CT (SkyScan 1172, Bruker,Germany). For SEM investigations of the 3D-printed PCL scaffolds, the outer layers of fibers were removed to reveal inner structures. To evaluate the homogeneity and thickness of the generated layers, we analyzed a series of SEM microscopic images taken using the FE-SEM InLens and SE detectors. Samples were fixed on the SEM imaging table with carbon tape and sealed with conductive adhesive (Ag glue) to create an effective ground for the object. Sputter coating was

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used to create conductive surfaces (sputter coating machine SCD 005/CEA 035; BAL-TEC, Pfäffikon, Switzerland). Analysis of nHAP layers was conducted using ImageJ software (developed by the National Institute of Health, USA). For cells morphology imaging PCL pellet samples coated with nHAP were used. Samples were fixed in cell lab using 4% paraformaldehyde and dehydrated by a graded ethanol series from

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10 to 100%, with three times 20 min incubation at each step. They were sputter coated with a

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gold layer before microscopic investigations.

To roughly assess layer adhesion to the PCL fibers, several fibers were manually

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stretched, and cut with scalpel before subsequent investigation by SEM. To evaluate the coating thickness, the scaffolds were brittle fractured in liquid nitrogen prior imaging.

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Surface topography examinations were performed by means of Atomic Force Microscopy (AFM) Bruker MultiMode (NanoScope 8, Bruker, Germany) equipped with NanoScope 5

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controller. The analysis was performed in oscillating mode of the microscope (Tapping Mode)

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using APP NANO (ACST type) scanning probe working at the frequency of 145 kHz (calibrated prior to examination). Roughness data were collected from 10 randomly selected areas of 5 μm

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x 5 μm and the average value - Ra was calculated. For the characterization of samples' surface mechanical properties at the nanoscale, the PeakForce Quantitative NanoMechanics (PF-QNM)

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AFM mode was used. It allows to evaluate samples’ stiffness, adhesion, dissipation and

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deformation simultaneously with topographical imaging of the surface. In the presented work, Derjaguin-Muller-Toporov (DMT) model was used for the quantification of the stiffness of tested samples [76-77]. The topographical images along with nanomechanical maps obtained during the acquisition were developed and analyzed in the NanoScope Analysis version 1.40 (Bruker, Germany) software. The wettability of the prepared non-coated and nHAP coated surfaces was measured by means of a contact angle system (OCA, Dataphysics, Germany) at ambient temperature with a

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10 µl water droplet (24h conditioning of samples before measurement). Measurements were collected from 10 randomly selected areas. Phase composition of the coating material was examined using (XRD) X-ray Diffraction (X’Pert PRO, PANalytical) diffractometer equipped with a Cu Kα1 anode and PIXcel1D detector.

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Data were collected in the range 2θ between 10° and 100° with a step of 0.03°, from nHAP powder

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samples before coating and from coated samples. The XRD patterns were compared with data

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from JCPDS-International Centre for Diffraction Data file for crystalline hydroxyapatite spectrum

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(JCPDS, 9-0432).

2.3. In vitro cell culture analysis

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In vitro experiments were performed using the human osteosarcoma cell line MG-63. The cells were cultivated in D-MEM media supplemented with 10% fetal bovine serum (FBS) and

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penicillin/streptomycin, in 24-well cultivation plates at 37°C under 5% CO2. Following cultivation,

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the cells were detached using 0.25% trypsin/EDTA solution. Next, 3D-printed PCL and TCP scaffolds uncoated and coated with nHAP6 and nHAP2 were washed with PBS (pH 7.4), and

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seeded with 105 cells in 200 µl cultivation medium. The seeded scaffolds were incubated for 1.5

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h at 37°C in a 5% CO2 atmosphere. Then cultivation medium was added to the well to completely cover the test material, and the samples were incubated for 5 days.

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After the 5-day culture, we used a CyQUANT cell proliferation assay kit to determine the cell numbers on the test material. This assay has a linear detection range extending from ≤50 cells to at least 50,000 cells. Higher cell concentrations can be assessed by adjusting the dye concentration.

2.4. In vivo tests

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The Animal Ethics Committee of Latvian Food and Veterinary Administration granted approval of an in vivo experiment in 15 six-month-old male rabbits (license Nr. 60/25.10.2013.). General anesthesia was administered by intramuscular injection of ketamine 10% (100 mg/ml; 30 mg per kilogram weight), xylazine 2% (20 mg/ml; 3 mg per kilogram weight), and atropine 0.1% (0.1–0.5 mg per kilogram weight). Under general anesthesia, a 2-cm incision was made on the

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iliac crest tuber sacrale of each leg, the. A trepan drill was used to create a 5-mm round hole in

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the bone, and scaffold was placed in this prepared artificial cavity. After scaffold implantation, the periosteum was stitched, and soft tissues were gradually closed using 5-0 Vicryl® sutures

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(Ethicon). The skin was sutured with 4-0 Prolene® synthetic monofilaments (Ethicon). After this operation, we administered daily subcutaneous injections of 5 mg doses of Enrobioflox 5% (50

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mg/ml solution; Vetoquinol Biowet, Poland) for 5 days. Euthanasia was carried out under general

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anesthesia, by injecting T-611 (1 ml/kg weight) into the lungs.

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2.5. Morphometric, histological, and immunohistochemical analysis of tissue samples At 3 months after implantation, scaffolds and surrounding tissue samples were harvested

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and histologically evaluated for bone regeneration. For morphometric analyses, samples were

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stained with hematoxylin and eosin. The preparations were visualized using a Leica DFC 420 camera and Leica DMRB microscope, along with the Leica Application Suite Software (version

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4.4.0). Quantitative analysis was performed using the Image-Pro Plus 7 Software (Media Cybernetics, USA). We measured the cross-sectional area of bone, biomaterial, and soft tissue. For this analysis, the PCL and β-TCP samples coated with nHAP2 and nHAP6 were integrated as a single group. We analyzed 10 samples from each sample group. Five rabbits were selected for immunohistochemical (IHM) analysis, each with a different implanted scaffold: uncoated PCL, uncoated β-TCP, nHAP-coated PCL, nHAP-coated β-TCP. The collected tissues were fixed using a mixture of 2% formaldehyde and 0.2% picric acid in 0.1 M phosphate buffer (pH 7.2). The samples were rinsed for 12 hours in a thyroid buffer containing

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10% saccharose, then embedded into paraffin, and cut into thin 6- to 7-µm sections. Control samples of host natural tissue were extracted from the rabbits’ bone at a distance of 1–3 mm from the implantation area. The samples were tested for the appearance of the following eight factors: osteocalcin (OC; 1:100; Abcam, UK), osteopontin (OP; 1:100; Abcam, UK), osteoprotegerin (OPG; 1:40;

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Santa Cruz Biotechnology, Inc.), bone morphogenetic protein (BMP 2/4; working dilution 1:100;

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RDS, UK), matrix metalloproteinase 2 (MMP2; rabbit, working dilution 1:100; Biobyt, UK), tissue inhibitor of matrix metalloproteinase 2 (TIMP2; mouse, 1:200 dilution; Santa Cruz Biotechnology,

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Inc., USA), interleukin 1 (IL-1; 1:50; Santa Cruz Biotechnology, Inc., USA), and interleukin 10 (IL10; 1:400; Abbiotec, UK). We evaluated the IHM reactions using a semiquantitative method. The

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following scale was used to indicate the number of immunoreactive structures observed in the visual field: 0, none; +–, few; ++–, a moderate number; +++–, numerous; and ++++–, an

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abundance.

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Spearman’s correlation was calculated to show how strongly one factor influenced another factor: 0.1 – 0.3, weak correlation; 0.3 – 0.7, moderate correlation; and 0.7 – 1.0, strong

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3. Results

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correlation.

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3.1. Scaffold structure

SEM images of PCL 3D-printed scaffolds have shown the organized structure of fibers (Fig. 2A-2D). On β-TCP ceramic scaffolds fracture surfaces, we observed a polycrystalline β-TCP material with a grain size of a few microns, and pores with diameters in the range of 20 – 200 m (Fig. 3A-3D). Representative μ-CT images of scaffolds cross-sections are shown in Figure 4. The β-TCP sample overall porosity was 49%, which is in agreement with the value of 52 ± 2.6% obtained using the Archimedes method. Using μ-CT, we determined that the average β-TCP pore

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size was 205 μm and pore wall thickness was 97 μm. The mean pore surface area of the 3Dprinted PCL scaffold was 0.18 mm2, which corresponds to mean pore diameter 420 μm, and the

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porosity volume fraction was 41.2% (Fig. 4), as determined by means of µ-CT scanning.

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Fig. 2. SEM images of the PCL scaffold, magnifications: A) 100x, B) 1000x, C) 10 000x, D) 25 000x

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Fig. 3. SEM images of the β -TCP scaffold, magnifications: A) 500x, B) 1000x, C) 10 000x, D) 25 000x

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3.2. Ultrasonic coating results

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Fig. 4. Cross-sections of the: A) PCL scaffold, B) β-TCP scaffold, visualized by means of µ-CT scanning.

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Homogenous coatings were obtained on both studied scaffolds (Figs. 5,7). The layer thickness ranged from 200 to 300 nm and surface roughness was in the range of 20 – 30 nm

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(Figs. 6,8). For the β-TCP scaffolds, we observed coating of all open pores. For the PCL 3Dprinted scaffolds, the fibers were coated from all sides, and equal coating was observed on the fibers near the sample surface and those in the interior. The coatings on both scaffolds were characterized by high homogeneity and almost 100% surface coverage. Figures 5 and 6 show SEM and AFM images of the 3D-printed PCL scaffold and the nHAP layer on its surface, revealing uniform coating of all fibers. Figure 7 shows the nHAP coating layer within pores of the β-TCP scaffold.

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Fig. 5. SEM images of the nHAP-coated 3D-printed PCL scaffold. Top view with magnifications: A) 100x, B) 1000x, C) 25 000x, and D) 100 000x cross-section view of the nHAP layer after polymer brittle fracture at liquid nitrogen temperature

Fig. 6. SEM and AFM surface topography images of nHAP-coated 3D printed PCL sample: A) SEM image of surface (scale bar 1m), B) AFM topography image of the surface, C) 3D phase image reconstruction. Ra mean = 20.4 nm  3.1 nm.

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Fig. 7. SEM images of nHAP coating on β-TCP scaffold. Top view with magnifications: A) 100x, B) 1000x, C) 25 000x, and D) 100 000x cross-section view of the nHAP layer after polymer brittle fracture at liquid nitrogen temperature.

Fig. 8. SEM and AFM surface topography images of nHAP-coated -TCP ceramic sample: A) SEM image of surface (scale bar 1m), B) AFM topography image of the surface, C) 3D phase image reconstruction. Ra mean = 26.3 nm  1.4 nm.

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Atomic Force Microscopy (AFM) studies of surface topography of the coating layer revealed a highly developed surface, composed of spherical nanoparticles. The mean surface roughness was in the range 20 – 30 nm (Figs. 6,8). As a result of AFM mechanical testing, data of the surface deformation and reduced Young’s modulus (DMT modulus) were collected. Brighter areas in Fig. 9 correspond to higher values. DMT modulus values increased significantly after nHAP coating

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of PCL (Figs. 9B - 9C). For uncoated PCL, DMTPCL was 501 MPa ± 18 MPa, while for the coated

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surface DMTnHAP/PCL=1771 MPa ± 158 MPa.

Fig. 9. AFM testing of DMT Modulus and Deformation. Images of: A) non-coated PCL sample, B) nHAP coated PCL sample at 5 m x 5 m scanning area, C) nHAP coated PCL sample at 500 nm x 500 nm scanning area.

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X-ray diffraction patterns of nHAP2 and nHAP6 before and after coating process revealed no major differences (Fig.10). XRD patterns were not affected by high-intensity ultrasounds applied during the process of sono-coating. The pattern of the as-synthetized nHAP samples before and after ultrasonic coating reveals only hydroxyapatite peaks. The ultrasonic treatment at 30oC for 15 min in water does not cause phase structure changes during nanoparticles deposition

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on substrates. The higher crystallinity was observed for nHAP6 in comparison with nHAP2 due to

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longer microwave synthesis, as described by Kuśnieruk et al. [65]. Samples were compared with data from JCPDS-International Centre for Diffraction Data file for crystalline hydroxyapatite

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(JCPDS, 9-0432).

Fig. 10 X-ray diffraction patterns of the nHAP6 and nHAP2 material before and after coating process.

Due to elastic nature of the polymeric scaffold it was important to know, whether the coating layer will remain stable during simulated operation manipulations similar to these expected at implantation procedure. For several nHAP-coated PCL fibers uniaxial stretching was performed.

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The fibers were also cut with a surgical scalpel. The layer of nanoparticles was not delaminating, and satisfactory adherence of the coating was proved (Fig.11). This is contrary to the situation of thick and brittle micron-size SBF-obtained layers, which tend to delaminate under loading and

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deformation [51,52].

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Fig.11. SEM images of scalpel blade cut on the single PCL fiber coated with nHAP, magnifications: A) 500x, B) 1000x, C) 5000x, D) 25 000x.

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3.3. In vitro cell culture results Cell culture cell proliferation tests after 5 days of revealed higher cell counts on polymer scaffolds coated with nHAP nanoparticles compared to uncoated materials (Fig. 12). The number of cells was up to 3 times higher on nHAP-coated PCL than on uncoated PCL. On the other hand, we observed excellent cell proliferation on uncoated β-TCP scaffolds, and a significantly lower

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cell proliferation rate on the nHAP-coated β-TCP scaffolds. The coated β-TCP samples and the

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coated PCL samples showed a similar cell count.

Moreover, SEM imaging revealed that cells spreading was more pronounced on the nHAP

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coated samples (Fig.13) than on the non-coated one. For the coated samples cells filopodial structures identified as nanopodia, were spreading in lateral directions (Fig.14). This can be

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attributed to the strong interactions of the filopodia with the nHAP coated nanostructured surface as well as surface increased hydrophilicity. The water contact angles were for uncoated PCL was

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65.3o ± 0.75 o and for nHAP- coated PCL it was 33.6o ± 7.13 o.

Fig. 12. Proliferation of MG-63 cells after 5 days on the following scaffolds: A) PCL, nHAP6-coated PCL, nHAP2-coated PCL, B) β-TCP, nHAP6-coated β-TCP, and nHAP2-coated β-TCP, n=5)

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Fig.13. Colorized representative SEM images of MG-63 cells on: A) PCL sample (magn. 1000x), B) PCL nHAP-coated sample (magn. 1000x), C) PCL sample (magn. 5000x), D) PCL nHAP-coated sample (magn. 5000x); optical camera representative images of water contact angle measurements: E) PCL sample, F) PCL nHAP-coated sample.

Fig.14. Colorized SEM image of MG-63 cell filopodia (arrows), on nHAP coated surface (magn. 50 000x).

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3.4. In vivo tests

Samples after preparation step were implanted in rabbit tibia bone as described in 2.4 In

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vivo tests section (Fig.15 A-D).

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Fig. 15. Images of a: A) trepan drill with raised periosteum tissue, B) created defect, C) PCL scaffold inserted in cavity D) β-TCP scaffold inserted in cavity,

Morphometric analyses revealed considerable new bone tissue formation after 3 months

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in all cases, except with the uncoated PCL scaffold (Fig. 16A–D). Bone formation and ingrowth

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into the scaffold were most pronounced in the nHAP-coated β-TCP samples, followed by the uncoated β-TCP samples and the nHAP-coated PCL samples. In the PCL scaffold samples,

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degradation was enhanced for the sonocoated samples (Fig. 16B).

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Fig. 16. Cross sections of A) PCL, B) nHAP-coated PCL, C) uncoated β-TCP scaffold, and D) nHAP-coated

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β-TCP scaffold. S, scaffold material; NB, new bone; *, soft tissue (Hem -Eos. staining, magnification x100).

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E) and F) µ-CT scans of nHAP-coated PCL scaffold in rabbit tibia bone after 3 months from implantation

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(scale bar 2mm).

Figure 17 graphically represents the results of quantitative morphometric analysis. Bone tissue

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comprised 19.0% ± 6.2% of the of the cross-sectional area of the uncoated β-TCP, 26.7% ± 3.7%

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for nHAP-coated β-TCP, 0.73% ± 0.1% for uncoated PCL scaffold and 10.9% ± 3.6% of the nHAPcoated PCL. The amount of bone did not significantly differ between the scaffold edges and the scaffold interior.

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A A

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B

Fig. 17. Diagrams of morphometric analysis: A) The volume % of scaffold material, new bone tissue, and soft tissue according to scaffold type. B) The volume % of the new bone tissue and soft tissues in the pores of the scaffolds according to scaffold type.

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In terms of the amount of new bone tissue, the best results were obtained with the nHAP-coated -TCP scaffold, where 68% of the volume of scaffold pores was filed new bone. The second-best results were obtained with uncoated -TCP, with 50% bone tissue and 50% soft tissue. With the nHAP-coated 3D-printed PCL scaffold, 33% of the volume of the pores was filled with new bone

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tissue.

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3.5. Immunohistochemical analysis results

Table 1 summarizes the results of immunohistochemical analysis. The upper row presents

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the results for new bone tissue adjacent to the implant material, while the lower row presents the results for control samples, i.e., natural bone adjacent to implanted materials. We found that OC

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was at a constant level, while OP values were variable.

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We used Wilcoxon test to estimate the statistical significance of the differences between the control samples and new bone tissue. For rabbit 6, with an uncoated 3D-printed PCL scaffold,

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the OPG and BMP2/4 levels in new formed bone were lower than in the control (P = 0.03). For rabbit 7, with an nHAP2-coated 3D-printed PCL scaffold, the OP and BMP2/4 levels in new formed

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bone were lower than in the control (P = 0.03). For rabbit 8, with an nHAP6-coated 3D-printed

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PCL scaffold, the OP and BMP2/4 levels in new formed bone were lower than for the control (P = 0.03). For rabbit 9, with an uncoated β-TCP implant, the OPG and MMP2 levels in new formed

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bone were lower than for the control group (P = 0.03). Spearman’s correlation revealed a moderate correlation between IL-1 and IL-10 (rs = 0.68; P = 0.04). The results of immunohistochemical investigations were in agreement with the results of morphometric investigations. All scaffold materials promoted cellular activity, except uncoated PCL. After implantation of a β-TCP scaffold, pro-inflammatory cytokine detection was prevalent, confirming the persisting (likely latent) presence of inflammation in newly formed bone. The bone

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regeneration process was most effective with the nHAP-coated ceramic β-TCP scaffolds (Fig. 17A-B), followed by uncoated β-TCP scaffolds, and PCL scaffolds coated with nHAP. Table 1. Immunohistochemical data regarding the relative appearance of factors in the bone tissue after implantation of different scaffolds in experimental rabbits. OC

OP

OPG

BMP2/4

MMP2

6

PCL Control

+++ +++

− −

−↓ +++

−↓ ++

++ ++

7

PCL+nHAP2 Control

+++ +++

+↓ ++

+++ +++

+↓ ++

8

PCL+nHAP6 Control

+++ +++

+↓ ++

+++ +++

+↓ +++

9

β-TCP Control

++ ++

0/+ +

++↓ +++

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β-TCP+nHAP2 Control

+++ +++

++ ++

++/+++ +++

5

β-TCP+nHAP6 Control

+++ +++

+/-

IL1

IL10

+++ +++

+/++ ++

++ ++

+ +/++

+ +/++

+++ +++

+++ +++

++ ++

+++ +++

++ ++

++ ++

+↓ ++

+/++ ++

++ ++

+ +

++ ++/+++

+ ++

+++ +++

+++ +++

+++ +++

+++ +++

++/+++ ++/+++

++/+++ ++/+++

++ ++

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TIMP2

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Material

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Animal code

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3. Discussion

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Significance of symbols in Table: −, No immunoreactive structures detected in visual field. +−, Few factor positive structures detected in visual field. ++, Moderate number of structures detected in visual field. +++, Abundance of structures detected in visual field. ↓, statistically significant difference. Abbreviations: OC – osteocalcin; OP – osteopontin; OPG – osteoprotegerin; BMP - bone morphogenetic protein; MMP2 - matrix metalloproteinase 2; TIMP2 - tissue inhibitor of matrix metalloproteinase; IL1 - interleukin and IL-10 interleukin 10.

In the present study, we addressed the issue of efficiency of a new sonocoating technology for coating of bone regrowth scaffolds at the temperature of 30oC and time of 15 minutes in aqueous solution. The scaffolds could be used to enhance bone tissue regeneration in bone defects caused for instance by tumor tissue extraction or bone fracture. Small defects might be repaired using bone cements that can both fill the small gaps and stimulate bone regrowth [5-7]. However, bigger size defects require repair using bone regrowth scaffolds that

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provide a structure to which cells can attach and multiply. The scaffold surface structure and properties are critical determinants of success in repairing large bone gaps. Our approach was to create a bone regrowth stimulating layer composed of nHAP particles with size in the range 10 to 43 nm using the sonocoating technique. Sonocoating for 15 min at the temperature of 30oC permitted to create a uniform layer with nano-size roughness and thickness in the range 200 to

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osteoconductive properties of porous bone regrowth scaffolds.

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300 nm. We examined whether this was an efficient method for enhancing the osteoinductive or

Previous studies demonstrated that ultrasonic energy can be used to coat various

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surfaces with nanoparticles [53,55,75]. This method has been described as the “throwing stones” technique since nanoparticles are thrown towards the material surface by a stream of liquid when

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cavitation bubbles collapse. Our present study is the first reported use of this technique to create

greater than the nanoparticle diameter.

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a uniform layer of extremely small nanoparticles, with the layer thickness an order of magnitude

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Compared to micro-HAP particles, nHAP particles have the potential advantage of a high

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specific surface area, and specific surface nanostructure. The AFM tests have shown that the nHAP coating is possess a unique nanotopography, with roughness in range of 20 – 30 nm. Given

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that the size of surface structures is comparable to those of proteins and cell organelles, enhanced interactions between the surface and corresponding biological structures are expected to occur.

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Further, it was previously reported that increasing stiffness of most of biomaterials in the range from MPa to GPa cause cells and especially cells filopodia to growth and spread. Therefore, nanostructure of surface is an important factor affecting cell adhesion and spreading. We observed enhanced cells spreading on PCL pellets coated with a HAP layer (Fig 12,13,14), which can be attributed both to enhanced hydrophilicity of such surfaces (Fig. 13), stiffness, and their nano-roughness (Fig. 8,9). This suggests that a nHAP material as a coating may provide a

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favorable environment for cells proliferation and migration, in agreement with studies on bulk nHAP pellets [66]. The challenge faced in the present study was how to use such small nanoparticles to create at temperatures close to room temperature conditions and in a short time a suitable coating on the porous scaffolds. Specifically, an effective coating must be a homogeneous layer, must

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uniformly cover the surfaces of the scaffold, do not delaminate during handling by a surgeon, and

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must be sufficiently stable to induce bone growth mechanisms before dissolving in the physiological fluids. The layer was created using the sonocoating technology. The results of this

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process could not be predicted due to the unique structure of the layers and nanoparticles [65,75]. In particular, one unknown factor was whether such a thin layer, with a thickness in the range of

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200 – 300 nm, would be either dissolved or metabolized by the organism before commencement

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of the bone regrowth process.

The sonocoating experiments in the present study resulted in the formation of a uniform

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nHAP layer both inside of the porous -TCP scaffolds and on the surface of the 3D-printed PCL scaffolds. The process parameters were precisely tuned to ensure that the surfaces were not

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melted or deformed during the cavitation process. The created layer was 5 to 10 times thicker

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than the nanoparticle size itself. Since the process was conducted at 30°C, it is unlikely that a chemical reaction occurred between the nHAP particles and the substrate. Such small

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nanoparticles can adhere to various substrates and to each other due to van der Waals forces. The produced sonocoating layer adhered to the PCL fibers even when they were stretched and cut (Fig.11). Advantages of the presently used process include the short coating time of 15 minutes, the low coating temperature of 30°C, and the use of water as a solvent. Additionally, the process required a very low concentration of nanoparticles (0.1 wt%) in the suspension. The coating parameters can be easily modified, including the coating time and temperature, the ultrasonic device parameters, and the distance of the samples from the sonotrode. Thus, the

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presented technology is “green”, and is scalable to mass production. The process is described in detail in the patent [75]. Further studies are needed to elucidate the mechanism of such layer creation, which is beyond the scope of the present paper. Our results indicated that sonocoating the scaffolds with nHAP is a highly efficient means

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of stimulating new bone tissue formation. Morphometric analysis of samples after 3 months

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implantation revealed positive results with both the -TCP porous ceramic scaffold and the nHAP-

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coated 3D-printed PCL scaffold. The fraction of new bone tissue in the space within the scaffold’s pores was significantly higher when using the coated -TCP scaffold (68%) compared to the

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uncoated -TCP scaffold (52%). Uncoated -TCP ceramics are already good bone-growth promoting materials and coating these scaffolds with active nHAP particles significantly enhanced

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bone tissue growth.

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Notably, both the coated and uncoated -TCP scaffolds performed better than the coated

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3D-printed PCL scaffolds in terms of new bone tissue production. The coated -TCP scaffold showed an approximately 3-fold higher efficiency compared to the 3D-printed scaffold. However,

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the 3D-printed scaffold has the advantage of being easily printed to any shape needed with

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engineered pore distribution [31,50,51,62-64]. The results of the in vitro and in vivo tests with the 3D-coated PCL scaffold were in

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agreement. Both indicate that nHAP coating increased cell proliferation and bone regrowth. On the other hand, in vitro tests with the -TCP scaffold showed that coating decreased cell proliferation to a level similar to with the coated PCL scaffold, while in vivo tests showed enhanced bone regrowth with the coated ceramic scaffold. It is seen on the -TCP example that in vitro tests are not always predicting the results of in vivo tests, where the results were opposite. The results of immunohistochemical analysis provided additional information about the quality of the newly formed bone tissue. From the bone ground substance proteins osteocalcin

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and osteopontin, osteocalcin is less dependent on sample type. Osteocalcin detection reflects bone tissue growth and mineralization [78]. Similar osteocalcin levels between the control samples and new bone tissue indicate that the mineralization of newly formed bone was similar to that in the control samples (Table 1). On the other hand, osteopontin acts as a hormone, and reflects the bone remodeling process that occurs in tissue [79,80]. We detected higher OP

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expression in the nHAP-coated samples compared to uncoated scaffolds. The variable

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appearance of OP suggests incomplete bone homeostasis in examined samples. These findings indicate that OP may be a prospective prognostic marker for bone ground substance quality in

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implantations.

All of the presently tested materials, except uncoated PCL scaffolds, demonstrated cellular

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activity in vivo. The lower MMP levels in the new bone tissue in samples with -TCP scaffolds

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indicated the occurrence of active extracellular matrix degradation and remodeling processes in the new bone tissue [81]. This result is in agreement with the higher bone growth activity observed

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for -TCP scaffolds than for PCL scaffolds. Active extracellular matrix degradation was also

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associated with high inhibitory TIMP levels in practically all cases. The nHAP-coated PCL scaffold resulted in better suppression of osteoclastogenesis than the uncoated scaffold, as evidenced by

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the higher osteoprotegerin (OPG) level. The major regulator of bone regeneration and growth, BMP, was expressed at the highest

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levels in cases with nHAP-coated and uncoated β-TCP scaffolds. This was in agreement with the finding that the fractions of new bone formation were highest with these scaffolds. As expected for a good bone regrowth scaffold, β-TCP scaffolds were associated with lower levels of tissue degradation enzymes (MMPs) and higher levels of tissue degradation inhibitors (TIMPs) in practically all cases, indicating active extracellular matrix degradation and remodeling processes. The dominant or equal expression of IL-10 relative to IL-1 further indicated the correct balance of

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immune response from the tissue side, with suppression of the main pro-inflammatory cytokine IL-1 in all cases. Comparing the 3D-printed PCL coated and uncoated scaffolds, we found a striking difference of tissue responses between them. For the uncoated PCL scaffold, only insignificant

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amount of bone tissue, occasionally on the surface of the polymer fibers was found (Fig. 16A),

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and the measured fraction of bone tissue volume was 0.7% only. Contrary, with the coated PCL

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scaffold, new bone tissue occupied 33% of the available space in the pores between the fibers (Fig. 16B). At the same time, degradation of the PCL fibers was more pronounced for the case of

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nHAP-coated samples.

It seems that bone cells are able to attach to the PCL fibers and create a thin bone layer.

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layer was a determinative condition.

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In the case of PCL scaffolds the nHAP coating was essential for stimulation of bone formation. In the case of the nHAP coated -TCP scaffolds, the nano-coating enhances

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performance comparing to the relatively good -TCP scaffold material, resulting in enhancement

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of bone regrowth.

The results of our in vivo experiments indicated that the best implantation material for bone

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regrowth was nHAP-coated β-TCP, followed by uncoated TCP ceramics, and then by nHAPcoated PCL scaffold. Our findings support the conclusion that a sonocoated layer of hydroxyapatite nanoparticles on 3D-printed PCL and porous β-TCP scaffolds remarkably increased the volume of newly formed bone, especially in the case of 3D-printed PCL implantation. Moreover, the quality of newly formed bone was comparable to that of the control bone samples, based on the similar expressions of bone ground substance proteins. Further studies are required to examine how nHAP particle size influences the osteoinduction or osteoconduction properties of scaffold coatings.

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As far as the plausible reasons for enhanced bone regrowth caused by the layer of nHAP particles produced by means of sonocoating, one may either consider enhanced proliferation of cells on the nanostructured surface or increased calcium ions concentration caused by a slow dissolution of the nHAP particles.

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As far as nanostructure effects are concerned, the observed surface roughness in the

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range of 20 – 30 nm (Fig 13,14,15) is comparable to the size of proteins and cell organelles.

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Therefore, interactions between the surface and corresponding biological structures are expected to occur. Surface topography is a factor affecting dynamics and other processes such as cell

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adhesion and spreading [82,83]. Also, it was previously reported that increased stiffness of biomaterials causes cells and especially cells filopodia to growth and spread [84]. AFM studies

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have shown an increased Young modulus for the coated surfaces comparing to the uncoated (Fig

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8,9). Fig 13 and 14 show that indeed the nHAP layer produced using sonocoating method enhances cells spreading over the surface of the polymer scaffold. This suggests that a nHAP

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coating may provide a favorable environment for cells proliferation and migration. Similar effects were demonstrated previously for pellets made from nHAP by pressing [66].

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On the other hand, in [65-66,85] it was shown that for nano-sized HAP there is a size

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dependent solubility and calcium ion concentration in test liquids. A reduced particle size leads to increased calcium ions level and may stimulate new bone formation into the scaffold’s pores [86-

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88]. The nHAP particles, contrary to the micron sized ones, are presumably gradually dissolved leading to a constant calcium release in the scaffold area, stimulating bone regrowth. New tissue was not progressively growing from the surrounding bone into the scaffolds, since we detected no statistically significant difference in the amount of new bone between the scaffold edges and interior. This could be explained by both rapid proliferation of cells from surrounding bone on the scaffold surface towards the interior (osteoconductivity), or the formation tissue in the scaffold interior (osteoinductivity). In each case, angiogenesis must have occurred

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as a consequence of new bone tissue formation. Future in vivo studies should be performed in different animal models to evaluate the species-specific osteoinductivity of the proposed materials, as well as their potential for healing critical-size bone defects [89,90]. The presently described ultrasonic coating method was not only stimulating substantial bone

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tissue growth on both types of tested scaffolds, but has also a substantial development potential,

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for low-cost industrial production of efficient bone scaffolds.

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since the process takes place in 15 minutes and at 30oC only in aqueous solution. It is suitable

4. Conclusions

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We demonstrated the successful use of sonocoating technology to coat 3D - printed PCL and porous β-TCP bone regrowth scaffolds with a layer of hydroxyapatite nanoparticles, which

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strongly enhanced new bone growth. nHAP nanoparticles were deposited on the scaffold surface

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creating a 200 to 300 nm thick layer with a unique nanotopography and surface roughness in the range of 20 to 30 nm.

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In vitro experiments revealed strongly increased osteoblast proliferation for nHAP-coated

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compared to uncoated 3D-printed PCL scaffolds. In vivo experiments in a rabbit model showed that sonocoating scaffolds with

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hydroxyapatite nanoparticles promoted up to 68 vol.% of new bone tissue growth inside the scaffold pores. The highest amount of new bone tissue growth occurred for the coated β-TCP

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ceramic scaffold (68 vol.%), followed by the uncoated β-TCP ceramic scaffold (50 vol.%), nHAPcoated 3D-printed PCL scaffold (33 vol.%). For uncoated PCL scaffold a thin layer only of new bone on the surface of the fibers was detected. Immunohistochemical analysis provided favorable results regarding the expression of bone tissue regeneration markers. The bone growth enhancing effect caused by the nHAP layer deposited on both scaffolds was attributed to the unique nanotopography of the coating surface and calcium ion release effect, resulting from gradual dissolution of the nHAP layer.

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Disclosure The Institute of High Pressure Physics PAS is an official holder of

“Method of

manufacturing bone implant and bone implant” (Polish granted patent PL226891 (B1), International Patent Application PCTIB2016/052566), USA Patent Application US 2018 / 0311407

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(A1).

Acknowledgements

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This work was financed by the Eranet project MATERA ERA-NET ‘SONOSCA’ MATERA/BBM-2557 “Sonochemical technology for bioactive bone regeneration scaffold

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production”. The research studies were conducted using equipment funded by the project CePT (reference: POIG.02.02.00-14-024/08), financed by the European Regional Development Fund

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within the Operational Programme “Innovative Economy” for years 2007–2013. We would like to

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personally acknowledge Mr. Jan Mizeracki, Barbara Ostrowska, PhD for the technical and

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scientific support of this research study.

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Highlights • Ultrasonic cavitation enabled deposition of nanolayers on porous materials. • nHAP coating on β-TCP and PCL bone regrowth scaffolds was deposited. • Nanometric nHAP layer promoted new bone tissue ingrowth in animal model in vivo. • Sonocoating was proved to be highly promising method for clinical applications

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