Bioresorbable Implants using Selective Laser Melting

Using bioresorbable materials implants can be manufactured which dissolve in the human body and are replaced by natural bone structure. For large implants an interconnecting porous structure needs to be integrated in the implant for a good vascularisation. Using additive manufacturing technology these internal structures can be directly manufactured. The structure can be designed by consequent following the guidelines of the medical expert. This paper describes the development of Selective Laser Melting to process bioresorbable materials Poly(D,L-lactide) and Tricalciumphosphate. The properties of the parts concerning microstructure, mechanical and biological properties after processing are analyzed in laboratory and animal tests. Possible applications are demonstrated and include individual bone substitute implants in cranio-maxillofacial surgery. Introduction In medical science rapid prototyping (RP) and rapid manufacturing (RM) are manufacturing technologies with rapidly growing influence and large potential for several applications. Some applications are already industrial standard with growing numbers of applicants. These are for example the manufacturing of teeth restorations out of Cobalt-Chromium-alloy or the manufacturing of complex orthopedic devices. The manufacturing of bone substitute implants using RM technology is another field of interest. Additive Manufacturing like Selective Laser Melting (SLM) (also described as Direct Metal Laser Sintering or LaserCusing) or Electron Beam Melting (EBM) has advantages to conventional manufacturing like milling or casting concerning geometrical freedom, flexible production, material consumption and manufacturing time. The manufacturing with additive technology becomes economical in two cases. The first case is the manufacturing of individual implants in small series. There is no need for the casting or deep drawing. Individual implants based on CAD-data can be directly manufactured without changes in machine set up. Secondly manufacturing becomes economical when new functionalities can be integrated in the implant which would not be possible by conventional manufacturing. When these functionalities result in better quality of life of the patient, higher production costs can be negotiated be better durability and functionality of the implant. These functionalities can be for example controlled porosity of the implant, adapted mechanical properties or new composite materials. There have been many investigations on the manufacturing of permanent medical implants out of titanium(alloys), cobalt-chromium-alloys or implant steel [1, 2]. There were already clinical studies and applications of implants in the human body which were manufactured by Rapid Manufacturing technologies [3]. In many cases regenerative methods would be much better suited for the treatment of bone defects. There is taking place a strategical change in implantology from permanent implants to regenerative methods using resorbable materials. The gold standard at the moment is human bone taken e.g. from the iliac crest. This method requires a second surgery with its secondary morbidity. To avoid this, bioresorbable bone substitute materials are developed. There are several products on the market using different materials and different processing technologies. The materials can be used by the surgeon as paste, granules or semi-finished parts like cuboids or wedges. At the moment there is no processing technique to manufacture individual implants out of bioresorbable materials. This is based on two reasons. Conventional manufacturing methods involve a sintering process. This requires a tool to shape the implant. Individual implants would therefore be expensive due to expensive tool production. The second reason is the internal structure of bioresorbable implants. For full vascularisation of large implants an interconnected porous structure needs to be integrated in the implant. By varying the pore size ingrowth of cell tissue into the implant is possible. To realise a full vascularisation the pore size and interconnectivity needs to be well defined. These two problems for manufacturing of individual bioresorbable implants can be solved by rapid manufacturing technologies. The geometrical freedom of additive manufacturing makes the production of individual geometries and the realisation of a defined interconnected porosity possible by integrating the porous structure in the design of the implant. There are several bioresorbable materials for example polymer-based [4], bioglass [5], resorbable metals or bioceramics [6, 7]. Bioceramics based on calcium phosphate (e.g. -TCP) have the highest acceptance in medical science since their chemical structure is closest to human bone. In this work the additive manufacturing with Selective Laser Melting (SLM) is applied. For fusion of the powder a melt phase is required. Materials with thermal instability (like -TCP) cannot be processed with SLM directly. Materials suitable for processing using melting of the material are polymer based materials e.g. polylactide (thermoplastic). During the resorption of polylactides an acidic environment is generated next to the implant due to the lactidic structure of the resorbable polymer. An acidic environment in human tissue results in inflammatory reactions [4]. For optimised properties of manufactured implants a combination of both materials is chosen: polylactide for processing using SLM and -TCP for optimised biological properties. The approach of combining two materials has other medical advantages. The acidic surrounding can be decreased by using a composite with alcaline materials like calcium phosphate ceramics. Using a composite of polymer and ceramic material the mechanical properties of the parts can be enhanced to pure polymer [8]. The main challenges for manufacturing bioresorbable implants using SLM are the preparation of the used materials for the SLM process and the processing of the material without damaging or changing its structure and chemical composition. This work deals with the process adaption of Selective Laser Melting to process the bioresorbable composite poly(D,L-lactide) and -tricalcium phosphate (PDLLA/ -TCP). This is a completely new approach using medical approved materials to generate parts by full melting of PDLLA and embedding of -TCP in a matrix of PDLLA. Materials and Methods Reflection of radiation The processing of the used material with SLM strongly depends on the optical properties of the material. The laser source (wavelength) is then used depending on the measured reflection spectrum. As the material was used as a powder, the reflection of radiation of different wavelengths is analyzed on powder material. For wavelengths ranging from 400-2600 nm an UV/VIS/NIR Spectrometer Lambda 9 from Perkin Elmer is used. For wavelengths from 1900-22000 nm a FTIR Spectrometer 1725 X from Perkin Elmer is used. Powder morphology -TCP powder (Biovision GmbH Biomaterial, Wiesbaden, Germany) and PDLLA (Boehringer Ingelheim, Ingelheim, Germany) was used to produce the composite powder. Both materials were filled in a polyethylene bottle. Zirconia milling balls were added and the bottle was put on a rolling platform for three weeks. The resulting powder was prepared with a sieve of 90 μm mesh. The prepared composite powder was characterized before and after the milling process. The particles size distribution was measured using a particle size analyzer (Mastersizer 2000, Malvern, Worcestershire, Great Britain). The milled particles were embedded in a resin and a cross-section was prepared and analyzed using scanning electron microscopy (LEO 440i, Carl Zeiss SMT AG, Oberkochen, Germany). SLM processing For the processing of the composite material PDLLA and -TCP using SLM, a process chamber developed at ILT was used together with a CO2-laser (FEHA 400S, FEHA, Germany). The laser beam was projected on the powder bed using a scanning device (hurryscan 20, Scanlab, Germany) and an F-Theta focusing lens. The CO2-laser operates in cw-mode at a wavelength of 10600 nm. To adjust the laser power for low intensities (0.2-10 W) a power attenuator (ULO optics, United Kingdom) was used. A schematic of the experimental set up is shown in Fig 1. To achieve a reproducible deposition of powder layers the powder deposition device was adopted. For qualifying the SLM process for the processing of the composite material an adaption of the process parameters laser power, laser beam diameter, scanning velocity, track distance as well as powder layer thickness was done. Fig 1: Schematic of the experimental setup Microstructural analysis One problem of powder based additive manufacturing is the appearance of pores and defects in the microstructure of the generated parts. This results in poor mechanical properties. Therefore the first aim for qualifying SLM to process the composite material PDLLA/ -TCP is to produce dense parts with a relative density >98%. Analysis of the density is done by optical analysis of cross sections of parts generated with different process parameters. An optical microscope and an analyzing software (Analysis, Olympus, Hamburg, Germany) is used to perform the measurement of the density. Mechanical properties The mechanical properties of bioresorbable bone substitute implants are tested with compression tests on porous test samples. This is in correlation to the literature [9, 10]. Cubes with varying pore diameter (500, 600 and 800 μm) and resulting porosity are manufactured using SLM and compressed until fracture of the part. Biological analysis The implementation of new manufacturing technologies for the processing of medical approved materials for series production of implants requires biological and medical analysis of manufactured parts. In biological tests (in vitro tests) the effect of the generated parts on bone cells (osteoblasts) is analyzed. Bioresorbable test geometries will be manufactured with an interconnecting pore structure to induce the ingrowth of bone into the implant for full resorption of the part. The pore diameter shows significant influence on the growth behavior of the cells. Test parts with different pore diameters are manufactured and the proliferation of the cells on the parts is measured using life / dead staining of the cells and optical analysis. To minimize the geometrical influence of the parts on the results, the same test geometry is used for in vitro and in vivo tests (animal tests). The geometry is a cylinder of 15.6 mm diameter and 5 mm thickness. Pore structure is designed using CAD software according to Fig 2 using multiplied unit cells. The pore diameter is varied between 500-800 μm. Fig 2: Schematic of the design of porous test geometries using multiple unit cells. The test geometries are manufactured using process parameters for manufacturing of dense parts without defects to achieve high strength of the parts. The powder remaining in the pore structure is then removed by blasting with alumina particals. The manufactured and blasted test parts are then cleaned in ethanol, -sterilized and populated with human osteoblasts (HOB). The analysis of the cultured cells takes place after 1, 3 and 6 weeks. The results are compared to reference material (poylvinylalcohol, PVA) and phase pure medical grade -TCP. The cells are then counted using optical analysis. Animal experimental studies Results concerning degradation and bone ingrowth into porous implants can only be achieved by suitable models of animal studies. To evaluate the properties of SLM implants with interconnected porous structure in vivo 16 implants as described in the previous section with a pore channel diameter of 600 μm are manufactured. The parts are then sand blasted, washed in ethanol and -sterilized as described earlier. The animals are 20 rabbits of the race Chinchilla. The bone defect is located in the skull (calvarium) of the animals as shown in Fig 3. Four defects are filled with autologous bone (control), 16 defects are provided with SLM implants. Fig 3: Schematic of the bone defect and the skin incision Histological analysis is performed on cross sections of the test area after 8 weeks.