3D printed bone implants for the treatment of infant tumors

• 1
  Why infant tumor reconstruction is not adult reconstruction
• 2
  How 3D printed bone implants are actually made
• 3
  Where CNC machining finishes what printing starts
• 4
  Material choice and its consequences
• 5
  Where 3D printed bone implants are the wrong answer
• 6
  Which route fits which feature

Osteosarcoma is the most common primary bone cancer, and roughly half of all cases appear in children and adolescents. The surgical plan is usually resection of the affected bone segment followed by reconstruction, either with a tumor prosthesis or with a patient-specific implant. In an adult, a fixed-size revision system can often be adapted on the table. In a young child, it usually cannot.

The reason is growth. A five-year-old femur may add several centimeters of length before skeletal maturity, and a standard implant sized for the current defect will not survive that change. The defect is also small, curved and close to open growth plates, so a bulky off-the-shelf component intrudes on structures that still have to function for another decade.

A third factor is load. Children heal fast, but their bone is softer and more compliant than adult cortical bone, and the implant shares load with a skeleton that is still mineralizing. Stress shielding matters more here than in a 60-year-old hip. Design margins that look conservative on paper can still loosen a screw in a growing pelvis.

That combination is what pushed surgeons toward patient-specific devices, and toward 3D printing as the shaping route. The geometry is free-form, the quantity is one, and the delivery window is measured in days. Those are exactly the conditions under which additive manufacturing beats casting and forging.

Almost all metallic patient-specific bone implants start from a CT scan, usually at 0.5–1.0 mm slice thickness. The DICOM stack is segmented into bone, then converted to a watertight STL or STEP surface. This step is where most error enters the process. Thresholding too aggressively smooths cortical bone and shrinks the model by 0.3–0.8 mm; thresholding too loosely leaves soft tissue attached and inflates it.

From the segmented model, the engineer mirrors the healthy contralateral side to reconstruct the resected volume, then designs the fixation plate, screw trajectories and any porous lattice in CAD. Porous regions are typically 300–800 μm pore size with 60–80% porosity, which sits in the range where bone ingrowth is reported and where the structure can still be printed without collapsing.

The build itself is usually laser powder bed fusion on Ti-6Al-4V (TC4) powder, 20–45 μm particle size, layer thickness 30–60 μm, in an argon atmosphere with oxygen held below 1,000 ppm. Solid regions are melted at high energy density for near-full density; lattice regions use lower energy density to keep struts intact.

After the build, the part is stress-relieved, cut from the plate, and the support structures are removed by hand and with fine tooling. Surfaces that interface with bone are often grit-blasted or acid-etched; surfaces that interface with soft tissue are polished. Anything that threads, seats or keys into another component is machined, and that is where a CNC partner enters the chain.

Printed titanium is not a finished interface. As-built laser powder bed fusion typically holds ±0.1 mm on external features, which is ten to twenty times looser than what a screw hole or a taper needs. The printed blank is therefore treated as a near-net shape, and the critical features are cut on a 5-axis machining center.

In practice we machine the following on pediatric bone implants: screw holes and their countersinks, taper and Morse connections to stems or extendable mechanisms, bone-contacting seating faces, and any flat datum used to align the implant during surgery. Threads below M2 and holes below Ø1.5 mm are the usual trouble spots, because printed material near the surface can be partially unmelted and will tear rather than cut.

Titanium is also a poor conductor. Heat concentrates at the cutting edge, and a worn tool will smear rather than shear. We run TC4 at 40–70 m/min surface speed with high-pressure coolant, carbide tools with sharp edges and generous rake, and we change inserts on a count rather than waiting for a bad finish. Deep holes get peck cycles.

Holding tolerance matters here because a screw hole that is off by 0.1 mm can miss the planned bone corridor. Our general machining tolerance is ±0.005 mm on turned and milled features, with surface finish down to Ra 0.2–0.8 μm where a polished interface is required. Those numbers apply to the machined features, not to the printed lattice.

Ti-6Al-4V (TC4) is the default for load-bearing pediatric implants. It has a proven clinical history, prints well, and machines to a predictable finish. Its elastic modulus is still roughly five to ten times that of cortical bone, which is why lattice regions are used to bring the effective stiffness down rather than changing alloy.

Commercially pure titanium TA1 and TA2 appear where strength is not the driver: cranial plates, small facial reconstruction, and any part that must be bent or contoured in the operating room. They machine more easily than TC4 and hold a cleaner edge, but they do not carry a femoral load.

Cobalt-chrome and stainless 316L are used in some trauma and spinal hardware. 316L prints and machines well, is cheaper, and is a reasonable choice for non-permanent fixation. Its corrosion behavior in long-term implantation is weaker than titanium, so we would not propose it for a device intended to stay in a growing child for ten years.

PEEK and carbon fiber reinforced PEEK are used for radiolucent components, so the surgeon can image through the implant during follow-up. PEEK machines cleanly with sharp tooling and air blast, but it does not osseointegrate. It is a spacer and a structural element, not a bone-bonding surface.

A printed implant is a single-use device with a validated process behind it. If the defect is a simple mid-shaft segment with a standard diameter, an off-the-shelf intramedullary nail or plate will be cheaper, faster and has years of published survival data. Printing adds cost and process risk without adding clinical value in that case.

Emergency trauma is another mismatch. Segmentation, design review, printing and machining of a patient-specific device cannot be compressed into a few hours without skipping verification steps, and skipping them is how a screw corridor ends up in the wrong place. For acute fractures, standard hardware is the correct answer.

Very small features are a hard boundary. A lattice strut thinner than about 200 μm will not print reliably on most laser powder bed systems, and a screw hole smaller than about Ø1.2 mm is difficult to drill in titanium without breaking. If the anatomy demands features below those sizes, the design needs to change rather than the process.

Finally, growth remains a design problem, not a manufacturing one. Some pediatric devices are built as extendable constructs with a replaceable midsection, so the fixation in bone stays put while the length is adjusted in a later procedure. That decision belongs to the surgical team. Our job is to hold the interface tolerances that make the extension mechanism work.