95gTc and 96gTc as alternatives to medical radioisotope 99mTc.

We studied 95gTc and 96gTc as alternatives to the medical radioisotope 99mTc. 96gTc (95gTc) can be produced by (p, n) reactions on an enriched 96Mo (95Mo) target with a proton beam provided by a compact accelerator such as a medical cyclotron that generate radioisotopes for positron emission tomography (PET). The γ-rays are measured with an electron-tracking Compton camera (ETCC). We calculated the relative intensities of the γ-rays from 95gTc and 96gTc. The calculated γ-ray intensity of a 96gTc (95gTc) nucleus is as high as 63% (70%) of that of a 99mTc nucleus. We also calculated the patient radiation doses of 95gTc and 96gTc, which were larger than that of 99mTc by a factor of 2-3 based on the applied assumptions. A medical PET cyclotron which can provide proton beams with energies of 11-12 MeV and a current of 100 μA can produce 12 GBq (39 GBq) of 96gTc (95gTc) for operation time of 8 h, which can be used for 240 (200) diagnostic scans.

Introduction

Various radioisotopes, such as 99mTc (half-life 6.02 h), 201Tl (half-life 3.04 d), and 133Xe (half-life 5.27 d), are used for single-photon emission computed tomography (SPECT) in medical diagnostic scans. In particular, 99mTc has become the most important medical radioisotope at present [1]. Over 30 commonly used radiopharmaceuticals are based on 99mTc; in addition, new radiopharmaceuticals have been developed (for example, Refs. [2, 3]). The 99mTc radioisotopes are supplied by 99Mo/99mTc generators, which continuously generate 99mTc through the β-decay of the parent nucleus 99Mo accumulated inside the generators. This supply method provides two excellent advantages. First, it is possible to transport 99Mo/99mTc generators from a production facility to any place in the world because the half-life of 99Mo is as long as 2.75 d. Second, when a 99Mo/99mTc generator is transported to a hospital, 99mTc can be produced fresh for up to 2 weeks by daily milking/elution from this 99Mo/99mTc generator. At present, the parent nucleus 99Mo is produced in nuclear reactors by the neutron-induced fission of 235U in highly enriched uranium (HEU) targets, in which the fraction of 235U is approximately 90%. However, some nuclear reactors that have supplied 99Mo require major repairs or shutdown, which may lead to a 99mTc shortage. Thus, many alternative methods to produce 99Mo or 99mTc without HEU have been proposed [4], and they can be classified into four groups. The first group involves nuclear fission using low-enriched uranium (LEU) with a fraction of up to 20% in nuclear reactors. This technology has been established, and some reactors such as the OPAL reactor [5] have started the production of 99Mo. The second group involves nuclear fission of 238U (or natural uranium) using high-flux neutrons/photons provided by accelerators [6]. This method has two advantages: 99mTc can be provided in the form of 99Mo/99mTc generators, and this method does not require LEU or HEU. However, the development of high-flux neutron/photon sources is a technical challenge. The third group involves β-decay from 99Mo produced by nuclear reactions without fission on uranium, for example (n, γ) reactions on 98Mo in nuclear reactors [7], (γ, n) reactions on 100Mo using accelerators [6, 8, 9], or (n, 2n) reaction with high energy neutron beams [10, 11]. The fourth group involves a production process using the 100Mo(p, 2n)99mTc reaction with proton beams provided by a compact cyclotron with medium energies of 18–24 MeV [12, 13, 14, 15, 16, 17, 18, 19, 20, 21]. For the last two methods, the plan is for 99mTc to be directly transported from a production facility to hospitals. In such cases, although the half-life of 99mTc is 6 h, it is expected that a production facility can cover an area with a radius of 400 km [1].

The September 11th terrorist attacks in Washington D.C. in 2001 also affected medical radioisotope production from the viewpoint of the safeguards of nuclear materials. The control of fissionable nuclides such as 235U and 239Pu is important for the safeguards of nuclear materials [22]. The International Atomic Energy Agency (IAEA) hopes to discontinue 99mTc production using HEU targets, which can be transmuted into nuclear weapons [23]. In the near future, 99mTc will be supplied by nuclear reactors using LEU targets in addition to HEU. The Nuclear Energy Agency (NEA) reported the prediction that the 99Mo/99mTc supply will be larger than the world demand when the scheduled nuclear reactors using LEU start 99Mo production [24]. However, the 235U fraction of 20% in LEU is still too high from the viewpoint of the safeguards of nuclear materials. In fact, the governments of the U.S. and Iran agreed to convert Iran’s LEU to relatively low-enriched uranium with a fraction of approximately 3.7%, which is usually used for light water reactors. Therefore, the development of an alternative method to supply Tc radioisotopes without LEU or HEU is expected.

Because the Tc chemistry is the same, all the radiopharmaceuticals based on 99mTc can, in principle, be applied to other Tc isotopes. There are five Tc isotopes with half-lives in the range from hours to days: 94mTc (half-life 52 m), 94gTc (half-life 4.88 h), 95mTc (half-life 60 d), 95gTc (half-life 20 h), and 96gTc (half-life 4.28 d), as shown in Fig. 1. Historically, various Tc isotopes have been studied as medical radioisotopes. In 1976, the biological half-life of Tc was measured using 95mTc instead of 99mTc [25]. Autoradiography using 95mTc and 96gTc was studied in the 1970s [26], and recently, 94mTc has been studied as a positron emitter for PET [27, 28, 29]. The half-life of 96gTc (4.28 d) is long enough for worldwide delivery from a production facility and lengthy use of up to 2 weeks in hospitals. 95gTc (20 h) can also be transported to a wide area and used for 3–5 days in hospitals. Thus, 95gTc and 96gTc are candidates for alternative γ-ray emitters. However, the decay rates of 95gTc and 96gTc are lower than that of 99mTc by a factor of 3.3 and 17, respectively, because the decay rate of a radioisotope is inversely proportional to its half-life. This fact leads to the question of whether these isotopes can work as 99mTc medical radioisotopes.

Fig. 1

Partial nuclear chart around Tc isotopes. The solid arrows show the (p, n) reaction. The dashed lines show the β-decays and internal decays with half-lives shorter than 5 days. The dotted lines show the β-decays with half-lives longer than 105 years. The large solid arrows show the β-decays after the nuclear fissions in nuclear reactors.

In the current study, we present the relative γ-ray flux of these isotopes with simple assumptions. We also estimate the patient radiation does per Tc-labeled tracer using the PHITS simulation code [30]. Various nuclear reactions that are production methods of Tc isotopes, such as (p, n) reactions [31, 32, 33], deuteron-induced reactions [34], and 96Ru(n, p)96gTc reactions [35, 36], were studied. We consider the production by the (p, n) reaction on an enriched Mo isotope. We also calculate the production rate using a typical PET medical cyclotron. Because the energies of decay γ-rays of these Tc isotopes are typically higher than 200 keV, they are not suitable for the traditional SPECT cameras. Thus, we also discuss the property of possible ETCC for high energy γ-rays.

Materials and methods

Relative γ-ray intensity

The question that we should ask here is the γ-ray flux of a Tc isotope relative to that of a 99mTc nuclide at a detection position outside of a human body. The relative γ-ray flux of a 96gTc nucleus can be approximately calculated using the equationwhere λ is the decay rate of the isotope, R is the decay factor of the isotope during the time between the isotope production and the diagnostic scan, M is the multiplicity of the emitted γ-rays per decay, I is the ratio of the isomer (the ground state) to the summation of the isomer and the ground state at which the nuclides are synthesized in the case that the isomer (the ground state) is used for radiotracers, and P is the probability of the γ-ray penetrating a human body. The aim of the calculation using this simple Eq. (1) is to clearly present the calculation details.

The decay factor R is defined by R = Nt1/Nt0, where Nt1 and Nt0 are the numbers of nuclides at the time of a diagnostic scan and at the end of the radioisotope production (beam irradiation or milking), respectively. We assume 1 h as a typical duration for 99mTc because it is typically produced every morning from 99Mo/99mTc generators in hospitals, whereas we assume 5 h as a typical duration for the other Tc isotopes that should be transported from production facilities to hospitals.

The multiplicity is the number of γ-rays emitted from a nucleus via decay. As shown in Fig. 2, 96gTc predominantly decays to two excited states in the daughter nucleus 96Mo through electron capture, and subsequently these two states decay to the ground state by the emission of three cascades of γ-rays with energies of 812(1127)–850–778 keV. Because the energies of these γ-rays except for the 1127 keV γ-ray are in the narrow energy range of 770–850 keV, it is possible to distinguish the 96gTc γ-rays with a gate on the energy window of 770–850 keV. We thus take the multiplicity of M = 3 for 96gTc. The nuclear structure of 94Mo is similar to that of 96Mo, and the multiplicity of 94gTc is also M = 3. 95mTc finally de-excites to the ground state of 95Mo with the emission of various γ-rays following β-decay, and we take M = 2 as a typical value for 95mTc. Because 95gTc predominantly decays with the emission of a γ-rays of 766 keV, we take a multiplicity of M = 1 for 95gTc.

Fig. 2

Partial level schemes for 94Tc (left) and 96Tc (right). The solid arrows with numbers denote decay γ-rays and their energies in unit of keV. The dashed arrows show population from meta-stable states.

One point to consider in the calculation of the γ-ray flux is the ratio of the isomer (or the ground state) to the summation of the isomer and the ground state if the isomer (or the ground state) is used for radiotracers. The population ratio of the isomer to the summation for 94Tc is as high as I(94mTc) = 0.8–1.0 and the ratio for 95gTc is I(95gTc) = 0.25–0.3 [37]. Because the isomer of 96Tc continuously decays to its ground state with a half-life of 52 m during beam irradiation, the final ground ratio depends strongly on the time of the beam irradiation and transport. With assumption of the beam irradiation time of 8 h, the cooling time of 5 h, and the isomer production ratio of I(96mTc) = 0.8 in proton induced reactions, we obtain I(96gTc) = 0.998 and I(96mTc) = 0.002. In the case of the 99mTc production using the 100Mo(p, 2n)99mTc reaction, the isomer ratio is approximately 0.3 [38]. When 99mTc is produced from 99Mo/99mTc generators, the fraction of 99mTc is not high because it continuously decays to 99gTc inside the generators. The fraction of 99mTc can be calculated aswhere λ and λ are the decay rates of 99Mo and 99mTc, respectively, and t is the time interval between two milkings. With the assumption that 99mTc in saline solution is taken from a 99Mo/99mTc generator by milking once per day (t = 24 h), the ratio of 99mTc/(99mTc + 99gTc) is 0.38 using Eq. (2). We use an isomer ratio of 0.38 for 99mTc.

The penetration of γ-rays depends on the depth of the radioisotope and the γ-ray energy. To discuss the γ-ray flux with dependence on the depth, we consider the penetration probability through H2O with thicknesses of 3 cm and 10 cm.

Dose

To estimate the patient radiation doses from 95gTc and 96gTc, we calculated the deposited energy of a 95gTc (96gTc) decay using the PHITS particle transport simulation code [30]. The deposited energy depends strongly on the geometry, namely the size of the patient body and the position of the accumulated radiotracers. Here, we assume that the 95gTc (96gTc) radioisotopes are located at a 3 cm or 10 cm depth inside a 40-cm diameter sphere that is filled with H2O. The relative dose is calculated by multiplying the energy deposited in the sphere by the decay rate. The relative doses are obtained aswhere E is the deposited energy per decay.

Production rate using medical compact cyclotron

Compact negative ion cyclotrons for producing various radioisotopes for PET have been developed. These cyclotrons can provide proton beams with currents of up to 100 μA and energies of 12–18 MeV. Furthermore, the advent of cyclotrons with self-radiation shields has enabled the creation of radioisotopes without requiring heavily shielded rooms in hospitals. As a result, over 200 medical PET cyclotrons have already been constructed in Japan. Thus, we consider the (p, n) production using typical medical PET cyclotrons. The (p, n) nuclear reaction cross sections on Mo targets typically have a maximum value in the range of 11–13 MeV. Therefore, medical cyclotrons can effectively produce the radioisotopes 95gTc/96gTc. The nuclear reaction cross section depends on the proton beam energy, which decreases via atomic processes inside the targets. Thus, we calculate the production rate with the PHITS code by considering the beam energy loss inside the targets. We assume that a molybdate oxide target with a thickness of 1 g/cm2 is irradiated by proton beams with energies of 7–15 MeV.

Results and discussion

γ-Ray intensities of 95gTc and 96gTc

Table 1 provides a summary of the calculated relative γ-ray intensities and the parameters in the case of a 3-cm thick layer of H2O. The intensities of 96gTc and 95gTc relative to 99mTc are 0.63 and 0.7, respectively. Table 2 shows the relative intensities in the case of a 10 cm depth, in which the relative intensity of 96gTc (95gTc) is 1.0 (1.1). The decay rates of 95gTc and 96gTc are much lower than that of 99mTc. However, the multiplicity, the ground state (isomer) ratio, and the penetrability of 95gTc (96gTc) are higher than those of 99mTc. As a result, the γ-ray intensity per nucleus becomes almost same. The current results show that 96gTc and 95gTc can be used as alternatives to 99mTc from the viewpoint of the γ-ray flux.

Table 1

Relative γ-ray intensity outside of a body.

Isotope	T1/2	Decay rate [1/h]	Residual rate after 5 h	Mγ	m/(m + g) or g/(m + g)	Penetrability, 3 cm H2O	Relative intensity
99mTca	6 h	0.17	0.89c	1	0.38e	0.63	1.0
99mTcb	6 h	0.17	0.56	1	0.3	0.63	0.50
96mTc	51.5 m	1.2	0.018	1	0.002d	–f	–
96gTc	4.28 d	0.0098	0.97	3	0.998d	0.79	0.63
95mTc	61 d	6.8 × 10−4	1.0	2	0.25	0.79	7.6 × 10−4
95gTc	20 h	0.05	0.84	1	0.75	0.79	0.70
94mTc	52 m	1.2	0.018	1	0.8	0.80	0.38
94gTc	4.9 h	0.20	0.49	3	0.2	0.80	1.4

99mTc produced by fission in nuclear reactors.

99mTc produced by (p, 2n) reactions in accelerators.

Residual rate after 1 h instead of 5 h.

The ratios after the beam irradiation of 8 h and the cooling of 5 h.

See main text.

Internal conversion dominates for the M3 transition with energy of 34 keV.

Table 2

Relative γ-ray intensity outside of a body through 10 cm of water.

Isotope	Penetrability, 10 cm H2O	Relative intensity
99mTca	0.22	1.0
99mTcb	0.22	0.5
96gTc	0.46	1.0
95mTc	0.46	0.013
95gTc	0.45	1.1
94mTc	0.47	0.64
94gTc	0.47	2.3

99mTc produced by fission in nuclear reactors.

99mTc produced by (p, 2n) reactions in accelerators.

The γ-ray penetrability increases with increasing γ-ray energy in the energy region of E < 2 MeV. The penetrability of the 700–900 keV γ-rays from 94–96Tc through a 10-cm thick layer of H2O is higher than that of the 141-keV γ-rays from 99mTc by a factor of approximately 2 (see Table 2). As a result, the relative intensities of these Tc isotopes increase to 1.0–1.1. In this way, these high-energy γ-ray emitters can provide an advantage in the imaging of deep positions in a human body.

Table 3, Table 4 show the calculated doses using Eq. (3) and the deposited energies of 95gTc (96gTc) decay in the sphere filled with H2O. Because the decay rates of 96gTc and 95gTc are much lower than that of 99mTc, it is expected that their radiation doses are lower than 99mTc. However, the calculated doses of 95gTc and 96gTc relative to that of 99mTc are in the range of 2.4–3.0. This is because the total energy of γ-rays emitted from 95gTc and 96gTc are as high as 1.1 MeV and 2.5 MeV, respectively, and hence the deposited energy per decay of 95gTc (96gTc) is higher than that of 99mTc by a factor of 5–15. This result suggests that the patient radiation doses of 95gTc and 96gTc are higher than that of 99mTc by a factor of approximately 2–3.

Table 3

Relative radiation dose in the case of a 3-cm depth.

Isotope	Decay rate [1/h]	Residual rate after 5 h	m/(m + g) or g/(m + g)	Deposited energy [MeV]	Relative dose
99mTca	0.17	0.89b	0.38c	0.052	1.0
96gTc	0.0098	0.97	1	0.88	2.8
95gTc	0.05	0.84	0.75	0.28	3.0

99mTc produced by fission in nuclear reactors.

Residual rate after 1 h instead of 5 h.

See main text.

Table 4

Relative radiation dose in the case of a 10-cm depth.

Isotope	Deposited energy [MeV]	Relative dose
99mTca	0.081	1.0
96gTc	1.17	2.4
95gTc	0.38	2.6

99mTc produced by fission in nuclear reactors.

Recently, a system combining SPECT and ordinary computed tomography (CT) has been introduced [39]. The radiation dose originating from CT is as high as that of the radiotracer. For example, the radiation is 6.3–8.9 mSv in SPECT/CT scanning using 99mTc-labled tracers with a radioactivity of 740–1100 MBq, whereas the X-radiation exposure from CT scanning is estimated to be 3.8–15.1 mSv [40]. When 95gTc or 96gTc is used as the radiotracer of the SPECT/CT scan, its dose would be higher than that from CT.

Fig. 4, Fig. 5 show the relative reaction rates calculated with PHITS. In the energy region of E < 14 MeV, 95gTc and 95mTc are the dominant products of the p + 95Mo reaction (see Fig. 4). Above 14 MeV, the production yields of 94gTc and 94mTc, which are produced by (p, 2n) reactions, suddenly increase. The productions of 96mTc and 96gTc are dominant at energies of E < 12 MeV in the p + 96Mo reaction (see Fig. 5). In both reactions, the production rates of the niobium isotopes are lower than that of the dominant product by at least two orders of magnitude in the energy region of E ≤ 13 MeV. We conclude that energies of 11–12 MeV are suitable for 95gTc and 96gTc production.

Fig. 4

Calculated relative reaction rates of the p + 95Mo reaction. The thickness of the 95MoO3 target is 1 mg/cm2.

Fig. 5

Calculated relative reaction rates of the p + 96Mo reaction. The thickness of the 96MoO3 target is 1 mg/cm2.

During the beam irradiation, 96mTc with a half-life of 51.5 m continuously decays to 96gTc. As a result, the residual isomer ratio, m/(m + g), is approximately 13% at the end of the beam irradiation of 8 h. After an additional cooling time of 5 h, the isomer ratio decreases to only 0.2%, whereas the ground state ratio increases to 99.8%. Although the 0.2% fraction of 96mTc is approximately 20% in activity, due its internal decay it will not cause significant change in the dose to the patients.

For one diagnostic scan, 74–740 MBq of 99mTc-labeled tracers are injected into a patient. We take 370 MBq as the typical value, corresponding to 1 × 1013 atoms. By dividing by the isomer ratio of 99mTc/(99mTc + 99gTc) = 0.38 (see Table 1), the number of 2.6 × 1013 atoms is obtained as the total number of the ground state and the isomer. As discussed previously, the same number of 96gTc (95gTc) atoms can work as alternative to 99mTc from the viewpoint of the γ-ray intensity. By considering the isomer ratio, the radioactivity of the 2.6 × 1013 atoms of 96gTc (95gTc) is approximately 49 MBq (190 MBq). In the following discussion, we assume that 96gTc (95gTc) isotopes with a radioactivity of 49 MBq (190 MBq) are injected into a patient for a diagnostic scan. When 96gTc (95gTc) is produced by a single cyclotron with an irradiation time of 8 h and a reaction rate at 11 MeV, which has been calculated by PHITS, the single cyclotron can produce 96gTc (95gTc) with a radioactivity of 12 GBq (39 GBq) each day (see Table 5). These 96gTc (95gTc) isotopes can be used for 240 (200) diagnostic scans.

Table 5

Quantities of Tc radioisotopes produced by a proton accelerator. The proton energies are 16 MeV and 11 MeV for 99mTc and 95gTc (96gTc) production, respectively. The irradiation time for 95gTc (96gTc) production is 8 h.

Isotopes	99mTca,c	99mTcb,c	95gTcd	95mTcd	96gTcd	96mTcd
Number of atoms [1/μA]	3 × 1013	5.1 × 1013	4 × 1013	2.1 × 1013	5.7 × 1013	7.4 × 1012
Radioactivity [MBq/μA]	964	1646	390	2.8	110	1700

Note that the isomer ratio 99mTc/(99mTc + 99gTc) decreases to 0.26 and 0.22 for the irradiation time of 3 h and 6 h, respectively [35].

99mTc produced by the irradiation time of 3 h.

99mTc produced by the irradiation time of 6 h.

These values are taken from Ref. [35], in which the target thickness has been tuned to decrease the proton energy from 16 MeV to 10 MeV.

These values are calculated with a target thickness of 1 g/cm2 without the cooling time.

Medical cyclotrons that have been operated to generate radioisotopes for PET can also be employed to generate 95gTc/96gTc. A medical PET cyclotron is typically employed for 2 (up to 4) runs per day. The maximum time of each run is 2 h because the half-life of 19F is 2 h, which is the longest among the major radioisotopes for PET. Thus, it is possible to operate these cyclotrons for 8 h per day to generate 95gTc/96gTc isotopes. If the beam energy is higher than 12 MeV, the energy can be decreased using metal absorbers. In Japan, more than 200 cyclotrons for PET have already been introduced. 200 cyclotrons can supply 96gTc (95gTc) tracers for 48,000 (40,000) diagnostic scans per day. Note that approximately 70,000 treatments using 99mTc-labeled radiotracers are administered worldwide per day.

Chemical separation

After the proton beam irradiation, the generated 95gTc (96gTc) should be separated from the large excess of molybdenum. Similar technologies have been developed for 99mTc production using the 100Mo(p, 2n)99mTc [12, 17, 18, 20], 100Mo(n, 2n)99Mo reactions [11], and 94mTc production with the 94Mo(p, n)94mTc reaction [28, 29]: they include thermal separation [11], ion exchange [18, 20], aqueous biphasic extraction chromatography [12], and solid-phase extraction with cross-linked polyethylene glycol resins [17]. Because the chemical behavior of the Tc isotopes is the same, these technologies can be applied to the production of other Tc isotopes from Mo targets. Compact automatic modules for the separation and purification of 99mTc produced by the 100Mo(p, 2n) reaction have been developed [12, 17, 20], and the use of such automatic modules is feasible.

All the radiopharmaceuticals developed from 99mTc can, in principle, be produced using the other Tc isotopes if the radionuclide purity and other chemical characteristics are adequate. Various pharmaceutical kits have been developed to produce 99mTc-labeled tracers from sodium pertechnetate (Na99mTcO4) eluted with saline solution. Thus, if 95gTc and 96gTc is delivered to hospitals in the form of sodium pertechnetate with saline solution, various radiotracers can be generated in hospitals using these pharmaceutical kits. In this scheme, 95gTc (96gTc) is kept for 3–5 days (2 weeks) in hospitals, in which 95gTc (96gTc) continuously decays to 95Mo (96Mo). For example, when 96gTc is kept for 7 days, approximately 81% of 96gTc decays to 96Mo. Thus, a chemical procedure to remove Mo from Tc with saline solution should be developed.

Purity

In the energy region of E < 12 MeV, (p, n) reactions are dominant for Mo isotopes (see Fig. 4, Fig. 5); the production of other elements such as niobium is negligible. Thus, when an isotope-enriched Mo target is contaminated with other Mo isotopes, the Tc medical radioisotope is contaminated with other Tc radionuclides produced by (p, n) reactions on the contaminated Mo isotopes (see the solid arrows in Fig. 1). Because the half-lives of 92Tc and 100Tc are as short as 4.23 m and 15.8 s, respectively, they almost entirely decay away during the period from the end of the beam irradiation to a diagnostic scan in hospitals. In contrast, the γ-ray yields of the 97Tc and 98Tc isotopes are relatively small because their half-lives are as long as 4.2 × 106 y and 4.2 × 106 y, respectively. Thus, the radionuclide purity is determined by the factions of the contaminated isotopes of 94Mo, 95Mo, and 96Mo, which can produce 94g/94mTc, 95g/95mTc, and 96gTc, respectively. The 95Mo (96Mo) isotope with an enrichment of approximately 97% has been widely used to study nuclear physics [29], in which the fractions of other Mo isotopes are usually lower than 1%. As the enrichment increases, the fraction of the contamination decreases. It should be emphasized that the highly enriched 100Mo targets with a fraction of up to 99.8% have been commercially provided, in which the maximum fraction of contaminated Mo isotopes is 0.17% of 98Mo [16]. If highly enriched 95,96Mo targets are commercially produced, the fraction of contaminated Mo isotopes is expected to be much lower than 1%. Note that the co-production of 95mTc (T1/2 = 61 d) with 95gTc decreases the purity of 95gTc, because it is practically impossible to separate 95mTc and 95gTc. Thus, 96gTc is more suitable for the medical radioisotope than 95gTc.

γ-Ray imaging

The energies of most γ-rays emitted from Tc isotopes are in the range of 700–1200 keV (see Fig. 2, Fig. 3). In this energy region, Compton scattering is the dominant interaction between photons and atoms. It is thus difficult to image using a detector system coupled with collimators to limit the angle of incident γ-rays, such as the conventional γ-ray detector for the SPECT. Therefore, we need a detector that can reconstruct the Compton scattering event by event. Compton cameras are well-known as such detectors. To determine the direction of an incident γ-ray, in principle, both the angles of the scattered γ-ray and the scattered electron should be measured. However, standard Compton cameras can measure only the angle of the scattered γ-ray. Recently, the ETCC was developed [41, 42, 43, 44, 45, 46]. The ETCC consists of a gaseous time-projection chamber (TPC) coupled to a micro pattern gas detector to measure precisely the track of the recoil electron generated by Compton scattering by an incident photon and a position-sensitive scintillation camera to detect the scattered γ-ray. As a result, the ETCC has an excellent feature: it can measure the angles of both the scattered γ-ray and electron, thereby obtaining the direction of the incident γ-ray. A previous study [41] revealed that the ETCC could provide a clear definition of the point spread function (PSF) and also showed that the standard Compton camera has a PSF of several tens of degrees, whereas the ETCC could give a PSF that is smaller than or equal to ten degrees. Furthermore, Tanimori et al. [43] showed that only the ETCC can perform imaging spectroscopy and that the PSF of the Compton camera intrinsically made this capability difficult due to geometrical optics. Therefore, the ETCC is the most suitable for γ-imaging in the Compton scattering energy region. Hatsukawa et al. [44] demonstrated the chemical separation of the 95mTc produced by the 95Mo(p, n)95mTc reaction and the measurement of γ-rays from 95mTc using an ETCC.

Fig. 3

Partial level schemes for the ground state of 95Tc (left) and the 95Tc isomer (right). The solid arrows with numbers denote decay γ-rays and their energies in unit of keV. The dashed arrows show population from meta-stable states.

We can accurately estimate the imaging performance of the ETCC from the PSF and the detection efficiency of an incident γ-ray. As mentioned in a previous study [43], the detection efficiency and PSF can be precisely estimated using the Compton scattering cross section of the gas and the stopping power of the scintillation detector used in the ETCC due to the simple structure of a cubic gas chamber and an array of scintillator pixels placed at the bottom of the gas chamber. A 30-cm3 ETCC with a 1-atm Ar gas TPC and one Gd2Si2O7:Ce (GSO) scintillation detector array has a PSF of 15 degrees and a detection efficiency of approximately 0.1% for an incident 511-keV γ-ray [42], and a compact 10-cm3 ETCC with 1.5-atm Ar gas and two GSO scintillator arrays gives a detection efficiency of 0.02% at 511 keV [45].

In 2017, Tanimori et al. expects to release a new type of compact ETCC with a 20-cm diameter-cylindrical TPC with 2-atm Ar gas and two GSO detector arrays, with a detection efficiency of approximately 0.1% at 511 keV [46]. The sensitivity of a γ-ray imaging module is approximately proportional to the detection efficiency and the effective area around a human body. An advantage of the ETCC is that it provides a wide field of view of 3 sr. The construction of ETCC-based γ-ray imaging modules with an effective area of 50% is planned for medical diagnostic scanning. Even if an ETCC module with an effective area of only 25% will be available, its sensitivity for the detection of γ-rays from radiotracers injected into a human body is approximately 200 cps/MBq at 511 keV, where we assume a multiplicity of M = 1 and an average penetration probability of 80%. This sensitivity is nearly equal to that of the typical SPECT (190 cps/MBq) [39]. Note that as a final goal of the ETCC, a previous study [41] suggested the possibility that a 30-cm3 ETCC with 3-atm CF4 gas will give a detection efficiency of a few % at 511 keV with a good PSF of <3 degrees. This high resolution PSP suggests the sensitivity of a γ-imaging module based on ETCC will be higher than that of the typical SPECT detector by an order of magnitude and the ETCC can provide clearer image. Finally, we would like to stress that the ETCC does not require collimators like SPECT. Thus, it is possible to develop lighter and smaller imaging modules than SPECT.

Feasibility of 96gTc and 95gTc production

The (p, 2n) reaction method [12, 13, 14, 15, 16, 17, 18, 19, 20, 21] and the present proposal have common features that both methods use medical cyclotrons and require the purification of Tc isotopes from Mo targets. Various basic technologies for this (p, 2n) reaction method were developed. As a result, automatic purification modules to separate 99mTc from a large excess of Mo after the proton beam irradiation were developed [12, 17, 20]. Furthermore, a new process for solid Mo targets based on the electrophoretic deposition of Mo powder onto Ta metal plates was developed [15]. A point for consideration in the economical production of 99mTc is the recycling of the enriched 100Mo targets after the chemical separation of 99mTc. However, high recovery yields of > 90% [12] and 95% [21] have already been achieved. At present, approximately 250 medical cyclotrons with an operating energy of 16.5 MeV, which can produce 99mTc by the (p, 2n) reaction, have been installed worldwide [19]. Therefore, the (p, 2n) reaction method is feasible for production without the use of either uranium targets or nuclear reactors. Because the chemistry of the separation of 99mTc from 100Mo and 96gTc (95gTc) from 96Mo (95Mo) is, in principle, the same, most technologies developed for the (p, 2n) reaction method can be applied to the 96gTc (95gTc) production. However, the 96gTc (95gTc) method requires following three technologies. First, we need a new chemical procedure to remove Mo, which is produced via β-decay of Tc, from Tc with saline solution. Second, it is expected that highly enriched 95,96Mo targets with fraction of approximately 99.8% are provided. Third, the 96gTc (95gTc) method requires new γ-imaging modules based on the ETCC, but it has two advantages: compactness without collimators and clearer image.

Conclusion

95gTc and 96gTc are studied as alternatives to the medical radioisotope 99mTc. The calculated γ-ray intensities of a 96gTc (95gTc) nucleus injected into a human body is 63% (70%) relative to that of a 99mTc nucleus. This result suggests that they can work as γ-ray emitters. 96gTc (95gTc) can be produced by (p, n) reactions on enriched 96Mo (95Mo) targets with proton beams provided by compact accelerators such as medical PET cyclotrons. A cyclotron with proton beams with energies of 11–12 MeV and a current of 100 μA can produce 12 GBq (39 GBq) of 96gTc (95gTc) for operation of 8 h, which can be used for 240 (200) diagnostic scans. Because the co-production of 95mTc (T1/2 = 61 d) with 95gTc decreases the purity of 95gTc, 96gTc is more suitable for medical radioisotopes than 95gTc. The estimated patient radiation doses of 95gTc and 96gTc are larger than that of 99mTc by a factor of 2–3 based on the applied assumptions. This method requires the development of a new γ-imaging module based on ETCC, which can provide clearer image than typical SPECT detectors.

Declarations

Author contribution statement

Takehito Hayakawa: Conceived and designed the experiments; Analyzed and interpreted the data; Wrote the paper.

Yuichi Hatsukawa: Analyzed and interpreted the data; Wrote the paper.

Toru Tanimori: Contributed reagents, materials, analysis tools or data; Wrote the paper.

Funding statement

This work was supported by JSPS KAKENHI Grant Numbers JP16K05025 and JP15H03665.

Competing interest statement

The authors declare no conflict of interest.

Additional information

No additional information is available for this paper.